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Spinal Mechanics during Drop Landing: Effects of Gender, Fatigue and Landing Technique

HIDE
 Title Page
 Dedication
 Acknowledgement
 Table of Contents
 List of Tables
 List of Figures
 Abstract
 Introduction
 Materials and methods
 Literature review
 Results
 Discussion
 Conclusion
 Future work
 Appendices
 References
 Biographical sketch
 

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1 SPINAL MECHANICS DURING DROP LANDING: EFFECTS OF GENDER, FA TIGUE AND LANDING TECHNIQUE By SOO-AN PARK A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL OF THE UNIVERSITY OF FLOR IDA IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF DOCTOR OF PHILOSOPHY UNIVERSITY OF FLORIDA 2006

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2 Copyright 2006 by Soo-An Park

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3 To my parents

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4 ACKNOWLEDGMENTS Many people have earned my gratitude for their guidance and support during my doctoral education and the completion of my dissertation. First I would lik e to thank my parents, SeungJae Park and Jung-Nim Kang. They have supported me in every path with love I have taken and they have provided me with the work ethic, value, and encouragement necessary to achieve my goals. I would also like to express my grat itude to my wife, H yunhee Kwon and my son, Joonsuh, for their love and patience they have provided over the past several years. The completion of my graduate education has been a joint endeavor and a shared achievement. I would like to expres s thanks to my committee members who have challenged me to become a better scientist and person. Dr. John C how has exponentially strengthened my research and dissertation and my abilities as a biomechan ist. Dr. Mark Tillman has provided great support and assistance in my research. Dr. Ronald Si ders has provided support in my teaching during graduation education, and Dr. False tti has assisted me in setting up the research hypotheses in my dissertation.

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5 TABLE OF CONTENTS page ACKNOWLEDGMENTS...............................................................................................................4 LIST OF TABLES................................................................................................................. ..........7 LIST OF FIGURES................................................................................................................ .........9 ABSTRACT....................................................................................................................... ............10 CHAPTER 1 INTRODUCTION..................................................................................................................12 Purpose of the Study........................................................................................................... ....16 Significance of the Study...................................................................................................... ..16 Hypotheses..................................................................................................................... .........17 Limitations.................................................................................................................... ..........18 2 MATERIALS AND METHODS...........................................................................................20 Subjects....................................................................................................................... ............20 Sample Size Justification...................................................................................................... ..20 Experimental Setup............................................................................................................. ....21 Testing Protocol............................................................................................................... .......21 Pre-Fatigue Landing Trials..............................................................................................22 Fatigue Procedure............................................................................................................22 Post-Fatigue Landing Trials............................................................................................23 Data Reduction................................................................................................................. ......23 Data Analysis.................................................................................................................. ........25 3 LITERATURE REVIEW.......................................................................................................34 Biomechanical Properties of Spinal Structures......................................................................35 Intervertebral Disc...........................................................................................................35 Vertebra....................................................................................................................... ....39 Spinal Ligaments.............................................................................................................40 Biomechanical Properties of Spinal Segments.......................................................................41 Multisegmental Mechan ics of the Spine.........................................................................41 Regional Mechanics of the Spine....................................................................................43 Biomechanical Performance of Spine In Vivo.......................................................................47 Trunk Posturing...............................................................................................................47 Weight Lifting.................................................................................................................48 Sitting and Standing........................................................................................................50 Walking........................................................................................................................ ...51 Running........................................................................................................................ ...53

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6 Biomechanical Etiology of Spinal Pain..................................................................................56 Vibration...................................................................................................................... ....57 Lordosis....................................................................................................................... ....59 Torsion........................................................................................................................ .....60 Biomechanical Performance of Painful Spine........................................................................61 Landing Biomechanics...........................................................................................................63 Biomechanical Performance of Lower Extremity Joints during Landing.......................64 Gender Difference...........................................................................................................66 Landing Stiffness.............................................................................................................67 Performance of Adapted Landing Biomechanics to Various Conditions.......................70 4 RESULTS........................................................................................................................ .......77 Effects of Landing Technique................................................................................................77 Effects of Knee Join t Muscles Fatigue...................................................................................79 5 DISCUSSION..................................................................................................................... ....92 Effects of Landing Technique................................................................................................92 Effects of Knee Join t Muscles Fatigue.................................................................................100 6 CONCLUSION.....................................................................................................................110 Effects of Landing Technique..............................................................................................110 Effects of Knee Join t Muscles Fatigue.................................................................................110 7 FUTURE WORK..................................................................................................................112 APPENDIX A INFORMED CONSENT......................................................................................................113 B PRELIMINARY STUDY.....................................................................................................118 C MANOVA and ANOVA TABLES......................................................................................119 Effects of Landing Technique..............................................................................................119 Effects of Knee Join t Muscles Fatigue.................................................................................125 LIST OF REFERENCES.............................................................................................................130 BIOGRAPHICAL SKETCH.......................................................................................................146

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7 LIST OF TABLES Table page 2-1 Geisser-Greenhouse Correction Detail Report..................................................................27 2-2 MANOVA table for the fatigue indices.............................................................................27 2-3 Collapsed mean and SD values of fatigue indices before/after th e fatigue procedure......27 3-1 Average neutral zones for a functional spin al units in different regions of the spine ( ).............................................................................................................................. .........73 3-2 Representative ranges of motion of C0-C1-C2 complex ( )..............................................73 3-3 Representative ranges and limits of moti on of the middle and lower cervical spines ( ).............................................................................................................................. .........73 3-4 Normal active cervical ranges of motion ( in vivo ) reported in the literatures ( )..............73 3-5 Representative ranges and limits of motion of the thoracic spine ( )................................74 3-6 Representative ranges and limits of motion of the lumbar spine ( ).................................74 3-7 Comparison of lumbar compression loads in various trunk postures without external loading........................................................................................................................ ........74 3-8 Average ranges of motion of the lumbar spine in normal walking and running in different studies ( )............................................................................................................75 3-9 Peak compression loads to the lo wer lumbar level during walking ( BW).....................75 3-10 Intradiscal pressure of low lumb ar level during various activities....................................75 4-1 Collapsed mean and SD values of differe nt landing variables for different genders and landing techniques.......................................................................................................82 4-2 Collapsed mean and SD values of to uchdown angle and extension ROM of each spinal region for different genders and landing techniques...............................................83 4-3 Collapsed mean and SD values of kine tic variables at L/S and C/T junctions for different genders and landing techniques..........................................................................84 4-4 Collapsed mean and SD values of differe nt landing variables for different genders and fatigue levels............................................................................................................. ..85 4-5 Collapsed mean and SD values of to uchdown angle and extension ROM of each spinal region for different genders and fatigue levels........................................................86

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8 4-6 Collapsed mean and SD values of kine tic variables at L/S and C/T junctions for different genders a nd fatigue levels...................................................................................87 B-1 Kinetic characteristics of L/S junction data for sa mple size justification........................118 C-1 MANOVA table for landing variables, touchdown angle variab les, extension ROM variables, and L/S and C/T kinetic variables in the study of landing technique effects..119 C-2 Univariate tests for the different landing variables in the study of landing technique effects........................................................................................................................ .......120 C-3 Univariate tests for the touchdown angle va riables in the study of landing technique effects........................................................................................................................ .......121 C-4 Univariate tests for the extension ROM variables in the study of landing technique effects........................................................................................................................ .......122 C-5 Univariate tests for the L/S kinetic variables in the study of landing technique effects........................................................................................................................ .......123 C-6 Univariate tests for the C/T kinetic variables in the study of landing technique effects........................................................................................................................ .......124 C-7 MANOVA table for the fatigue indices in the study of knee joint muscles fatigue........125 C-8 Univariate tests of within-subjects effect s for the fatigue indices in the study of knee joint muscles fatigue........................................................................................................125 C-9 MANOVA table for landing variables, touchdown angle variab les, extension ROM variables, and L/S and C/T kinetic vari ables in the study of knee joint muscles fatigue........................................................................................................................ ......126 C-10 Univariate tests of between-subjects eff ects for the landing variables in the study of knee joint muscles fatigue................................................................................................127 C-11 Univariate tests of between-subjects effect s for the touchdown a ngle variables in the study of knee joint muscles fatigue..................................................................................127 C-12 Univariate tests of between-subjects eff ects for the extension ROM variables in the study of knee joint muscles fatigue..................................................................................128 C-13 Univariate tests of within-subjects effect s for the kinetic variab les of L/S junction in the study of knee joint muscles fatigue............................................................................128 C-14 Univariate tests of between-subjects a nd within-subjects effects for the kinetic variables of C/T junction.................................................................................................129

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9 LIST OF FIGURES Figure page 2-1 Correlations of power and sample si ze for each combination of variables.......................28 2-2 Experimental setup......................................................................................................... ....29 2-3 A subject with markers on.................................................................................................30 2-4 Overview of the experimental procedures.........................................................................31 2-5 Marker placement (left) and definition of regional a ngles of the spine (right).................32 2-6 Kinematic variables defined by the critic al instants identified from kinematic and forceplate data................................................................................................................ ....33 3-1 The load-displacement curve of a functiona l spinal unit (FSU) is generally nonlinear and biphasic [neutral zone (N Z) and elastic zone (EZ)]....................................................76 4-1 Significant interactions of the touchdow n angle and C/T kinetic variables between gender and landing technique............................................................................................88 4-2 Representative kinematic s of each spinal region...............................................................89 4-3 Representative kinematic s of each spinal region...............................................................90 4-4 Significant interactions of the kinetic variables at L/S and C/T junctions between gender and fatigue level.....................................................................................................91

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10 Abstract of Dissertation Pres ented to the Graduate School of the University of Florida in Partial Fulfillment of the Requirements for the Degree of Doctor of Philosophy SPINAL MECHANICS DURING DROP LANDING: EFFECTS OF GENDER, FA TIGUE AND LANDING TECHNIQUE By Soo-An Park December 2006 Chair: John W. Chow Major: Health and Human Performance Objective: To investigate the kinematics of th e spinal column and the kinetics of lumbosacral (L/S) and cervicothoracic (C/T) junc tions during the drop landing, and to evaluate the effects of gender, landing technique, and fatigue. Methods: Thirteen male and 13 female healthy y oung volunteers were tested. To track the kinematics of different spinal regions, surface ma rkers were placed on skin over selected spinous processes. Data were collected using a 3-D motion capture system and a forceplate. The subject performed 3 drop landings using his/her own land ing technique (NL) and 3 soft landings with instruction (SL; SL1). During each trial, the subject descended fr om a 50-cm height platform and landed on a forceplate with the le ft foot at the center of forc eplate. After completing isokinetic knee flexion/extension exercises and a 30-minute run on a motorized treadmill, the subject performed 3 more soft landings in a knee joint muscle fatigue st ate (SL2). Kinematic variables included touchdown angle and initial extension ra nge of motion of different spinal regions. Kinetic variables included joint resultants at L/ S and C/T junctions comp uted using an inverse dynamics approach. Multivariate analyses of variance (MANOVA) and follow-up univariate analyses of variance (ANOVA) were used to exam ine the effects of gende r, landing technique, and fatigue on differe nt kinematic and kinetic variables.

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11 Results: Females exhibited a significantly exte nded thoracolumbar landing posture and greater thoracic motion than males in all 3 expe rimental conditions. Thoracic and lower cervical extension motions increased and most joint result ants decreased significantly when going from NL to SL. Posterior shear force in males and anterior shear force in females were significantly greater than their counterpart s at C/T junction during NL. Fe males exhibited significantly increased joint resultants from SL1 to SL2, while males did not. Conclusion: The spinal column is more activel y involved in energy absorption during drop landings in males, and the thoracolumba r region could be more loaded by hyperextension during soft landing in females comp aring to males. Repeated drop la ndings may cause injuries to the cervical spine by different mechanisms in each gender. For females, soft landings under fatigue condition can be a risk factor of spinal injury.

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12 CHAPTER 1 INTRODUCTION Spinal function in vivo has been generall y determined by radiographs, showing the kinematics of each spinal segment and spinal region (Wong et al., 2006) or by inclinometer measurements (Ng et al., 2001). Most previous investigations about spinal kinematics were based on postures of the spinal column in static conditions and static spinal kinematics. Active spinal motion can be identified as the end range of motion (ROM) achieve d by the subjects, and the passive motion is the end ROM obtained by applying external forces to a fully motioned spine. Both active and passive ROM of the spine ha ve been used to interp ret the functional status of the spine in clinical and labor atory studies (Dvor ak et al., 1988). Several researches have evalua ted the kinematics of the spin e in dynamic situations like walking (Callaghan et al., 1999; Crosbie et al ., 1997) and running (Schache et al., 2002). Although results from both walking and running anal yses may have clinical implications, these two locomotive tasks may not be the most adequate tasks to reveal spin al function in dynamic situation. The age of patients who deve lop spinal degeneration and undergo spinal surgeries is getting younger and this population is getting larger (Kjaer et al., 2005). The prevalence of spinal pain and degeneration is higher in active in dividuals (Bono, 2004). Th e biomechanics of the spine in a static condition are quite different from those in a dynamic condition. Demands for developing a functional ev aluation of the spine in vivo have increased, because conventional radiographic study of spinal ROM is not good en ough to estimate the spinal function of every subject. Also with the development of the su rgical technique of sp inal arthroplasty, the restoration of the original function of each patient has become a primary goal of the surgery.

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13 Therefore, normal spinal mechanics in vivo in various tasks and the mechanical factors contributing to spinal pain and degeneration are useful information to medical practitioners. It has been suspected that most disabling and chronic back pain arises from intervertebral disc degeneration. With disc de generation, biologic as well as biomechanical changes follow. Disc degeneration occurs most commonly in the third to fifth decades of life. Aging causes definite changes in the morphology and composition of spinal tissues that are mostly unrelated to pain. However, many studies reported that aging weakens the intervertebral disc tissues due to decreased cell number, high apoptosis rate, and th e different response to biologic environments. Impaired function of the intervertebral discs may make people more vulnerable to mechanical injuries, which can initiate further struct ural and symptomatic disc degeneration. Nonphysiologic loading to spinal structures can contribute to in tervertebral disc degeneration. With the increased loading to the trunk, spinal shrinkage was found double to the unloaded condition (Fowler et al ., 1994). A flexion posture signi ficantly increases extensor muscle activity when compared with a standi ng neutral posture (Arjmand & Shirazi-Adl, 2006). Forward leaning of the trunk causes the vertebrae in anterior translati on, and disc loads and stresses were significantly incr eased most markedly at the L5/S 1 level (Harrison et al., 2005). When the torso is fully flexed during repetitive lif ting tasks, fatigue failure of spinal tissues can occur rapidly (Wrigley et al., 2005). On the other hand, abnormally low loading causes atrophy in muscle, cartilage and bone, l eaving them less able to resist high loads (White & Panjabi, 1990). Abnormalities in the lower extremities can affect spinal mechanics. For example, patients who have leg length discrepancy due to lower ex tremity disorders demonstrate different spinal kinematics during gait when compared with normal subjects suggesti ng greater risk of

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14 developing spinal disorders (Kakushima et al ., 2003). However, a walking task may not be sensitive enough to detect the e ffects of altered mechanics in the lower extremity upon spinal mechanics. Gender differences in spinal kinematics duri ng trunk posturing hinted that there may be different mechanical causes of spinal degene ration in different genders. During a prolonged sitting, males exhibit more flexion of lumbar spine and trunk than females. Males and females may be exposed to different loading patterns during certain prolonged pos tures and can develop different injuries or degeneration mechan isms of the spine (Dunk & Callaghan, 2005). Various in vivo and in vitro biomechanical techniques have been developed to investigate spinal mechanics, but they all have limitations. Physical properties of the human spine may be obtained from studies of living subjects, whole cadavers, isolated whole cadaveric spines, and isolated spinal segments. A living subject provides realistic but less accurate measurements. An isolated spinal segment lacks muscles, but can pr ovide accurate data and allows the possibility of studying the effects due to trauma and surgical stabilizations (Wh ite & Panjabi, 1990). The force applied to the spine depends on body weight (Rodacki et al., 2005), external loads (Lawrence et al., 2005), and internal musc le forces (Arjmand & Shirazi-Adl, 2006) which can be varied during dynamic activities (Chow et al., 2003; Tully et al., 2005). Direct measurement of spinal compression force pione ered by Nachemson (1966; 1964), obtained by inserting a pressure needle into the lumbar intervertebral discs of living subjects, has been benchmarked and compared to many other studies that measured forces applied to the spine in vivo (Ledet et al., 2005; Sato et al., 1999). However, most researchers who have investigated spinal mechanics utilized indir ect techniques on living subjects and cadaveric spinal segments because of the invasiveness and limited localization of direct measurements.

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15 Normal spinal mechanics during activities of living subjects and each spinal segment as functional spinal unit (FSU) and spinal column ar e usually studied using indirect measurements. Spinal mechanics during various movements walking, forward/backward bending, sit-tostand/stand-to-sit, etc. have been investigat ed using image-based motion analysis systems. Using surface electromyography (EMG) techniques, spinal muscle activity during different spinal movements has been investigated. McGi ll (1992) estimated the moments generated by trunk muscles using an EMG-driven musculoske letal model during trunk posturing movements. By applying inverse dynamic techniques to a rigi d segment model, joint resultants at different lower extremity joints and spinal motion segments can also be calculated. However, many spinal vertebrae included in one trunk segment may mimic the join t resultant values at the trunk segment, and the kinematics of the spinal region in vivo could not be accesse d in details (Khoo et al., 1995). Using forceplate and kinematic data, invers e dynamic techniques are commonly used to calculate joint resultants at lower extremity jo ints during various activities (Kernozek et al., 2005). Jumping and landing were commonly adopted for measuring joint resultants at lower extremity joints simulating active and vigorous movements. A recent study completed by the author indicated that limiti ng trunk movements caused changes in the mechanics of lower extremity joints during drop la ndings (Park et al., 2006b). Th is finding suggests that body segments proximal to the hip joints could be in volved in regulating the force transmitted to the lower extremity joints during the landing phase. Conversely, different mech anical configurations of lower extremity joints may affect spinal mechanics during landing and the findings from spinal mechanics for each specific configuration may provide insights into the spinal function in dynamic situations.

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16 Purpose of the Study The purpose of this study was to investigate the kinematics of different regions of the spinal column and the kinetics at the lumbos acral (L/S) and cervicot horacic (C/T) junctions during drop landings. Specificall y, this study evaluate d the effects of gender and landing techniques (self-selected landing and instructed soft landing techniques) on spinal mechanics during drop landing in the first part of this study. In the second pa rt of this study, the effects of gender and muscular fatigue of knee flexor/extensor on spinal mechanics during drop landings were examined. Significance of the Study Previous researchers have evaluated the mechanical causes of spinal degeneration in vivo and in vitro while simulating various activities using di rect and indirect measurement techniques. Most activities employed in these biomechanical studies were simple, everyday activities that do not demand much spinal movement. However, th e populations who have spinal degeneration and undergo spinal surgeries are getti ng larger and younger. These indi viduals want to be physically active and participate in activities that may de mand vigorous spinal move ments. Drop landings have been widely used to examine coordination and mechanical stress at different joints under dynamic situations. However, most landing studies were confined to the biomechanics of lower extremity joints and very few studies eval uated mechanics of upper body movements during landings. In addition to lower extremity joints, this study attempted to identify mechanical characteristics of different spinal regions duri ng drop landings. The results might provide insight into the effects of gender, la nding technique and fatigue of knee joint muscles on spinal mechanics during drop landings.

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17 Hypotheses In the absence of extensive pilot data and re levant data reported in the literature, the following questions were raised and null hypotheses were tested to address the aims of this study: To determine the effects of gender and landi ng technique on spinal column kinematics and loads to the lumbosacral (L/S) and cervicot horacic (C/T) junctions during drop landings. Q1: Would there be significant differences in sp inal column kinematics and loads to the L/S and C/T junctions during drop landings betw een males and females regardless of landing technique? 1a. There would be no significant differences in motion characteristics of the spinal column between males and females. 1b. There would be no significant differences in peak resultant forces and moments transmitted to the L/S and C/T junctions between males and females. Q2: Would the self-selected and soft landing techniques cause signifi cant differences in spinal column kinematics and loads to the L/S and C/T junctions during drop landings? 2a. The conditions for landing technique woul d not cause significant differences in motion characteristics of the spinal column. 2b. The conditions for landing technique would not cause significant differences in peak resultant forces and moments transmitted to the L/S and C/T junctions. To determine the effects of gender and fati gue of knee joint muscles on spinal column kinematics and loads to the L/S and C/T junctio ns during drop landings using a soft landing technique.

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18 Q3: Would there be significant differences in sp inal column kinematics and loads to the L/S and C/T junctions during drop landings us ing a soft landing technique between males and females regardless of fatigue condition? 3a. There would be no significant differences in motion characteristics of the spinal column between males and females. 3b. There would be no significant differences in peak resultant forces and moments transmitted to the L/S and C/T junctions between males and females. Q4: Would fatigue of knee joint muscles caus e significant differences in spinal column kinematics and loads to the L/S and C/T j unctions during drop landings using a soft landing technique? 4a. Fatigue of knee joint muscles would not ca use differences in motion characteristics of the spinal column. 4b. Fatigue of knee joint muscles would not cause differences in peak resultant forces and moments transmitted to the L/S and C/T junctions. Limitations Measurement errors of forceplate and digital video cameras are always present but they are considered acceptable within the sp ecifications of the manufacturers. Marker placement was controlled cautiously to minimize errors. Sagittal spinal kinematics relative to the ad jacent spinal region based on spinal marker locations would have some errors due to skin movement, but will be considered acceptable.

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19 The center of L/S junction was assumed to be located at the midpoint between both posterior iliac crest markers at L5/S1 level. The center of the C/T junction was assumed to be located at the midpoint between the two acromial process markers. Mechanical characteristics of the lower extr emities were assumed to be symmetrical and only the data collected from the left leg were used to calculate joint resultants at the L/S and C/T junctions.

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20 CHAPTER 2 MATERIALS AND METHODS Subjects Thirteen male (age: 21.4 1.3 yrs, mass: 74.2 10.2 kg, height: 174.8 5.2 cm) and 13 female (age: 21.1 1.3 yrs, mass: 58.6 6.8 kg, height: 165.5 5.3 cm) healthy and physically active individuals participated in this study. They were free from any cardi o-respiratory diseases that would prevent them from completing the fati gue procedures and musculoskeletal diseases or injuries that could influence spinal and lower ex tremity mechanics. Before testing, each subject carefully read and signed a written informed c onsent approved by the Institutional Review Board of the University of Florida (Appendix A). Sample Size Justification To simplify the calculation of sample size de termination, data of a dependent variable (peak extensor moment at the lu mbosacral junction) from a pilot study were used to set up a 2 2 [gender landing technique: normal landing and soft landing] ANOVA with repeated measures on the last factor. Based on the pre liminary data collected from one male and one female, the means of male and female data were 2.74 and 2.58 N m kg-1 BH-1, respectively, and those for NL and SL were 3.05 and 2.27 N m kg-1 BH-1, respectively (Note. BH: body height). There was a 39% difference between male and female, and a 26% difference between normal and soft landing conditions. The mean standard deviation value was 0.3 (Table B-1 in Appendix B). Based on these values, the current study was de signed to detect at least 12% changes in peak extensor moment in between-subject group and 10% difference in within-subjects groups with alpha = 0.05 and beta = 0.2 (80% power). A Geisser-Greenhouse corre ction report indicated that power values were over 80% for all the te rms with a sample size of n=13 (Table 2-1, Figure

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21 2-1). Therefore, at least 12 subj ects were required for each with in-subject group to determine the effects of the treatment. All the calculat ions were performed using PASS 2005 (Number Cruncher Statistical Systems, Kaysville, Utah). Experimental Setup A forceplate (Type 4060-10, Bertec Corporat ion, Columbus, OH) operating at 1,200 Hz was set up at the center of the Biomechanics Res earch Laboratory (Figure 2-2). A 50-cm height platform was placed behind and slightly to the right of the forceplate. Seven Hawk digital cameras (Motion Analysis Corp., Santa Rosa, CA) we re stationed around the forceplate to collect kinematic data and were 3-4 m from the land ing area. Kinematic data were captured at a frequency of 100 Hz. The system was calibrated pr ior to each testing session according to the procedures specified by the manufacturer. Testing Protocol To expose the lower extremity and the back of the trunk, subjects were asked to wear only short pants (both males and females) and sports bra (for females only) (F igure 2-3). They wore their own sports shoes during testing. Upon co mpleting the measurements of body weight and height, each subject jogged on a treadmill and stretched with self-selected exercises for 10 minutes as a warm-up (Figure 2-4). Reflective markers (1.0 cm in diameter) were placed on the left second metatarsal head, dorsal navicular surface, heel, lateral malle olus, lower-shank, mid-shank, lateral tibial epicondyle, lower-thigh, mid-thigh, greater troc hanter, anterior superior iliac spines, 2nd spinous process on median sacral crest of sacrum to trac k the locations of the left lower extremity and pelvis (Kadaba et al., 1990; Kada ba et al., 1989). Another set of reflective markers was applied over the subjects spinous processes for measuri ng the kinematics of the spinal column (C4, C6, T1, T3, T6, T10, T12, L2, L4, both posterior ili ac crests at L5/S1 level) (Figure 2-5).

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22 After marker placement, drop la nding practice trials were prov ided to ensure consistent landing during experimental trials. In each tria l, the subject descended from a 50 cm-height platform and landed on a forceplate with his/her left foot at the cen ter of the forceplate and right foot on a wooden platform of the same height. Pre-Fatigue Landing Trials After practice drop landings, th e subject completed 3 trials of drop landing using his/her own landing technique (normal la nding: NL). The subject was then instructed on how to perform a drop landing using the soft landing technique (s oft landing: SL). He/She was instructed to try soft landing by touching the balls of feet on the ground at initial impact, dela ying heel contact, and using more knee flexion after the landing. Three trials of soft landing were collected (SL1). The averages over the three trials for each landin g technique were used in subsequent analyses. In each trial, kinematic and GRF data were samp led for 4 s. If the subject did not maintain balance after landing, that trial was discarded and repeated. Fatigue Procedure The subject was asked to s it on the chair of a KinCom dynamometer (Chattanooga Group, Inc., Hixson, TN) and perform 30 repetitions of is okinetic reciprocal knee flexion/extension with full ROM and maximal effort at 60 /s to induce muscular fatigue of knee joint flexor/extensor. The exercise was repeated at the speed of 180 /s to measure the isokinetic strength for the purpose of quantifying the fatigue level. After the isokinetic exercises, the subject ran on a motorized treadmill at 4-6 mph fo r 30 minutes. The running intensity was lower than the typical daily exercise for developing and maintaining fitn ess (Fletcher et al., 200 1; Pollock et al., 1998). Also, the speed and duration of running used in this study was known not to elicit cardiorespiratory fatigue (Hardin et al., 2004). If the subject became exhausted before the end of the 30

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23 minutes run, the running speed was reduced to 4.0 mph so that he/she could complete the run. Isokinetic knee flexion/extension exercises were repeated immedi ately after running to quantify the fatigue state of knee joint muscles. Post-Fatigue Landing Trials Immediately after the fatigue procedure, the subject performed 3 trials of drop landing using the soft landing technique from the same platform (soft landing performed under fatigue condition: SL2). The averages over the 3 tria ls were used in subsequent analyses. Data Reduction Kinematic data were processed using EVaRT 4.6 software (Motion Analysis Corp., Santa Rosa, CA). The animation of reflective markers in each trial was examined qualitatively by the investigator. Positional data were smoothed usin g a Butterworth low pass filter with a cutoff frequency of 10 Hz. Three-dimensional kinematic a nd kinetic data for the left ankle, knee, and hip joints, pelvis and trunk, and the kinematic da ta of spinal column markers were calculated using Kintrak 6.2 software (Mo tion Analysis Corp., Santa Rosa, CA). Locations of spinal column markers were used to define 5 spinal regions: lower cervical (LC), thoracic (TH), thoracolumbar (TL), lumbar (L) and sacral (S) regions (Figure 2-5). Joint resultants at the left knee and hip joints and L/S and C/T junctions were computed based on the kinematic and forceplate data us ing an inverse dynami cs approach. Assuming symmetry in lower extremity mechanics, mechanical characteristics of the right leg were the same at the left leg for the purpose of computing joint resultants at the L/S and C/T junctions. To minimize the variation due to individual di fferences in physique, force variables were normalized to the subjects body mass, and moment variables were normalized to body mass and body height.

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24 Muscle fatigue was quantified by the fatigue index. Fatigue index was determined by the decline in peak torque in 30 repetitions, and calculated by the following formula to yield a percent decrease for each is okinetic torque value: Fatigue index = 100 [(last 5 repetiti ons/highest consecuti ve 5 repetitions) 100] For each subject, the highest consecutive five repetitions were determined by the values attained from the two repeti tions immediately prior to, a nd following, the single highest repetition value. If the single hi ghest repetition value was observed within the first 3 repetitions, the first 5 repetitions were us ed to calculate the fatigue i ndex (Pincivero et al., 2003). Fatigue levels of knee joint muscles before/aft er the fatigue procedure were evaluated with the fatigue indices of knee flexor and extensor muscles. Repeated measures MANOVA revealed a significant main effect of knee joint muscles fatigue for the fatigue index variables (p=0.001), but did not reveal any significan t main effect of gender (p=0 .528) and interaction between gender and fatigue level (p=0.589) (T able 2-2). The univariate contra st procedures indicated that the fatigue indices of both knee flexors and ex tensors increased significantly by the fatigue procedure (Table 2-3). Only those subjects who demonstrated increased fatigue indices in both knee flexors and extensors (12 males and 10 females) were included in subs equent analyses (SL1 vls. SL2). The landing phase was defined using the critical instants identified fr om the kinematic and GRF data (Figure 2-6), and the cr itical instants are as follows: The instant when the vertical ground reaction fo rce (VGRF) starts to increase, the initial touchdown (the beginning of landing phase ), was identified from GRF data. Maximal knee joint flexion afte r initial touchdown was identified from the kinematic data (the end of landing phase). At the completion of data reduction, the depende nt variables were divi ded into five groups:

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25 Landing variables: peak vertical GRF (PVGRF), time for landing phase (t(LP)), knee flexion angle at touchdown ( TD(KFA)), ROM of knee flexion from touchdown to initial peak of knee flexion ( P(KFA)), hip flexion angle at touchdown ( TD(HFA)), ROM of hip flexion from touchdown to initial peak of hip flexion ( P(HFA)). Touchdown angles: lumbar re gional angle at touchdown ( TD(L/S)), thoracolumbar regional angle at touchdown ( TD(TL/L)), thoracic regional angle at touchdown ( TD(TH/TL)), lower cervical regional angle at touchdown ( TD(LC/TH)). Extension ROMs: extension ROM of lumbar region from touchdown to initial peak during landing phase ( P(L/S)), extension ROM of thoracolumbar region from touchdown to initial peak during landing phase ( P(TL/L)), extension ROM of thoracic region from touchdown to initial peak during landing phase ( P(TH/TL)), extension ROM of lower cervical region from touchdown to initial peak during landing phase ( P(LC/TH)). Kinetic variables at L/S junction: peak axial compressive force [AxF(L/S)], peak anterior shear force [ShF(L/S)ant], peak posterior shear force [ShF(L/S)post], peak flexor moment [FlxM(L/S)], peak extensor mome nt [ExtM(L/S)] after touchdown. Kinetic variables at C/T junction: peak axial compressive force [AxF(C/T)], peak anterior shear force [ShF(C/T)ant], peak posterior shear force [ShF(C/T)post], peak flexor moment [FlxM(C/T)], peak extensor mome nt [ExtM(C/T)] after touchdown. KFA was defined as the angle between the line of shank axis and thigh axis. HFA was defined as the angle between th e line of thigh axis and pelvis axis. The negative angles of TD(KFA) and TD(HFA) mean the flexions of knee and hip joints at th e touchdown. For P(KFA) and P(HFA), absolute values were used. A regional angle of the spine was defined as the angle between th e lines representing a spinal region and its lower adjacent region. A positive touchdown angle indicates the spinal region is in an extended state or extension mo tion relative to the lower adjacent region and a negative angle indicates the spinal region is in a flexed state or flexion motion (Figure 2-5). Data Analysis For the non-fatigued data (NL and SL data), the 5 groups of dependent variables were submitted to five separate 2 2 (Gender Landing type) MANOVA with repeated measures on the last factor. For the soft landing data (SL1 a nd SL2 data), the 5 groups of dependent variables

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26 were submitted to 5 separate 2 2 (Gender Fa tigue level) MANOVA with repeated measures on the last factor. Follow-up univa riate analyses were conducted when appropriate. Bonferroni adjustments were used during follow-up testing. A priori alpha level was set at 0.05 for all statistical procedures. All sta tistical tests were performed using SPSS 13.0 for Windows (SPSS Inc., Chicago, IL).

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27 Table 2-1. Geisser-Greenhous e Correction Detail Report. Term (levels) Power Alpha F Lambda df1|df2 Epsilon E (Epsilon) G1 n = 13 N = 26 Means 1 Gender (B: 2) 0.8310 0.05 4.26 9.25 1|24 1 1 0 Landing (W: 2) 1 0.05 4.26 43.33 1|24 1 1 0 BW 1 0.05 4.26 43.33 1|24 1 1 0 Table 2-2. MANOVA table fo r the fatigue indices. Effect Roy's Largest Root F Hypothesis df Error df Sig.(p) Observed Power (a) Gender 0.057 0.657 2 23 0.528 0.147 Fatigue* 0.875 10.067 2 23 0.001 0.971 Fatigue Gender 0.047 0.542 2 23 0.589 0.129 Significant main effect or interaction (p<0.05) Table 2-3. Collapsed mean and SD values of fati gue indices before/after the fatigue procedure. Before After Fatigue indices Mean (SD) Mean (SD) Gender: p Fatigue: p Fatigue Gender: p Knee extensors (%) 14.9 (9.0) 20.2 (11.2) 0.253 (0.2) 0.033 (0.58)* 0.763 (0.06) M 16.4 (9.3) 22.4 (8.2) F 13.4 (8.8) 17.9 (13.4) Knee flexors (%) 10.6 (8.8) 18.1 (10.7) 0.887 (0.05) <0.001 (1.0)* 0.3 (0.17) M 9.5 (9.0) 18.8 (13.1) F 11.7 (8.7) 17.5 (8.1) Significant main effect or interaction (p<0.05)

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28 Figure 2-1. Correlations of power and sample si ze for each combination of variables. Note: B (between-subject variable), W (within-subj ect variable), BW (i nteraction between B and W).

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29 Figure 2-2. Experimental setup.

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30 Figure 2-3. A subject with markers on.

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31 Figure 2-4. Overview of th e experimental procedures.

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32 Figure 2-5. Marker placement (lef t) and definition of regional angles of the spine (right).

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33 Figure 2-6. Kinematic variables defined by the cr itical instants identified from kinematic and forceplate data.

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34 CHAPTER 3 LITERATURE REVIEW The purpose of this study was to investigate the mechanical characteristics of the spine during drop landings using differe nt landing techniques and fatigue status of lower extremity muscles. A biomechanical model of the lower extremity and spine was employed to study the kinetics and kinematics of the spine and lower ex tremity joints. Research investigating spinal mechanics using in vitro and in vivo techniques is quite extensive while studies of spinal mechanics in jumping and landing are quite limited. Previous in vitro research focused on range of motion of spinal segments in various conditions. Previous in vivo research focused on the spinal mechanics during common, everyday activ ities without considering lower extremity activities. More specific knowledge about spinal mechanics dur ing vigorous physic al activities is necessary to have a better unde rstanding of spinal mechanics when the spine is under dynamic loadings. A description of spinal mechanics during drop landings is not currently available in the literature. Early studies to investigate spine biom echanics were mostly focused on defining mechanical properties of spinal structures and segments in various conditions using different instrumentations. Recent in vitro research focuses on the cause -and-effect relationship and mathematical modeling of specific situations with cadaveric spinal segments using computational programming and statistical procedur e to explain the complicated situation and to develop the spinal instrumentation by minimi zing and simplifying the mechanical conditions. Recent in vivo research focuses on estimating and verifying the results from in vitro and clinical studies using optoelectrical system s, EMG, forceplate and so on. To assist the reader in unders tanding the current issues of research on spinal mechanics, the literature review wi ll be presented under the following he adings: biomechanical properties of

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35 spinal structures, biomechanical properties of spinal segments, biomechanical performance of spine in vivo biomechanical etiology of spinal pain, biomechanical performance of painful spine, and landing biomechanics. Biomechanical Properties of Spinal Structures The mechanical properties of the bony structures (including facet joints), ligaments, spinal muscles and intervertebral disc of each spinal le vel have been thoroughly identified in previous studies. The intervertebral disc ha s received more attenti on than the other spinal structures due to its specific anatomical and biomechanical features Defining disc characteri stics is the first step in reviewing the mechanical properties of the spine. Intervertebral Disc An intervertebral disc consists of three components: the nucleus pulposus, the annulus fibrosus, and the cartilaginous end-plate. However, there is no clear landmark to differentiate the nucleus pulposus and annulus fibrosus because the peripheral region of the nucleus pulposus merge with the inner region of the annulus fibros us. The nucleus pulposus is a centrally located mucoid material in semi-fluid state, and its water content of which ranges from 70-90% (Panagiotacopulos et al., 1987). The annulus fibr osus forms the outer region of the disc, and consists of collagen fibers in a highly ordere d pattern. The collagen fibe rs are layered in 10-20 sheets called lamellae and arranged in a helicoid manner. They run in the same direction in a given lamella but in opposite directions in adja cent lamella. The lamellae are thick in anterior and lateral regions of the annulus, and thin poste riorly (Inoue, 1981). The end-plate is composed of hyaline cartilage about 0.6-1.0 mm thick which se parates the other two components of the disc from the vertebral body (Roberts et al., 1989). The e nd-plate covers the entire nucleus pulposus, but does not cover the entire annulus fibrosus peripherally. Instead, the ring apophysis, which is part of a vertebral body, covers the peripheral region of the annulus fibrosus. Because of the

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36 attachment of the annulus fibrosus to the end-pl ate, the end-plate is strongly bound to the disc, but weakly attached to th e vertebral body (Inoue, 1981). The basic functions of the disc are to transm it loads from one vertebral body to the next and to allow movement between vertebral bodies All components of the disc are involved in weight-bearing. When an axial compressive load is applied to a nucleus the nucleus tends to reduce the height and expand radial ly towards the annulus fibrosus. This radial expansion exerts a pressure on the annulus which tends to stretc h its collagen lamellae outwards. However, the tensile properties of the collagen resist this stretch, and the lamellae oppos e the outward pressure exerted by the nucleus. Application of a 400 N load to an intervertebral disc causes only 1 mm of vertical compression and only 0.5 mm of radial expansion of th e disc (Hirsch & Nachemson, 1954). The nucleus pressure is also towards the e nd-plates, and constrained by the end-plates and vertebral bodies. The pressure on th e end-plates serves to transmit the part of applied load from one vertebra to the next, and the radial pressure on the annulus fibrosus braces it and prevents the annulus from buckling (Roaf, 1960b). Brown and colleagues (1957) conducted static te sts to compare the relative strength of the disc with that of the vert ebral body in compressive loads, without the posterior elements. They found the first stru cture to fail in such a construct was the vertebra, instead of the intervertebral disc, because of the fracture of the end-plates. They also observed no difference between the vertebrae with normal discs and those with degenerate d discs. The mode of failure by the pure compressive loads seemed to be mostly dependent on the condition of the vertebral body (osteoporosis of the vertebrae), not on the condition of the disc. During distraction, all points on one ve rtebral body move an equal distance perpendicularly from the upper surface of the ot her vertebral body. Consequently, every collagen

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37 fiber in the annulus fibrosus is equally strained, and resists distraction. Ho wever, the disc is not often subjected to the tensile loads under normal physiologic activ ities. Also, even under the distraction of the spine, the discs are under the compression load due to spinal muscle activity. However, the annulus fibrosus is subjected to te nsile stresses in various physiologic activities. In addition to compression, any direction of the be nding (flexion, extension, and lateral bending) moves the instantaneous axis of ro tation to lie outside of the disc and the disc is subjected to the tensile stress at the opposite side of the bending (White & Panjabi, 1990). Bending involves lowering one end of the vertebral body and raising the opposite end. This causes distortion of the a nnulus fibrosus and the nucleus pulposus. In forward bending, the anterior annulus is compressed and the disc tend s to bulge anteriorly. Th e nucleus pulposus is also compressed anteriorly, but the elevation of the posterior end of the vertebral body relieves the pressure on the nucleus posteriorly (Brown et al., 1957; Shah et al ., 1978). However, Roaf (1960a) did not find any changes in shape or posi tion of nucleus pulposus on the nucleographs of the disc during flexion/extension. This supports th e relevance of maintain ing a slightly flexed lumbar spine posture as a treatment and prophylax is for the patients with low back pain. The increase in disc pressure observed in vivo during bending of the lumbar spine may not be just from bending, but from the result of compressive loads applied to the di scs by the action of the spinal muscles which are involved in bending motion (Ortengren et al., 1981). During torsional movement of the inter-body joint, only the co llagen fibers in the annulus in the direction of movement have their poin ts of attachment separated. Thus the annulus fibrosus resists torsional moveme nts with only half number of lamellae. Farfan et al. (1970) conducted experiments using cadaveric vertebra -disc-vertebra construct including posterior structures to examine the effect of torsional load. They found that the failures occurred at the

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38 annulus in the final phase of the loading. The average failure torque for the normal discs was found to be 25% higher than that for the degenera ted discs, and the average torsional angles at failure were 16 and 14.5 for normal and degenerated discs, respectively. Torsion of inter-body joints seems be the most likely mechan ical factor to injure the annulus. In pure shear movements of the inter-body joint, only half of the fibers in the annulus are strained, and the shear stress is ra ised mostly at the side of the loading. The shear stiffness in the horizontal plane was found to be about 260 N/mm, a nd this could be the large force to cause an abnormal horizontal displacement in the normal disc (White & Panjabi, 1990). This means that it is rare for the annulus to fail cl inically due to pure shear loading. In addition to the load characteristics, th e intervertebral disc demonstrates the time dependent behavior which is called the viscoelastic property: creep and hyst eresis. If a constant force is applied to a viscoelastic structure for a prolonged time, further movement will be detectable after the end of physio logic motion. If this movement is small in amplitude, occurs slowly, almost imperceptible, then it is called as creep. Kazarian ( 1975) performed compression creep study on functional spinal uni ts (FSUs) and differentiated th e disc specimens into four grades according to the degree of degeneration. He found the creep and degeneration grade of the disc are related. The non-degenerated discs creep sl owly and reach their final deformation after a long time, compared with the degenerated discs. This means that degenerated disc loses the capability to attenuate shock and to distribute th e load uniformly over the entire end-plate. Viscoelastic structures also show differences in mechanical behavior during loading and unloading. Restoration of the initia l length of a structure from unloa ding occurs at a slower rate and to a lesser extent than did the deformati on from loading. This difference in mechanical behavior is referred to as hyste resis, and reflects the amount of energy lost compared with

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39 structure at initial loading. Virg in (1951) observed that hysteres is is largest in young people and smallest in the middle-aged ones, and the lo wer thoracic and upper lumbar discs show less hysteresis than the lower lumbar discs. He also found hysteresis decreased when the same disc loaded repeatedly. This means that the disc is less protected against repetitive loads. When forces are repeatedly applied to a material, it does not behave the same way each time. Each application produces a certain amount of hysteresis, and the material is altered slightly. Following many repetitions small weaknesses accumulate and weakness in the material becomes apparent. After several frequent repetitions of a stress, the material may fail at a certain stress which is less than that required to dama ge the material following a single application of a force. This is referred as the fatigue failure, and the fatigue tests of the disc were developed to identify the number of load cycles that can be tolerated before disc failure develop. Brown et al. (1957) conducted a fatigue test on the disc with a small constant axial load and a repetitive forward bending of 5 The disc failure started to o ccur after 200 cycles of bending, and complete failure occurred after 1000 cycles. Ho wever, the fatigue tolerance of the disc in vivo is not known. Vertebra The basic morphology of the vertebrae in various regions of the spine from C3 to L5 is approximately the same. The size and mass of the vertebrae increase from C1 to L5 vertebra. This is the mechanical adaptation of the vert ebrae to the progressively increasing compression loads. Chalmers et al. (1966) observed that the strengths of vertebral ca ncellous bone of each lumbar segment are almost the same. This means the variation of the vert ebral strength according to the spinal level is mostly due to the size of vertebrae. However, ve rtebral strength decreases with age. Bell et al. (1967) found that a small loss of osseous tissue produces much loss of

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40 vertebral strength: a 25% loss of the osseous tissue results in a more than 50% decrease in the vertebral strength. The bone mineral density of fema le vertebrae is less than that of male at any corresponding age, but the rate of decrease is not different between males and females (Hansson & Roos, 1980). The vertebral body is designed for load-bearing of large compression. Fi rst, the vertebral body is not a solid bone block, but a shell of cortical bone and can cellous core. Second, the space between the trabeculae in the cancellous core can be used as the channels for the blood supply and venous drainage of the vertebral body. Li kewise, the presence of bone marrow in the intertrabecular spaces acts as a useful elemen t for transmitting the loads and absorbing the force (White & Panjabi, 1990). Rockoff et al. ( 1969) conducted compression strength tests on two groups of vertebrae without posteri or elements: the vert ebral bodies with ce ntral hollow and the other ones without outer shell. They found the loss of strength in both specimens, and the sharing of the compressive load by the cortical shell was about 65%. Facet joints are important stabilizing structur es and carry about 18% of total compressive load in the lumbar spine region (Nachemson, 1960) However, King et al. (1975) pointed out that the load-sharing between facet joints and disc is much complex in their dynamics study with whole cadavers. They observed that the load sh aring carried by the facets could be 0-33% and the value depended on the spine posture. Facet joints also contribute to tors ional strength of FSU. Farfan (1970) observed that th e vertebral body-disc-body with l ongitudinal ligaments share the torsional strength equally with the two facet s and capsular ligaments, about 45% each. The remaining 10% was carried by the interspinous ligaments. Spinal Ligaments The ligaments from C2 to the sacrum are similar, and seven ligaments are generally referred to as the spinal ligaments in this region: the anterior a nd posterior longitudinal

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41 ligaments, the intertransverse ligament, the caps ular ligament, the ligamentum flavum, and the interspinous and supraspinous ligam ents. The anterior and posterio r longitudinal lig aments attach to the anterior and posterior edge s of vertebral bodies and discs fr om basiooccipital to the sacrum and coccyx. These ligaments become deformed by the relative separation between adjacent vertebrae and by the bulging of the disc. Tkaczuk (1968) observed that the anterior longitudinal ligament was twice stronger than the posterior longitudinal ligament. The intertransverse ligament is known to have no mechanical significan ce in the lumbar region because of its small cross-sectional size (Chazal et al., 1985). The ca psular ligaments contribute to the flexion stability in the cervical spine region (Panjabi et al., 1975). The ligamentum flavum has a high percentage of elastin (80%) when compared to th e other ligaments. This allows a large extension without permanent deformation (Yahia et al., 1990). Nachemson and Evans (1968) found that the ligamentum flavum has pre-tension, and this produces resting compression of the disc. Consequently, the high elasticity and pre-tens ion of the ligamentum flavum minimizes the chances of any impingement to the spin al canal during sudden spine motions. Biomechanical Properties of Spinal Segments Multisegmental Mechanics of the Spine The basic motion segment of the spine is refe rred to as the functiona l spinal unit (FSU), which consists of two adjacent vertebrae, inte rvertebral disc, and the connecting ligaments without the spinal musculatures. Generally, a FSU exhibits similar biomechanics to those of the entire spine, and can be used as a common testing specimen in vitro The range of motion (ROM) of a FSU is represented by the sum of two distinct phases: neutral zone and elastic zone. The neutral zone is defined as the low-load re sponse of FSU near the neutral position, and the elastic zone is defined as the spinal behavior beyond the neutral zone up to the end of the physiologic limit (Figure 3-1) (White & Panjabi, 1990).

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42 Generally, the neutral zone is referred to as a quantitative measurement of the laxity around the neutral position of a FSU It is known to increase with degeneration, surgical injury, repetitive cyclic loads, and high-speed trauma In flexion/extension and lateral bending, the neutral zone is the larg est in the lower cervical region. In axial rotation, th e neutral zone is the largest in the C1-C2 region (Table 3-1). By defining the neutral and el astic zones in the load-displacement curve of an FSU, the coefficient of flexibility and stiffness can be calc ulated. The flexibility coefficient is defined as the ratio of the displacement produ ced to the load applied. The stiffness coefficient is defined as the ratio of the resistance offere d to the displacement imposed. Ho wever, the load characteristics of the spine is quite complex ( nonlinear, biphasic, and viscoelastic) and cannot be demonstrated by a single number. Previous studies showed that FSUs are more flexible in tension than in compression in all regions of the spine. The shear flexibility is not qui te different in each direction (anterior, posterior, or la teral). The spine is more flexible in flexion than in extension in all regions except the sacroiliac joint. Flexibilit y values for lateral bending are in between the values of flexion and extension (Panja bi et al., 1988; Pa njabi et al., 1976). Axial rotation is generally known to be more harmful to the disc than the other motions, except for a combination of axial rotation and lateral be nding (Farfan et al., 1970). In the cervical region, the spine is about 37% as flexible in torsion as compar ed with flexion. In the upper thoracic region, the torsional stiffn ess is about the same as in flexion. The torsional flexibility of the lumbar region is about 27% of flexion flexib ility, which is the lowest value in all regions. However, the torsional flexibility is much greater at the lumbosacral joint (55% of flexion) and sacroiliac joint (250 % of flexion) (White & Panjabi, 1990).

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43 Individual physiologic motions of FSUs are inherently conne cted, and which is called the coupling. Coupling of spinal motions is due to th e geometry of individual vertebrae, connecting ligaments, discs, and the curvature of the spin e. The motion produced by an external load is defined as the main motion, and the accompanying motions are called the coupled motion. In the thoracic region, there is a st rong coupling between all the motions in the sagittal plane (translation and rotation) (Panjabi et al., 1976) The coupling of axial rotation with lateral bending is a very common physiologic motion in the cervical and lumbar regions (Moroney et al., 1988; Panjabi et al., 1977). Regional Mechanics of the Spine The entire spine is divided into cervical (upper C0-C1-C2, middle C2 C5, lower C5 T1), thoracic (upper T1 T4, middle T4 T8, lo wer T8 L1), lumbar (L1 L5), lumbosacral (L5 S1) and sacroiliac regions, based on the kinematic, kinetic, a nd clinical characteristics. The upper cervical region is composed of occi pital-atlanto-axial join ts (C0-C1-C2) and is the most complex region of the spine, anatomica lly and kinematically. Most of the axial rotation and some of the flexion-extension and lateral be nding of the head come from the upper cervical movements. The dominant atlant ooccipital (C0-C1) motion is mo stly flexion/extension, some lateral bending, and tiny axial rotatio n (Table 3-2). The atlantoaxial (C1-C2) articulation consists of four joints: two atlantoaxial lateral joints, the atlantoaxial median joint (between anterior arch of the atlas and dens axis), and a joint betw een the posterior surface of the dens and the transverse ligament. The lateral atlantoaxial joint capsule is loose and allows a great deal of axial rotation, in which the vertical ax is of dens acts as a pivot abou t the atlas rotation. The possible atlantoaxial motions are also summarized in Table 3-2. The anatomical structures and the function of the middle and lower cervical regions are quite different to those of the upper cervical re gion. The dominant motion in the lower cervical

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44 spine is flexion/extension. Lysell (1969) observe d the routes of each cervical vertebra in the sagittal plane from flexion to extension or vice versa. The movement is a combination of translation and rotation, and he cal led that movement the top angle, which indicated the arch steepness of the route, generated during the fl exion/extension. The arch es were flat at C2, steepest at C6 and followed by C7. The average RO M in healthy adults in the middle and lower cervical regions are summarized in Table 3-3. Dvorak et al. (1992) performed an in vivo test to measure the cervical ROM, based on the modified inclinometer technique, to define the ageand gende rrelated differences. They found a general tendency of decreasing cervical ROM as the age increased. The most drastic decrease in cervical motion occurred at th e age of 30-39 and 40-49 years. The cervical motions that did not decrease with age were the rotation out of maximum flexion and the upper cervical rotation. Females showed greater ROMs in all planes of cervical motion. However, there were no significant differences between genders for the group over 60 years. Th e typical cervical ROM values are presented in Table 3-4. The human thoracic spine is a unique spinal re gion to be adapted to an erect posture and load-bearing. The predominant posture of the th oracic spine is a kyphotic curve while the last region (T11 L2) is almost straight in the sa gittal plane. Thoracic kyphosis may arise from postural and structural factors. Postural factor come from positioning of the spine due to the ligamentous tension and muscle t one, as well as the disc configur ation. The shorter ventral height of thoracic vertebral bodies than dorsal one contri butes to the structural kyphotic curve in the thoracic region. Due to the kyphotic curve of the t horacic spine, the axial load applied to this region generates a bending moment to cause furthe r flexion. As a stability of this region, dorsal tension-band capacity by the posterior ligamen ts and ventral weight-bearing by the vertebral

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45 bodies are the vital combination to prevent sp inal deformities in this region (Benzel & Stillerman, 1999). There are two transitional regions in th e thoracic spine: cervicothoracic and thoracolumbar junctions. The biomechanical ch aracteristics of these junctional regions are described as a blending of two adjacent regions. The upper thor acic region has a very limited flexion/extension, whereas the caudal region from T10 has a larger range of flexion/extension. The sagittal orientation of the facets in the lo wer thoracic region severely limits axial rotation and, to a lesser extent, laterals bending. The f acets of the upper thoraci c spine are similar in orientation to those of the cervi cal spine, and similar motion ch aracteristics occurred at the cervical and upper thoracic regions. Likewise, the facets of the lower thoracic spine are similar to those of the lumbar spine and similar motions are seen in both lower thoracic and lumbar regions. Representative ROMs of different mo tion segments of the thoracic spine are summarized in Table 3-5. The pattern of motion in the sagittal plane for the thoracic spine is similar to that in the cervical spine. To describe the mo tion of the thoracic vertebra in the frontal and sa gittal planes, the top angle was also employed. In the sagittal plane motion, the arch is quite small, and there is no variation according to the level. The arch in the frontal plane is also flat, but greater than that in the sagittal plane. Also in the frontal plane, the arch tends to increase from T1 to T12. There is also coupling of latera l bending and axial rotation in the thoracic region. The pattern of coupling in the thoracic region is similar to the one in the cervi cal region. However, the coupling pattern in the lower thoracic regi on is not as strong as that in the cervical region (White, 1969). Movements in the thoracic spine are greatly limited by the facet orientation and by the rib cage. The costotransverse and cost overtebral articulati ons provide strong, stable attachment of

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46 the thoracic vertebrae to the ribs. The costosternal articulations also contribute to the stability of the thoracic spine (Pal & Routal, 1987). Oda et al. (1996) found a signi ficant increase of flexion/extension in thoracic moti ons after resection of the posteri or elements. With the removal of the costovertebral joints bilaterally, large in creases in lateral bending and axial rotation were observed. They concluded that the integrity of the costovertebral join ts and the rib cage significantly contribute to the spinal stability of the thoracic region. In the sagittal motion of the lumbar spine, there is a cephalocaudal increase in flexion/extension. The lumbosacral joint provide more sagittal plane motion than the other lumbar motion segments do. For the coronal moti on, the ROM for each lumbar level is about the same, except for the lumbosacral region, which demonstrates limited lateral bending. Limitation of lumbosacral motion is about the same for th e axial rotation. Repres entative ROMs of the lumbar spine are summarized in Table 3-6. A nother important kinematic component in the lumbar spine is the sagittal plane translation. In general, 2.0 to 2.8 mm is referred to as the normal limit of anterior translation ofr a lumbar spinal vertebra, and 4.5 mm is an evidence of clinical instability for a lumbar motion segment. There are several patterns of coupling motion in the lumbar spine. Pearcy et al. (1984) observed coupling of slight axia l rotation and lateral bending w ith flexion/extension, based on their stereoradiographic study. Another coupling pa ttern is that of la teral bending and axial rotation. The direction of lateral bending with ax ial rotation in the lumb ar region is that the spinous processes point to the same direction as the lateral bending. It is opposite to the pattern occurred in the cervical and thoracic regions. However, the coupling pa ttern of lateral bending with axial rotation at the lumbosacr al joint is the opposite of that in the lumbar region and similar to the cervical and thoracic re gions (Pearcy & Tibrewal, 1984).

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47 The sacroiliac joint is partly synovial and partly syndesmotic. It is known to be completely ankylosed in 76% of the subjec ts over 50 years of age (White & Panjabi, 1990). However, studies of the motion about the s acroiliac joint have pr oduced a wide range of results. Miller et al. (1987) performed a kinematic study of the s acroiliac joints in cada veric specimens. They measured the displacements of the sacrum in re lation to the ilium with each plane of loading. They observed that lateral bending of one side was 1.4 anterior translation 2.74 mm, and axial rotation of one side 6.21 in their study. Walheim et al. ( 1984) observed 2-3 mm of vertical translation and 3 of rotation of the pubis at the sy mphysis pubis with one leg standing. Sturesson et al. (1989) al so observed tiny motions of sacroiliac joints in their stereoradiographic study. Biomechanical Performance of Spine In Vivo Trunk Posturing Trunk postures during various activities are rela ted to the risk of developing a low back disorder (Granata & Wilson, 2001). Spinal comp ression below 3,400 N may be considered as a safe margin to prevent low back disorder in occupational population (Konz, 1982). However, a spinal injury associated with the instability can occur at a lo w compressive load (Granata & Marras, 1999). This means that the appropriate recruitment of the spinal and trunk muscles provide the stable support for th e large load applied to the spin e, but some postures may limit the ability of the muscles to maintain spinal stability (Wilke et al., 1995). Nachemson (1966; 1981) conducted studies to ve rify the compression load applied to the disc in vivo by measuring the intradiscal pressure at L3 /L4 level in various postures. The lowest compression was observed in the supine position ( 300 N), which is about 50% of that in standing (700 N) without external loads. During sitting without a back support, the compression went up

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48 to 1,000 N. Forward bending of the trunk generated the largest compression force to the spine up to 1200 N. Ledet et al. (2005) meas ured intradiscal loads in baboons in vivo and found similar results to Nachemsons. Takahashi et al. (2006) also performed the same testing at L4/L5 level with young, healthy subjects, and found similar re sults. However, their values in specific postures are different from those reported by others. The values of compression loads in different studies are summarized in Table 3-7. The data from Takahashi et al. (2006) were greater than those from Nachemson (1966; 1981). Takahashi et al. explained the differences with several r easons. Their subjects were all young and normal subjects. Their techniques seemed to be much more developed than the old one used in Nachemson studies. The specific level used in intr adiscal measurement was lower than that in Nachemson studies. Takahashi et al. concluded that the risk of intervertebral disc injuries or degeneration could be induced by a simple repetitive forward bending of the trunk in everyday movements. However, slightly flexed, relatively straight or nonlordotic position of the lumbar spine during standing is used more freque ntly in the populations who complain less often about back pain. Also, the flexion of the hip re duces the tension of th e psoas muscles and the lordosis of the lumbar spine, resulting in reduced load on the lumbar spine. In addition to the forward bending of the tr unk, twisting, lateral bending and asymmetric posture combined with lifting are also known to increase the risk of lo w back problems (Kelsey et al., 1984; Marras et al., 1993). So me EMG studies explained that the high loads to the spine in the unstable and asymmetric postures are from th e co-contraction of agoni stic and antagonistic spinal muscles (Cholewicki et al., 1997). Weight Lifting Lifting and bending episodes accoun ted for 33% of all work-related causes of back pain (Damkot et al., 1984). Increasing we ights anterior to the spinal column greatly increases the

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49 forces which are exerted on the lumbar spine. This is due to the forces developed in the spinal muscles in order to maintain the equilibrium. The resultant forces applied to the fulcrum, which is the lower lumbar region, are very high. The distance of the weight from the body is directly related to th e joint reaction force (high disc pressure) to the lower back, the gr eater force required by th e erector spinae muscles (high electromyographic activ ity), and a need of gr eater truncal support to protect the spine (high truncal pressure). This indicates the significance of closer distance of the object to the body as the proper lifting technique. Anothe r factor to determine a proper li fting is the back posture. For the optimal lifting methods, the squa t lift (knee bent and back strai ght) is generally considered to be safer than the stoop lift (knee straight and b ack bent) in bringing the load closer to the body, and reducing the back muscle demands to c ounterbalance the additional moments. These techniques consider the posture of the back in addition to the distance of the objects. However, many workers prefer the stoop lift ov er the squat lift. There is an increased physiologic cost and more rapid fatigue development in a squat lift. And the squat lift is not always possible because of the lift setup and load size. Likewise, the ri sk of developing low back pain by lifting tasks depends rather on the lumbar pos ture than the choice of lifting techniques (van Dieen et al., 1999). There is a conflict about the favorable lu mbar postures during lifting tasks. Some advocate lordotic and straightening lumbar posture because they believe increased erector spinae activity is beneficial in augmenting spinal stab ility and decreasing anterior shear force on the spine (Hart et al., 1987). Howeve r, others favored the kyphotic lift (flexed lumbar spine), because they believe passive ligaments of the lumbar spine can relieve the active extensor muscles (Gracovetsky et al., 1981). Cholewicki et al. (1992) tested professional power lifters to

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50 evaluate the kinematics of the lumbar spine and resultant posterior ligament lengths during lifting tasks. They observed significantly smalle r lumbar flexion and increased lengths of the ligaments during the lifting when compared with the full flexion of the trunk. They concluded that the back muscles were substantially res ponsible for resisting trunk flexion moments during heavy lifting. Arjmand and Sirazi-Adl (2005) tested the kine matics of the lumbar spine and activity of selected spinal and trunk muscle s in healthy subjects during a stat ic lifting. They examined the lumbar spine postures (lordosis, kyphosis) during the lifting procedures and found the lordotic posture increased extensor muscle forces, axia l compression and shear fo rces at L5/Sl. They recommended the moderate flexion posture of the lumbar spine as a posture of choice in static lifting tasks. Lifting capacity is generally used to determine the degree of spinal impairment state, and the back strength and aerobic capacity are known to be the contributing factors (Matheson et al., 2002). Sitting and Standing The back rest and lumbar support is known to decrease the loads applied to the spine during sitting. Andersson et al. (1977) performed a study to estim ate intradiscal pressure of L3/L4 under different backrest inclinations and lumbar suppor t conditions. They found the highest intradiscal pressure in the sitting with no lumbar support and a 90 backrest inclination, and the lowest disc pressure and the least electr omyographic activity of the paraspinal muscles in the sitting of 120 -inclination and 5-cm widt h of lumbar support. More specifically, the lumbar support has the greater influence on lumbar lord osis and the backrest inclination has more influence in reducing the loads on the lumbar disc (Andersson et al., 1979). The arm rest is also known to reduce intradiscal pre ssure (Kelsey & Hardy, 1975).

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51 The ability to stand up from sitting is a prer equisite for walking and other independent function. Sit to stand demands coordinated movements of linked body segments to transport the center of body mass in a horizontal then vertical direction while mainta ining balance over a small base of support, the feet. Previous studi es reported these kinema tic characteristics of standing from sitting: flexion of the trunk and hi ps bring the center of mass forward, followed by bilateral extension of th e lower limb joints, and trunk extens ion to raise the body in a vertical direction over the feet (Doorenbosch et al., 1994) Tully et al. (2005) studied the kinematics of the body segments including the thoraco-lumba r region during standing from sitting. They divided the standing movement into phases before and after the lift-off, based on the relation of position between the buttock and the sitting object. Before the lift-off, they observed a forward leaning of the trunk, accomplished by concurrent lumbar and hip flexion (1:3). As the lumbar spine flexed the thoracic spine extend ed, resulting in a trunk angle of 45.7 at lift-off with respect to the horizontal plane. Following the lift-off, th e hip and lumbar spine extended and the thoracic spine flexed, with the standing thoracic angle appr oximating the initial thoracic posture in sitting. Walking The biomechanical function of the trunk during walking has been investigated extensively. Earlier studies examining the trunk kinematics du ring walking considered only the entire trunk motions with respect to the pelvis. Later studi es examined the movements of the lumbar and thoracic spines or pelvis. Cr osbie et al. (1997) studied the patterns of spinal motion during walking using a model including upper and lower trunk, lumbar and pelvis segments. They used three spatial surface markers in each spinal segment on the back surface of the subjects. The pattern of flexion/extension of each segment wa s generally biphasic throughout the gait cycle. The pelvis rotated into negative pelvic tilt at heel strike. This was followed by a counter-motion

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52 to a maximum positive pelvic tilt in the si ngle support phase. The lumbar spine reached maximum flexion at heel strike. This was followed by a rapid extens ion to neutral until the single support phase. The lower thoracic segment extended maximally at heel strike, and returned to a neutral at mid-stance, then extended again through the late stance. The counter-motions occurred between the lumbar and lower thoracic segments at heel strike. They concluded that spinal segments demonstrate complementary movements to the motion of the pelvis, and pelvic motion responds to the need of advancing the lower li mb and transferring the body weight from one supporting side to the other. Lumbar ROMs dur ing walking and running are summarized in Table 3-8. Thousands of repetitive low level loadings are applied to the spine in everyday activities. During normal walking, activation of the spinal muscles, accel eration of the trunk, combined with the external loadings result in cyclic spin al loads. Some studies i nvestigated the magnitude of these loads, in conjuncti on with the spinal motion and mu scular activities during walking. Callaghan et al. (1999) conducted a biomechanical study using two models to estimate loads applied to the L4/L5 level: linked segment mode l with EMG technique and rigid segment model with inverse dynamic technique. The joint load ing (at L4/L5) calculated by the EMG model resulted in large increases in the maximum comp ressive forces, compared with the joint reaction forces calculated using inverse dynamics. Including the muscular component resulted in a more than three-fold increase in joint load. The comp ression loads applied to the lower lumbar level during walking from different studies are summarized in Table 3-9. Unlike the differences in compressive load, the joint shear forces (anterior/posterior, lateral) obtained using the two techniques were similar. The flexion/ex tension moment had two peaks throughout the gait cycle. At heel contact there was a peak fle xor moment followed by a

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53 peak extensor moment around toe-off. During a fa ster speed gait, the flexion/extension moment at L4/L5 shifted to the extension side and demo nstrated a high extensor moment around toe-off. The lumbar spine motion with respect to the pe lvis was quite consistent within and between subjects. The sagittal motion of the lumbar spine showed several dominant phases. Following heel contact a flexion phase was present unt il the relative spine motion reached maximum flexion just following toe off. And the spine re mained in a constant posture and additional extension phase during the single stance. They concluded that the loads and motions for the lumbar spine during gait depended on the walkin g speed. Increasing walking speed increases the lumbar spine ROM, activation of spinal and trun k muscles, and anterior/posterior shear forces. Running Running typically requires the spine moving through only a limited range of motions. Acute injuries to the spine directly from runni ng activities are relatively infrequent and amounted approximately 11 13% of all injuries sustai ned (Walter et al., 1989). The frequent and significant spine injuries related to running are largely due to the repetitive axial compressive loading of the spine which occurs during the foot strike in each stride. A typical distance runner who runs 130 km per week in training might subject to about 40,000 f oot strikes per week (Cavanagh & Lafortune, 1980). Several case studies ha ve highlighted the overuse injuries of the lumbar spine and pelvis in running (Guten, 1981; Koch & Jackson, 1981). Schache et al. (2002) conducted a kinematic study of lumbar spine and pelvis during running. The average ROMs of the lumbar spine and pelvis on each plane are summarized in Table 3-8. They found significant inverse correlati ons between flexion/extension of the lumbar spine and anterior/posterior tilt of the pelvis and lateral bend of the lumbar spine with obliquity of the pelvis. Essentially, as anterior tilt of the pelvis increased during the terminal stance, extension of the lumbar spine also increased. When the lumbar spine was laterally bent to the

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54 same side of the foot contacted to the ground, the pelvis was lowered on the opposite side. When the lumbar spine began to laterally bend toward s the opposite side of the foot contacted, the pelvis began to elevate on the opposite side of the foot contacted. They also found a significant positive correlation between the axial rotations of lumbar spine and pelvis. However, coordination between the axial rotations of the lu mbar spine and pelvis was out of phase by 21% of the running cycle. The kinematic pattern of axial rotation of the pelvis during running is different to that during walking. At the initial co ntact of one foot during walki ng, the pelvis showed maximal rotation to the opposite side (W hittle & Levine, 1999). This move ment helps in augmenting the stride length at that time. With the loss of the double support phase during running, the pelvis along with lower extremities are no longer required to be enga ged in a stride lengthening mechanism. At the initial contac t during running, the pelvis was rota ted to the same side of the foot contacted. This movement was suggested as minimizing the horizontal braking force at the initial contact to avoid potent ial loss of running speed (Novach eck, 1998; Schache et al., 1999). Schache et al. (2003) performed another ki nematic study of lumbo-pelvic-hip complex during running to define the gender differences They found that females displayed a shorter stance time, swing time, stride time and stride le ngth, and a higher stride rate than males. Mean waveforms were different in the peak-to-peak oscillations and the offset of pelvis anterior/posterior tilt. Females displayed grea ter amplitudes of lumbar spine lateral bend and axial rotation, pelvis anteri or/posterior tilt, obliquity and axia l rotation, and hip adduction/abduction than their male counterparts. The mean positions of anterior pelvic tilt across the running cycle were 20.2 for females and 16.9 for males. The prevalence of pelvic-

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55 femoral stress fractures in female runners might be explained by these fi ndings (Bennell et al., 1996; Pavlov et al., 1982). Wilke et al. (1999) studied the intradiscal pressure of non-degenerated disc at L4/L5 in a 45-years old man during va rious activities. They observed intr adiscal pressure of 0.5 MPa during relaxed standing and 0.35-0.85 MPa during jogging with tennis s hoes. Rohlmann et al. (2001) also recorded a peak intradiscal pressure 0.85 MPa while jogging on the treadmill of, which was 170% of the pressure noted in standing. The intr adiscal pressures recorded in different studies are summarized in Table 3-10. There are several factors suggest ed in previous studies that affect the spin al posture and loading during running. Exhaustive running, as fatigue occurs, has shown biomechanical changes in the legs that lead to a lower effective body mass during heel strike (Der rick et al., 2002). This resulted in increased peak leg impacts and in creased shock attenuation, which may change the spine loading forces with fatigue. Therefore, th e loads applied to the spine may vary during a run, depending on the lower extrem ity activity and the level of fatigue (Lennard & Crabtree, 2005). Another factor that affects the spinal loading during running is the type of shoes and insole materials utilized. An impulsive shock wave is gene rated at heel strike th at is transmitted from the lower extremities through the spine. The use of shock absorbing insoles has been used to treat low back pain patients to lower the shock wave at low back level. On the other hand, the development of external force and the transmi ssibility of impact fo rces through the human body are increased by wearing soft so les. Ogon et al. (2001) perfor med a study to investigate the behavior of low back muscles du ring jogging barefoot and wearing id entical athletic shoes. They observed that wearing shoes and insert materials decreased the rate of shock transmission to the

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56 lower back and reduced the time interval betw een peak acceleration of lower back and peak spinal muscle response in jogging. It is from the increased latency between heel strike and peak acceleration at the lower back by wearing shoes. They suggested weari ng shoes decreases the time interval between maximum external (peak acceleration at lower back) and maximum internal force (generated by the spinal musc les) in the lower back during running. Wen and associates (1997) found a signifi cant correlation between leg lengt h discrepancy and onset of lumbar pain within 12 months of running in a marathon training program. However, the clinical importance of leg length discrepancy in short distance running is not clear. Biomechanical Etiology of Spinal Pain Spinal pain can be caused by trauma, infecti on, tumor, and systemic diseases. However, the term spinal pain is generally used to refer to the cervical, t horacic, and lumbar pain that is not related to these injuries and diseases. The common neck pains with or without arm pain and the back pain with or without leg pain which are frequently encountered in daily livings and cause the spinal degeneration are called spinal pain. Although ther e are specific considerations associated with spinal pain in different regions of the spine, there is much similarity in different spinal regions. Spinal pain usually occurs in th e more mobile and lordo tic portions of the spine and onsets at 30-50 years of age. Spinal pain occu rs most often in the lumbar region, followed by the cervical region, with the lowe st incidence in the thoracic re gion. Spinal pain can come from direct irritation of the nerves that innervate most of spinal structures. The pos terior annulus fibrosus, the posterior vertebral body and the pos terior longitudinal ligam ent are innervated by the sinu-vertebral nerve, which is considered the most common origin of spinal pain (Bogduk & Twomey, 1997). Another type of sp inal pain is the indirect re ferred pain which is not fully understood and explained at this moment. Spinal pain is a major socioeconomic problem and approximately 80% of all back problems are of unknown origin (Vogt et al., 2001).

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57 Specific motion, force, and high-quantity repeti tive loading, or any combination of these may serve as mechanical stimulus to the spine, and which can be referr ed to as the abnormal mechanical causes of spinal pain and degenera tion either quantitatively or qualitatively. There are many biomechanical factors known to contribut e to spinal pain: vibr ation, lordosis, torsion, driving motor vehicles, material handling, leg le ngth discrepancy, etc. However, there are still controversies on the roles of mechanical factors relative to spinal pain. Vibration Epidemiologic studies reported increased spinal pain and/or disc disease in those who drive more than 3 hours per day or operate vibrating equipment (Frymoyer et al., 1983). Vibration, particularly in the fr equency domain of 5 to 15 Hz in which resonance of the spine can occur (Goel et al., 1994 ), is considered a key etiologic factor in low back pain (Magnusson et al., 1996), neurovestibular disorders (Sei del et al., 1988), and a causal factor in circulatory disorders such as Raynauds syndrome (Dandanell & Engs trom, 1986). Thus, the industries such as the transportation and construction, as well as the military are working toward minimizing occupational exposure to potenti ally noxious mechanical stim uli (Bongers et al., 1988; de Oliveira et al., 2001). Brumagne et al. (1999) perfor med a study to test the proprio ceptive changes in response to vibration in human. They applied a vibra tion (70 Hz, 0.5 mm amplitude) to the multifidus muscle for 5 seconds and measured the lum bosacral repositioning accuracy. They found a significant increase in directional error during vibration of the para spinal muscles. The subjects had the illusion during vibration that their pelvis was more pos teriorly tilted, and accordingly, they undershot the target position. They explained their results with the reflex inhibition of the muscle spindle, which play an important role in proprioception. They conc luded that vibration on the spine and trunk muscle can re sult in damage to or dysfunc tion of the muscle spindles.

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58 Consequently, a decreased muscle spindle input could jeopardize spin al proprioception and segmental stability, and likely to make the spine more vulnerable to injuries and low back pain. On the other hand, whole-body vibration has been utilized as an exercise therapy for musculoskeletal problems in sports, geriatrics, an d rehabilitation laboratories (Bosco et al., 1999; Rittweger et al., 2000). Vibration is thought to elicit muscle activity via stretch reflexes (Clark et al., 1981). Rittweger et al. (2002) observed that metabolic power increased during the wholebody vibration from a ground platform (amplitude of 2.4 mm, 5.0 mm, 7.5 mm, frequency of 18 Hz, 26 Hz, 34 Hz, duration of 4 minutes) and th at this metabolic power is augmented by the application of additional axial lo ads. Ritweger et al. (2002) pe rformed another study on patients of chronic lower back pain devoid of specific spin e diseases to verify the therapeutic effects of vibration exercise (maximum amplitude of 6 mm, frequency of 18 Hz, duration of 4 to 7 minutes). They compared whole-body vibration ex ercise with lumbar extension exercise, and observed a significant reduction in pain sensa tion and pain-related disability in both groups. Lumbar extension torque increased in the vibrat ion exercise group, but significantly more in the lumbar extension exercise group. They concluded th at a well-controlled vibr ation can be the cure rather than the cause of lower back pain. Some animal studies indicated that brief (<20 min) daily durations of extremely low-level (0.5g), high-frequency (15-90 Hz) vibration can be strongly an abolic to the trabecular bone, increasing bone mineral density, trabecular wi dth and number in the weight-bearing skeleton (Rubin et al., 2001; Rubin et al., 2002). These st udies suggested the osteogenic potential of extremely low-level mechanical stimuli as a treatment for osteoporosis. Rubin et al. (2003) studied transmissibility of high-frequency ( 15-35 Hz) ground-based, whol e-body vibration to the proximal femur and lumbar vertebrae. They obser ved 30-130% of transmissibility of the loading

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59 vibration, regardless of the target region, freque ncies, and posture of th e subjects. In addition, transmissibility at the hip was different to the sp ine, mostly at the lowest frequencies in this study. For the loading frequency less than 20 Hz, the resonance at the hi p exceeded 100% in the erect standing and relaxed standing postures. Ho wever, the resonance at the lumbar vertebrae was lower than that at the hip, a nd it was constantly maintained at the high frequency of loading vibrations and the other sta nding postures (relaxed standing, be nd knee). They mentioned that slight changes in posture can ha ve significant influence on the degr ee of vibration to be delivered to the spine or hip regions. They also emphasi zed the possible undesirable side effects of using whole-body vibration, against the prevention strategy for the oste oporosis. They suggested that vibration which approaches 1 g (9.8 ms-2), even at beneficial high fr equencies, should be avoided considering the risks to many physiologic systems. Lordosis Reduced or flattened lumbar lo rdosis has signified lumbar spine problems in previous studies (Adams & Hutton, 1985; Evcik & Yucel, 2003; Farfan et al., 1972). On the other hand, the cultural groups who spend considerable time in lumbar flexed position are known to suffer less low back pain and disc degenerati on (Fahrni & Trueman, 1965). Although there are controversies raised in previous studies, the biomechanics of lumbar lordosis and back pain were found to be closely related in some instances. Frymoyer et al. (1984) performed a radiologic study to investigate the lumbar lordosis in low back pain patients. They found no association of lumbar lordosis with low back pain. Murrie et al. (2003) conducted MRI assessments of lumbar lordosis in between patients of low back pain and normal controls, and did not find any significant difference in both gr oups. Instead, they observed that lumbar lordosis is more prominent in women and those with a higher body mass index.

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60 However, there are some cogent observations on the significant changes in the facet joints and discs associated with spine extension and hy perlordosis. Dunlop et al (1984) reported that each degree of increased extension of the spine leads to a 4% increa se in peak articular pressure in the facet joints. Yang and King (1984) reported that arthritic fa cet joints may bear up to 47% of the load transmitted. Thus, these increased loads to the facet joints lead to abnormal motion to the inferior facet articulation and damage to the joint structures, which can cause low back pain. The available evidence does not support strong co nclusions, but there seem to be disadvantages to hyperextension of the lumbar spine if there is facet joint arthriti s or disc pathology. Roussouly et al. (2005) perfor med a radiologic study to classi fy the normal variation of sagittal lumbar spine and pelv is. They found that sagittal alig nment of the human spine and pelvis was highly variable in different individuals in a standing position. The angle of the superior end plate of S1 with respect to the horizontal axis varied between 20 and 65 the angle of global lumbar lordosis varied between 41 and 82 and the number of vertebral bodies in a lordotic orientation varied from 1 to 8. The char acteristics of the lumbar lordosis were most dependent on the orientation of the sacral slope and the pelvis. The upper ar c of lumbar lordosis remained relatively constant, with an average value of approximately 20 In contrast, the lower arc of lordosis was the most important determinan t of the global lordosis They concluded that changes in the specific sagittal alignment and lumbar lordosis are potentially responsible to indicate degenerative changes and symptomatic back pain, instead of utilizing the vague term of lordosis. Torsion Axial rotation of the spine i nvolves torsion of the interver tebral discs and impaction of the facet joints. It is considered to be a possible mech anism of spinal damage and pain, especially

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61 to the lumbar spine (Hadjipavlou et al., 1999). The assumed injury mechanism is that shear loading of the annulus and the damage of the facet joints and ligament structures may lead to the segmental damage. However, the axial torque was not considered a significant factor to contribute to the disc degeneration in previous st udies. Because the facet joints limit the torsion to small range (1-2 ), it does not appear to allow enough st ress to lead to the disc damage (Adams & Hutton, 1981). Likewise, in an intact disc the facet joints and posterior ligaments are known to protect the intervertebral disc from tors ional loading. Because the axis of rotation of a lumbar vertebra passes through the posterior part of the vertebra l body, all the posterior elements of the moving vertebra swing around the axis during the axial rotation (Cossette et al., 1971). Quantitative analysis revealed that the disc contri butes 35% of the resistance to torsion, and the remaining 65% are being exerted by the posteri or elements (Farfan et al., 1970). Another experimental study showed that the facet joints c ontribute 42 to 54% of th e torsional stiffness in an FSU (Asano et al., 1992). In contrast, a study by Liu et al. (1985) supported that cyclic torsional loads can lead to the failures in the disc, facet, laminae, and cap sular ligaments. The anterior and posterior components are probably damaged or irritated by the axial torque of the lumbar spine. The initial torsional loading on the vertical axis are borne by the posterior elements (White & Panjabi, 1990). Biomechanical Performance of Painful Spine Patients with low back pain demonstrate a change in the mobility of the spine (Pearcy et al., 1985), deficits in reaction tim e, coordination (Luoto et al., 1996), and postural control with reduced velocity (Marras & Wongsam, 1986) when compared with healthy subjects. Differentiating the mechanical performance of pathologic spine with the normal variation

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62 associated with gender, aging, and physical f unction will be the basic step to determine the treatment methods for patients with spinal pain. However, previous studies have been limited to identifying the behaviors of the spinal kinema tics of low back pain patients, instead of approaching the various biomechani cal performance of the spine. Vogt et al. (2001) performed a kinematic study of lumbar spine during the treadmill walking in patients of chronic low back pain. They found significantly sh orter stride times and stride-to-stride variability in al l anatomic planes in these patients, as compared with healthy subjects. Decreased stride time, suggesting smaller steps of the patients, c ould be interpreted as a rigid or more cautious walking pattern and a pr otective way to reduce or avoid the pain. The increased between-subject variabilit y could be interpreted as patie nts individual adaptations and adjustments in walking behavior. These findings were interpreted as changed neuromuscular strategies to maintain an effective manner of locomotion which could be mediated by the altered proprioception in the lower back region. Neverthe less, the overall pattern of angular spinal displacements in patient group was shown to be w ithin normal limits. It was suggested that pain of musculoskeletal origin had no significant effect on the magnitude of lumbar angular displacements. Shum et al. (2005) performed a kinematic st udy of the lumbar spine and hip during sitting and standing movements in patients of low back pain. They found that th e mobility and velocity of the spine and hip were significantly limited and the contribution of the lumbar spine relative to that of the hip was reduced in back pain subjects. The patients with low b ack pain, in particular those with positive straight le g raising (SLR) sign, had altered hip-spine coordination. Shum et al. (2006) performed another spine kinematic st udy for a picking up activity, and found the same reduced mobility of lumbar spine an d hip in low back pain subjects.

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63 Al-Eisa et al. (2006) completed a kinematic st udy of the trunk in pa tients of unilateral nonspecific low back pain during sitting. They found a significant corre lation between pelvic asymmetry and asymmetric trunk motion in the pa tient group, and suggested that people with low back pain may have a distinct compensatory mechanism, secondary to the pelvic asymmetry from the unilateral low back pain, which put the lumbar spine under higher stresses. They concluded that movement asymmetry, rather than ROM, may be a better indicator of disturbed function for people with low back pain. McGill et al. (1999) studied the motions of th e spine and trunk muscle activity in normal elderly subjects (without back pain) during trunk posturing move ments. They found the elderly exhibited slower motion, reduced ROM in full fl exion and lateral bending. Furthermore, there was more coupled motion in the twisting effort s and abdominal muscles appeared to become more active earlier in the lateral bending move ment. The earlier activa tion and increased cocontraction suggested that elderl y people might be seeking greater stabilization either for general balance or for actual spine stabilization. Landing Biomechanics Previous studies on landing mechanics were mostly descriptive in biomechanical properties of various landing conditions. In majo rity of previous studies focused on defining injury mechanism (Fagenbaum & Darling, 2003; Wikstrom et al., 2006), gender differences (Kernozek et al., 2005), and condi tional variability (Schot et al., 2002) of landing mechanics, associated with knee and ankle joints. There were several different techniques and models developed for the landing biomechanics of lowe r extremity joints (Baca, 1999; Nagano et al., 1998; Spagele et al., 1999). Additionally, there were several landing techniques tested to compare the efficiencies and injury rates duri ng landing activities (McNai r et al., 2000; Onate et al., 2005). Because landing from a variety of jumping is a good example of active movements,

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64 the ability of a subject to cont rol kinetic and kinematic changes applied to his/her body segments from landing can be a good measure to determine the functional status. However, most studies focused on the performance of lower extremity joints, and to our best knowledge none were developed to evaluate the pe rformance of upper body segments above the hip joints during landing so far. Biomechanical Performance of Low er Extremity Joints during Landing Landing involved in many activit ies are vigorous and violen t in nature. There are many reports about landing injuries from various acti vities (Ford et al., 2003; Salci et al., 2004). Many studies have been conducted to define the biom echanical etiology of lo wer extremity injuries during landing procedures (James et al., 2003; McNitt-Gray, 1993; Sp agele et al., 1999). Vertical ground reaction force (GRF) from the ground imp act of landing ranged from 1.0 to 14.0 times of body weight for normal subjects (Caulfield & Garrett, 2004; Decker et al., 2002; Fritz & Peikenkamp, 2001; James et al., 2003). Landing cond itions investigated include landing height (McNitt-Gray, 1993; Zhang et al., 2000), landing posture and techni que (Eloranta, 1996; Kovacs et al., 1999), joint stiffness of lower extremity joints (Devita & Skelly, 1992; Horita et al., 2002), and the number of legs involved in landing (C aulfield & Garrett, 2004; Hass et al., 2003). Devita and Skelly (1992) studied biomechani cal variables applied to the lower extremity joints from drop landing with a fall-height of 59 cm, comparing soft a nd stiff landing conditions. Soft and stiff landings were defi ned with knee flexion angle afte r ground impact as greater and less than 90 The shapes of GRF, moment, and power curves were identical between both landings. Larger GRF, hip extensor, knee fle xor, and ankle plantar flexor moments were observed during descent in the stiff landing, which produced a more erect body posture and a flexed knee position at impact. Also the stiff landing exhibited larg er power in ankle muscles,

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65 while the soft landing showed larg er powers in knee and hip muscle s. They concluded that ankle plantar flexors absorbed more energy in the stiff landing, wher eas the hip and knee extensors absorbed more energy in the soft landing. To compare gymnasts and recreational athl etes, McNitt-Gray (1993) evaluated lower extremity joints kinetics during drop landing with different landing heights. They found gymnasts chose to dissipate the impact loads by using the larger ankle and hip extensor moments at higher impact velocities than recreational athletes who chose to adjust their strategy by using greater degrees of hip flexion and longer landing phase durations than the gymnasts. The greater demands placed on the ankle and hip extensors by the gymnasts, as compar ed to the recreational athletes, was explained by the need to maintain balance during competitive gymnastics landings or, by the inability of recr eational athletes to produce larger ex tensor moments at the ankle or hip during landings from the great heights. Zhang et al. (2000) studied lower extremity jo ints kinetics with different landing heights (0.32 m, 0.62 m, and 1.03 m) and di fferent landing techniques (sof t, normal, and stiff landings). They found increases in peak GR F, peak joint moments, and pow ers with increases in landing height and stiffness. The mean eccentric work were 0.52, 0.74, and 0.87 J/kg by the ankle plantar flexors, 1.21, 1.63, and 2.26 J/kg by the knee exte nsors, and 0.94, 1.31, and 2.15 J/kg by the hip extensors, for heights of 0.32, 0.62, and 1.03 m, respectively. They concluded that knee extensors were consistent contributors to ener gy dissipation, while th e ankle plantar flexors contributed more in the stiff la ndings and the hip extensors did more in the soft landings. This shift from ankle to hip strategy was observed as landing height increased. They explained that larger volume of proximal muscles (knee and hip jo int muscles) of the lower extremity was more capable of energy absorption compared with the ankle muscle group. Another study suggested

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66 that biarticular muscles are used effectively for power transportation during locomotion (Bobbert & van Ingen Schenau, 1988). Gender Difference The majority of knee injuries occurred duri ng landing from various activities are caused by the non-contact mechanism. Numerous studies have found females have a higher rate of noncontact anterior cruciate ligament (ACL) injuri es compared to males (Decker et al., 2003; Kernozek et al., 2005). Many studi es had been conducted to identify different mechanical properties of landing in each ge nder and etiology of th is gender disparity. Anatomically or intrinsically, small cross-sectional area of ACL (Feagin & Lambert, 1985), narrower intercondylar notch (Souryal et al., 1988), grea ter quadriceps angle, and greater knee laxity (Malinzak et al., 2001) of females have been sugg ested to contribute to the higher ACL injuries in females than in males. Except for the intrinsic anatomical factors, the differences in level of conditioning, level of muscle strength, and motor c ontrol strategies in fema les were suggested as the extrinsic factors, related with gend er disparity (Delfico & Garrett, 1998). Chappell et al. (2002) compared knee kinetics of male and female recreational athletes performing forward, vertical, and backward stopjump tasks. They observed females exhibited greater proximal anterior shear force, greater knee extension and valgus moments than males did during the landing phase. During the takeoff phase, males showed greater proximal tibia anterior shear force than females. They concluded th at females might have altered motor control strategies that result in knee positions in which ACL injuries might occur. Decker et al. (2003) compared the biomechan ical variables of lower extremity joints between male and female subjects performing a 60 cm drop landing. They found females demonstrated a more erect landing posture and utilized greater hip a nd ankle joint range of motions and maximum joint angular velocities when compared to males. Females exhibited

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67 greater energy absorption and peak powers from the knee extensors and ankle plantar-flexors compared to males. Energy absorption contribut ions revealed that the knee extensor was the primary shock absorber for both genders. The ankle plantar flexor was the second largest contributor to energy absorption for the females and the hip extensor was for the males. The different shock absorption strategy used in fema les was proposed to prov ide a greater potential risk for non-contact ACL injury for fe males under certain landing conditions. Kernozek et al. (2005) compar ed gender differences in frontal and sagittal plane biomechanics during a 60-cm drop landing. They obs erved that females exhibited greater peak hip and knee flexion, and ankle dorsi flexion angles in the sagittal plane, and greater peak knee valgus and ankle pronation angles in the frontal plane. Females e xhibited greater peak vertical and posterior GRF, and reduced varus moment th an males. They noted the importance of gender differences in the frontal plane variables in addition to those in sagittal plane. Landing Stiffness During landing from various activ ities, the actions of different musculoskeletal structures, including muscles, tendons, and ligaments, are inte grated together so that the overall skeletons behave like a spring. As a result landing from a jump can be modeled by using a simple springmass system (Asmussen & Bonde-Petersen, 1974). Stiffness of the leg spring represents the overall stiffness of the integrated musculoske letal system during the landing phase, which is referred to as leg stiffness. Leg stiffness infl uences the kinetics and kinematics applied to the whole body during landing. Leg stiffness is greatly determined by the knee joint stiffness, which is mostly affected by knee extensor muscles. Join t stiffness is calculated by a linear regression of the knee joint moment/angle relationship. The in teraction between leg stiffness and reflex activities plays a major role in regulating musc le power and performance in preand post-

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68 landing to absorb or dissipate the large energy de veloped from the ground impact (Horita et al., 1996). Horita et al. (2002) studied the interaction between pre-la nding activities and stiffness regulation of the knee joint during a drop landing followed by a countermovement jump by examining landing motions, GRF, and EMG activity of the vastus lateralis during the preand post-landing. They divided the contact phase into three phases from initial contact to takeoff of the countermovement jump: (1) init ial impact to initial peak of knee joint moment, (2) initial peak to onset of pushoff, and (3) concentric phase until takeoff. Dr op landing performance was evaluated using the takeoff velocity, average contact time, and knee joint moment. A positive correlation was found between positive peak powe r of knee joint and the knee joint moment. However, they did not find any significant rela tionships between any drop landing performance parameters and ankle measures. They explaine d leg stiffness with a combination of precontraction of the vastus lateralis muscle and knee joint a ngular velocity at touchdown. They proposed two types of landing motions with regard to the pre-landing motion of the knee joint. The proper pre-landing movement could be characterized by th e knee flexion just before touchdown, which is associated with a high initial joint stiffness coupled with the high joint power. This was called bouncing type in their st udy, which is close to the plyometrics. On the other hand, an inadequate pre-landing movement associated with incomplete knee flexion induced subsequent deep-knee flexion after t ouchdown was called the absorbing type of landing and was regarded as the poor type. The absorb ing movement comes too late and demand longer contact time and lower takeoff velocity. The pre-activation of landing is initia ted by the centrally pre-programmed motor commands of the required landing ta sk (Dyhre-Poulsen et al., 1991), and increase the sensitivity

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69 of the muscle spindle to enhan ce the stretch reflex (Gottlieb et al., 1981). The stretch reflex enhances muscle stiffness, and thus, improved s tiffness regulation could be attained by proper pre-landing muscle activati on (Allum & Mauritz, 1984). GRF acting on the body during landing has been implicated in injuries to the lower extremity, and controlling peak GRF at impact is directly associated with lowering the landing stiffness. Apparently, the movement of lower extr emity joints can influence the magnitude of the impact forces. It is generally known that subjects who land on the balls of th eir feet and flex their knee and ankle joints more have lower peak GRF. It was suggest ed that more joint movements allowed the body mass to decelerate over a longer period of time thus the impact force and time to peak force were decreased. McNair et al. (2000) reported that GRF coul d be decreased immediately after instruction of landing technique about the lower extremity jo ints kinematics. Additionally, they commented that instruction could be more effective if the subjects attention was drawn to distinct cues (sound of soft landing impact) in the environment. Cowling et al. (2003) assessed th e efficacy of verbal instruct ion about landing techniques to change impact force. They used instructions to increase knee flexi on, to recruit hamstring muscles earlier, and muscle bursts immediately befo re landing. Only the instruction to increase knee flexion resulted in signifi cantly greater knee flexion at initial ground contact and lower GRF, compared with the ot her instruction conditions. Park et al. (2006a) compared mechanical charac teristics of soft land ing between male and female subjects performing vertical leaps and drop landings. They found significant increases in ankle plantar flexion at touchdown, knee flexio n motion during the landing phase, times for maximal flexion of all joints, and decreases in pe ak vertical GRF, axial hip joint force, and knee

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70 extensor moments in soft landing. Peak vertical GRF of males was significantly greater than that of females in soft landing. Pre-land ing extension of the distal joints in vertical le ap and that of proximal and distal joints in drop landing were suggested to activate th e soft landing, with precontracting the muscles to prepar e the proper landing. They noted that the soft landing strategy for males was fit for plyometrics and that of females was for absorbing type of landing. Performance of Adapted Landing Bi omechanics to Various Conditions Landing is ideally suited for a performance study of weight-bearing segments of body, as it requires large eccentric muscle forces during the control of joint flexion and mimics the muscular stresses experienced during the landing phase. Theref ore, numerous landing studies have been conducted to test the performance of normal subjects in various conditions, as well as to test subjects of ACL defici ency, ankle instability, fatigue of knee joint muscles, and so on. A lesion of ACL is a major trauma of the knee joint, and mostly treated with a ligament reconstruction. ACL reconstruction demands sophisticated rehabil itation program to regain the original function before the injury. Decker et al. (2002) studied the biomechanics of lower extremity joints in fully rehabilitated ACL r econstructed and healthy subjects performing drop landing from a 60-cm height. At initial touchdown, the ACL gr oup demonstrated greater hip extension and ankle plantar flex ion, compared to the healthy gr oup. The peak vertical GRF was not different between groups, but the ACL group delayed the time to its occurrence. The knee extensors provided the major energy absorpti on function for both groups; however, the ACL group performed 37% more ankle plantar flexor wo rk and 39% less hip extensor work compared with the healthy group. They conc luded that the ACL group utilized a different landing strategy adapted to the ACL reconstruction which employe d the hip extensor muscles less and the ankle plantar flexors more.

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71 Doorenbosch et al. (2003) studied EMG activ ity of the quadriceps and hamstrings in patients with ACL deficiency and healthy subjec ts during the vertical jump and landing. They observed significantly higher co-contraction i ndex of quadriceps/hams tring muscles in ACL deficiency subjects. This was suggested that hi gher levels of co-contra ction of quadriceps and hamstrings during movements in ACL deficient subj ect help to compensate for the loss of the passive constraining structure. Caulfield et al. (2004) studied jump and landing performan ce in subjects with functional instability of the ankle joint a nd normal control. The subjects performed five single leg jumps onto a forceplate. They observed la teral and anterior force peaks o ccurred significantly earlier in subjects with ankle instability. These changes occurred immediately post-impact and too early for the reflex correction. Madigan et al. (2003) studied th e effect of lower extremity fatigue on the performance of lower extremity joints during drop landing. EMG da ta were used to confirm fatigue in the quadriceps muscles. They observed a decrease in peak vertical GRF, an increase in maximum flexion of knee and ankle joints occurred early in a fatigue landing, while significant changes in vertical GRF impulse and time to maximum knee flexion occurred during the middle or late stages of a fatigue landing. For the first half of a fatigue landing, hip extensor impulse increased, knee extensor impulse did not change, and a nkle plantar flexor impulse decreased. These changes were explained with a distal-to-proximal redistribution of extensor moments, which allowed the larger proximal mu scles to contribute more to re sisting collapse during landing. They also suggested active insufficiency of gast rocnemius, since it crosses the knee and shortens as knee flexion increases. The shortening of this muscle diminishes the ability to produce plantar flexor moment at the ankle. The increase in knee flexion upon landing during the first half of the

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72 fatigue landing may have reduced the contribution of the gastrocnemius muscle to plantar flexor impulse and resulted in an overall decrease in plantar flexor impulse. This decrease also has dictated an increase in hip ex tensor impulse to generate th e necessary support moment for landing deceleration. Park et al. (2006b) studied the effects of limited lower back motion on soft landing mechanics of lower extremity joints with subjec ts wearing various low back braces, simulating different low back stiffness conditions. They found that knee and hip joint flexions were decreased and peak vertical GRF and axial hi p force were increased in stiff brace condition, compared with no brace and soft brace conditions Additionally, typical sequential joint flexion from distal to proximal was disrupted in fema les wearing the stiff brace, comparing to male counterpart. They concluded that limited spinal motions by the brace caused alterations in knee and hip joint motions during the landing phase and an increase in impact force. They emphasized that lower back motion is one of the factors in determining landing mechan ics, and a stiff lower back is associated with a s tiff landing. With limited trunk mo tion, more decelerating torque might be concentrated on the knee extensors for females, and more axial loads be transmitted to the proximal body segments of males during the soft landing.

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73 Table 3-1. Average neutral zones for a functional spinal units in different regions of the spine ( ). Region Flexion/Extension (total) Lateral bending (one side) Axial rotation (one side) C0 C1 1.1 1.6 1.5 C1 C2 3.2 1.2 29.6 C3 C6 4.9 4 3.8 C7 T1 / T11 T1 1.5 2.2 1.2 L1 L2 / L3 L4 1.5 1.6 0.7 L5 S1 3 1.8 0.4 Note: Cited from the work of White and Panjabi (1990). Table 3-2. Representative ranges of motion of C0-C1-C2 complex ( ). Level Reference Flexion/Extension (total) Lateral bending (one side) Axial rotation (one side) C0-C1 Penning (1978) 35 10 0 Goel et al. (1988) 23 3.4 2.4 Panjabi et al. (1988) 24.5 5.5 7.2 C1-C2 Penning (1978) 30 10 70 Goel et al. (1988) 10.1 42 23.3 Panjabi et al. (1988) 22.4 6.7 38.9 Table 3-3. Representative ranges and limits of motion of the middle and lower cervical spines ( ). Region Flexion/Extention Lateral bending (one side) Axial rotation (one side) Middle C2-C3 10 (5 16) 10 (11 20) 3 (0 10) C3-C4 15 (7 26) 11 (9 15) 7 (3 10) C4-C5 20 (13 29) 11 (0 16) 7 (1 12) Lower C5-C6 20 (13 29) 8 (0 16) 7 (2 12) C6-C7 17 (6 26) 7 (0 17) 6 (2 10) C7-T1 9 (4 7) 4 (0 17) 2 (0 7) Note: Cited from the work of White and Panjabi (1990). Table 3-4. Normal active cervical ranges of motion ( in vivo ) reported in the literatures ( ). Motion Dvorak (1992) Lanz (1999) AMA (2001) Flexion/Extension 141.3 116.1 110 Lateral bending 91.4 84.1 90 Axial rotation 175 144.2 160 Rotation from flexion 81.4 Rotation from extension 165

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74 Table 3-5. Representative ranges and lim its of motion of the thoracic spine ( ). Level Flexion/Extension (total) Lateral bending (one side) Axial rotation (one side) T1 T2 4 (3 5) 5 (5) 9 (14) T2 T3 4 (3 5) 6 (5 7) 8 (4 12) T3 T4 4 (2 5) 5 (3 7) 8 (5 11) T4 T5 4 (2 5) 6 (5 6) 8 (5 11) T5 T6 4 (3 5) 6 (5 6) 8 (5 11) T6 T7 5 (2 7) 6 (6) 7 (4 11) T7 T8 6 (3 8) 6 (3 8) 7 (4 11) T8 T9 6 (3 8) 6 (4 7) 6 (6 7) T9 T10 6 (3 8) 6 (4 7) 4 (3 5) T10 T11 9 (4 14) 7 (3 10) 2 (2 3) T11 T12 12 (6 20) 9 (4 13) 2 (2 3) T12 L1 12 (6 20) 8 (5 10) 2 (2 3) Note: Cited from the work of White and Panjabi (1990). Table 3-6. Representative ranges and lim its of motion of the lumbar spine ( ). Level Flexion/Extension Late ral bending Axial rotation L1 L2 12 (5 16) 6 (3 8) 2 (1 3) L2 L3 14 (8 18) 6 (3 10) 2 (1 3) L3 L4 15 (6 17) 8 (4 12) 2 (1 3) L4 L5 16 (9 21) 6 (3 9) 2 (1 3) L5 S1 17 (10 24) 3 (2 6) 1 (0 2) Note: Cited from the work of White and Panjabi (1990). Table 3-7. Comparison of lumbar compression loads in various tr unk postures without external loading. Nachemson (1966; 1981) Takahashi et al. (2006) Ledet et al. (2005) Trunk posture N % of standing N % of standing body weight % of standing Supine 300 25 1.9 95 Standing 700 100 645 100 2 100 Sitting 1000 140 2.5 125 Standing flexed 2.6* 130* 10 1277 198 20 1200 150 1922 298 30 2305 357 Sitting flexed 185 2.8 140 The trunk flexion angles were not specified in Ledet et al. (2005).

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75 Table 3-8. Average ranges of motion of the lu mbar spine in normal walking and running in different studies ( ). Source Flexion/Extension Late ral bending Axial rotation Walking Crosbie et al. (1997) 3.5 9 4.5 (Lower thoracic) 2.5 7.0 4.0 (Pelvis) 3.5 6.0 4.0 Callaghan et al. (1999) 6.2 6.7 7.1 Van Herp et al. (2000) 2.3 4 6.6 Running Schache et al. (2002) 13.3 18.5 23.0 (Pelvis) 7.6 10.6 13.9 Table 3-9. Peak compression loads to th e lower lumbar level during walking ( BW). Source Model Compression force Cappozzo (1983) Single muscle equivalent model 1.0 2.5 Cromwell et al. (1989) EMG model 1.0 Khoo et al. (1995) Single muscle equivalent model 1.5 2.1 Callaghan et al. (1999) EMG model 0.9 3.5 Inverse dynamic model 0.2 1.0 Nachemson (1964) Direct intradiscal pressure measurement 850 (N) Table 3-10. Intradiscal pressure of low lumbar level during various activities. Wilke et al. (1999) Rohlmann et al. (2001) Nachemson (1966; 1981; 1987) Position Intradiscal pressure (MPa) Percentage of standing (%) Percentage of standing (%) Lying supine 0.1 20 25 Relaxed standing 0.5 100 100 Standing with forward bending 1.1 216 150 Sitting without backrest 0.46 90 140 Sitting with maximum flexion 0.83 185 Standing up from a chair 1.1 Walking 0.53 0.65 130 121 Jogging 0.35 0.65 170 Jumping 240 380 157 Lifting 20 kg, squat lift 1.7 300 Lifting 20 kg, stoop lift 2.3 460 486

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76 Figure 3-1. The load-displacement curve of a func tional spinal unit (FSU) is generally nonlinear and biphasic [neutral zone (NZ) and elasti c zone (EZ)]. Range of motion (ROM) is the sum of the neutral and elastic zones. Average flexibility coefficient (FC) is the elastic zone divided by the maximum physiological load.

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77 CHAPTER 4 RESULTS Effects of Landing Technique Overall landing characteristics were evaluate d using different landing variables (Table 41). Repeated measures MANOVA revealed signi ficant main effects of gender (p<0.001) and landing technique (p<0.001) for th e landing variables. However, no significant interaction was found between gender and landing technique (p=0.057) (Table C-1 in Appendix C). Follow-up univariate contrast procedures revealed that t(LP), P(KFA), and P(HFA) increased significantly and PVGRF decreased significantly when going from NL to SL condition. For touchdown angles, both TD(KFA) and TD(HFA) were significantly more flexed when going from NL to SL condition. Female subject s exhibited significantly greater P(HFA) than male subjects across both landing conditions (Table 4-1). Kinematic characteristics of the spinal co lumn after touchdown of drop landing were evaluated with touchdown angle and extension RO M variables of each spinal region. Relative to research questions Q1 and Q2 and associated hypotheses (1a and 2a), repeated measures MANOVA revealed a significant main effect of gender (p=0.011) and a significant interaction between gender and landing technique (p=0.025) for touchdown angle vari ables, and significant main effects of gender (p=0.013) and landing technique (p<0.001) for extens ion ROM variables. However, no significant main effect of la nding technique was found for touchdown angle variables (p=0.236), and no signi ficant interaction was found between gender and landing technique for extension ROM va riables (p=0.232) (Table C-1). Follow-up univariate contrast procedures for touchdown angl e variables revealed that TD(TL/L) of females was significantly greater than that of males (i.e., females demonstrated significantly more extended thoracolumbar regi onal angle at touchdown than males), and a

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78 significant interaction was found in TD(TH/TL) (Table 4-2). The signi ficant interaction of TD(TH/TL) between gender and landing technique means that males demonstrated more flexed thoracic regional angle at touchdown from NL to SL condition, while females did not demonstrate differences across different la nding techniques (Figure 4-1). Follow-up univariate tests for extension ROM variables identified a significant main effect of gender for tP(TH/TL) and significant main effect s of landing technique for P(TH/TL) and P(LC/TH). The contrast procedures indicate d that females exhibited signif icantly greater extension motion in the thoracic region than males did, and extens ion motions in the thoracic and lower cervical regions increased significantly from NL to SL condition (Table 4-2). The overall kinematic characteristics of each spinal region reveal that the lumbar and thoracolumbar regions exhibit flexion and the thoracic and lower cervical regions show extension during the landing phase However, the thoracolumbar region undergoes a short period of extension followed by flexion during the landi ng phase in selected subjects (Figure 4-2). Kinetic characteristics of L/S and C/T junctio ns after touchdown dur ing drop landing were evaluated with the kinetic variables at L/S and C/T junctions. Relative to research questions Q1 and Q2 and associated hypotheses (1b and 2b), repeated measures MANOVA revealed a significant main effect of landi ng technique for the kinetic vari ables at L/S junction (p<0.001), and significant main effects of gender (p=0.031 ) and landing technique (p<0.001) for the kinetic variables at C/T junction. There was also a significant interaction be tween gender and landing technique for the kinetic vari ables at C/T junction (p=0.018). However, no significant gender effect (p=0.403) or interac tion (p=0.196) between gender an d landing technique was found for the kinetic variables at L/S junction (Table C-1). The univariate contrast procedures revealed that

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79 all the kinetic variables at L/S junction decrease d significantly from NL to SL condition (Table 4-3). For the kinetic variables at C/T junction, all but ShF(C/T)ant decreased significantly from NL to SL condition and ShF(C/T)ant and ShF(C/T)post demonstrated significant interactions between gender and landing technique (Table 43). The significant inte ractions between gender and landing technique for ShF(C/T)ant and ShF(C/T)post indicate that ShF(C/T)post of males was greater than that of females during NL, but no gender difference was observed during SL condition. On the other hand, ShF(C/T)ant of females was greater than that of males during NL, but not different from each other during SL condition. Females demonstrated decreased ShF(C/T)ant from NL to SL condition, but not for males (Figure 4-1). Effects of Knee Joint Muscles Fatigue Overall landing characteristics were evaluate d using different landing variables. Repeated measures MANOVA identified a significant main effect of gender (p=0.012) for the landing variables. However, no significant fatigue eff ect (p=0.559) or interaction (p=0.104) between gender and fatigue level was found for the landing variables (Table C-9). Follow-up univariate tests revealed that female subjects exhibited a significantly greater TD(KFA) than male subjects across both landing conditions (i.e., females demons trated more extended posture of knee joint at touchdown than males) (Table 4-4). Kinematic characteristics of the spinal co lumn after touchdown of drop landing were evaluated with touchdown angle and extension ROM variables. Relative to research questions Q3 and Q4 and associated hypotheses (3a and 4a), repeated measures MANOVA revealed significant main effects of gender for the touchdown angle (p=0.036) and extension ROM

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80 (p=0.006) variables. However, there was not any si gnificant fatigue effect or interaction for both touchdown angle and extension ROM variables (Table C-9). Univariate contrast procedures revealed that females demonstrated a significantly extended posture of thoracolumbar region at touchdown (greater TD(TL/L) in females than in males) and a greater extension motion (greater P(TH/TL) in females than in males) during landing phase than males did (Table 4-5). The overall kinematic characteristics of each spinal region reveal that the lumbar region exhibits flexion and thoracic a nd lower cervical regions show extension during the landing phase. However, the thoracolumbar region undergoe s a short period of extension followed with flexion during the landing phase in selected subjects (Figure 4-3). Kinetic characteristics of L/S and C/T junctio ns after touchdown of drop landing were evaluated with the kinetic variables of L/S and C/T junctions. Relative to research questions Q3 and Q4 and associated hypotheses (3b a nd 4b), repeated measures MANOVA revealed significant interactions between ge nder and fatigue level for the ki netic variables at L/S junction (p=0.033) and C/T junction (p=0.043) (Table C-9). Follow-up univariate tests identified significant interactions of gender fatigue level for AxF(L/S), ShF(L/S)post, ExtM(L/S), and AxF(C/T) (Table 4-6). The significant interaction for AxF(L/S) indicated that females exhibited incr eased AxF(L/S) from SL1 to SL2 condition, while males did not (Figure 4-4). The si gnificant interaction of ShF(L/S)post revealed that females exhibited increased ShF(L/S)post from SL1 to SL2 condition, while males did not. Also, ShF(L/S)post was greater in males than in females fo r SL1, but was greater in females than in males for SL2 condition. The significant interaction of AxF(C/T) indicated that females exhibited increased AxF(C/T) from SL1 to SL 2 condition, while males di d not. Also, AxF(C/T)

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81 was greater in males than in females for SL1, but was greater in females than in males for SL2 condition. The significant interaction of ExtM(L/S ) revealed that female s exhibited increased ExtM(L/S) from SL1 to SL2 condition, while male s did not. Lastly, ExtM(L/S) was greater in males than in females for SL1, but was greater in females than in males for SL2 condition.

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82 Table 4-1. Collapsed mean and SD values of di fferent landing variable s for different genders and landing techniques. NL SL Landing variables Mean (SD) Mean (SD) Gender: p Landing: p Landing Gender: p PVGRF (N kg-1) 23.91 (3.87) 15.77 (3.03) 0.877 <0.001* 0.723 Male 23.67 (3.22) 15.83 (2.91) Female 24.15 (4.55) 15.70 (3.27) t(LP) (s) 0.222 (0.084) 0.355 (0.139) 0.148 <0.001* 0.536 Male 0.203 (0.071) 0.320 (0.075) Female 0.241 (0.095) 0.390 (0.179) TD(KFA) ( ) -7.3 (6.8) -10.5 (6.4) 0.063 <0.001* 0.268 Male -9.2 (4.6) -13.2 (5.5) Female -5.3 (8.2) -7.8 (6.2) P(KFA) ( ) 53.0 (10.4) 69.9 (14.2) 0.087 <0.001* 0.131 Male 51.7 (12.6) 71.8 (16.0) Female 54.2 (7.9) 68.0 (12.6) TD(HFA) ( ) -61.5 (12.0) -66.6 (12.3) 0.853 <0.001* 0.747 Male -60.9 (10.7) -66.4 (11.7) Female -62.1 (13.6) -66.9 (13.4) P(HFA) ( ) 33.3 (13.7) 51.2 (10.8) 0.001* <0.001* 0.113 Male 24.8 (11.5) 47.0 (7.2) Female 41.7 (10.2) 55.5 (12.3) Note: NL (self-selected normal landing), SL (s oft landing), PVGRF (peak vertical GRF), t(LP) (time for landing phase), TD(KFA) (knee flexion angle at touchdown), P(KFA) (ROM of knee flexion angle from touchdown to initial peak of knee flexion), TD(HFA) (hip flexion angle at touchdown), P(KFA) (ROM of hip flexion angle from touchdown to initial peak of hip flexion), (significant in univariate tests)

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83 Table 4-2. Collapsed mean and SD values of touchdown angle and extension ROM of each spinal region for different genders and landing techniques. NL SL Kinematic variables Mean (SD) Mean (SD) Gender: p Landing: p Landing Gender: p Touchdown angle of each spinal region TD(L/S) ( ) 15.1 (8.8) 15.0 (9.0) 0.215 0.85 0.433 Male 13.3 (9.0) 12.6 (9.5) Female 17.0 (8.5) 17.4 (8.2) TD(TL/L) ( ) 11.9 (7.8) 11.9 (8.7) 0.003* 0.933 0.28 Male 7.6 (5.1) 7.2 (6.1) Female 16.1 (7.8) 16.7 (8.4) TD(TH/TL) ( ) -21.9 (6.9) -22.6 (7.8) 0.196 0.219 0.032* Male -23.1 (6.2) -25.1 (8.2) Female -20.7 (7.6) -20.1 (6.9) TD(LC/TH) ( ) -1.6 (12.4) -0.3 (13.4) 0.91 0.451 0.844 Male -1.4 (10.6) 0.1 (14.6) Female -1.7 (14.5) -0.8 (12.8) Extension ROM of each spinal region P(L/S) ( ) 0.5 (1.3) 0.0 (0.1) 0.167 0.061 0.179 Male 0.2 (0.3) 0.0 (0.1) Female 0.9 (1.7) 0.1 (0.1) P(TL/L) ( ) 2.3 (2.0) 2.0 (2.3) 0.070 0.392 0.526 Male 1.8 (1.9) 1.2 (1.7) Female 2.9 (2.0) 2.8 (2.6) P(TH/TL) ( ) 4.6 (3.7) 8.0 (5.1) 0.012* < 0.001* 0.057 Male 3.3 (2.8) 5.4 (3.9) Female 6.0 (4.2) 10.6 (4.9) P(LC/TH) ( ) 10.3 (7.8) 16.3 (10.4) 0.591 0.006* 0.557 Male 10.0 (9.7) 14.8 (10.8) Female 10.5 (5.7) 17.7 (10.1) Note: NL (self-selected normal landing), SL (soft landing), TD(L/S) (lumbar regional angle at touchdown), TD(TL/L) (thoracolumbar regional angle at touchdown), TD(TH/TL) (thoracic regional angle at touchdown), TD(LC/TH) (lower cervical regional angle at touchdown), P(L/S) (extension ROM of lumbar region from touchdown to initial peak during landing phase), P(TL/L) (extension ROM of thoracolumbar region from touchdown to initial peak during landing phase), P(TH/TL) (extension ROM of thoracic region from touchdown to initial peak during landing phase), P(LC/TH) (extension ROM of lower cervical region from touchdown to initial peak during landing phase) (significant in univariate tests)

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84 Table 4-3. Collapsed mean and SD values of kinetic variables at L/S and C/T junctions for different genders and landing techniques. NL SL Kinetic variables Mean (SD) Mean (SD) Gender: p Landing: p Landing Gender: p L/S junction AxF(L/S) (N kg-1) 8.52 (2.70) 5.77 (1.93) 0.182 <0.001* 0.268 Male 7.72 (2.43) 5.55 (1.77) Female 9.33 (2.81) 5.99 (2.13) ShF(L/S)ant (N kg-1) 1.73 (1.01) 0.83 (0.51) 0.085 <0.001* 0.022 Male 2.16 (1.24) 0.82 (0.55) Female 1.30 (0.45) 0.84 (0.48) ShF(L/S)post (N kg-1) 9.73 (3.54) 4.35 (2.32) 0.297 <0.001* 0.476 Male 10.48 (3.21) 4.66 (2.45) Female 8.97 (3.81) 4.05 (2.23) FlxM(L/S) (N m kg-1 BH-1) 1.37 (0.44) 0.91 (0.43) 0.583 <0.001* 0.961 M 1.41 (0.37) 0.95 (0.36) F 1.33 (0.50) 0.87 (0.51) ExtM(L/S) (N m kg-1 BH-1) 3.44 (1.15) 1.94 (0.70) 0.31 <0.001* 0.418 Male 3.68 (1.06) 2.02 (0.64) Female 3.19 (1.23) 1.85 (0.77) C/T junction AxF(C/T) (N kg-1) 4.98 (4.07) 1.65 (2.28) 0.749 0.001* 0.303 Male 4.36 (4.0) 1.96 (2.25) Female 5.60 (4.20) 1.34 (2.36) ShF(C/T)ant (N kg-1) 3.22 (2.28) 2.56 (1.73) 0.197 0.117 0.042* Male 2.35 (1.84) 2.56 (1.58) Female 4.08 (2.42) 2.56 (1.94) ShF(C/T)post (N kg-1) 5.54 (1.28) 4.26 (0.81) 0.026* <0.001* 0.001* Male 6.26 (1.42) 4.35 (0.97) Female 4.83 (0.53) 4.17 (0.65) FlxM(C/T) (N m kg-1 BH-1) 2.54 (0.91) 1.75 (0.81) 0.758 <0.001* 0.342 Male 2.40 (0.77) 1.80 (0.70) Female 2.68 (1.04) 1.70 (0.93) ExtM(C/T) (N m kg-1 BH-1) 3.85 (1.33) 2.24 (0.71) 0.302 <0.001* 0.145 Male 4.21 (1.4 2.25 (0.60) Female 3.50 (1.19) 2.24 (0.83) Note: NL (self-selected normal landing), SL (soft la nding), AxF (peak axial compressive force), ShFant(or post) (peak ant. or post. shear force), FlxM (peak flexor moment), ExtM (peak extensor moment), (L/S) (for lumbosacral junction), (C/T) (for cervicothoracic junction), (significant in univariate tests)

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85 Table 4-4. Collapsed mean and SD values of di fferent landing variable s for different genders and fatigue levels. SL1 SL2 Landing variables Mean (SD) Mean (SD) Gender: p Fatigue: p Fatigue Gender: p PVGRF (N kg-1) 15.72 (3.03) 16.47 (3.43) 0.312 0.087 0.007 Male 15.8 (3.04) 15.2 (3.36) Female 15.62 (3.19) 17.99 (2.98) t(LP) (s) 0.335 (0.147) 0.342 (0.163) 0.304 0.628 0.666 Male 0.307 (0.088) 0.308 (0.093) Female 0.368 (0.197) 0.382 (0.220) TD(KFA) ( ) -10.2 (6.7) -11.0 (7.1) 0.008* 0.322 0.281 Male -13.2 (5.8) -14.6 (6.5) Female -6.6 (6.2) -6.6 (5.2) P(KFA) ( ) 70.5 (15.0) 71.7 (15.9) 0.271 0.44 0.112 Male 72.8 (16.2) 75.9 (18.0) Female 67.6 (13.8) 66.5 (11.9) TD(HFA) ( ) -65.2 (10.9) -64.1 (15.5) 0.094 0.608 0.648 Male -66.3 (11.8) -64.2 (18.4) Female -64.0 (10.2) -63.9 (12.2) P(HFA) ( ) 52.0 (10.8) 53.6 (12.9) 0.094 0.264 0.749 Male 48.1 (6.3) 50.0 (8.5) Female 56.8 (13.3) 57.9 (16.3) Note: SL1 (soft landing before the fatigue procedur es), SL2 (soft landing after the fatigue procedures), PVGRF (peak vertical GRF), t(LP) (time for landing phase), TD(KFA) (knee flexion angle at touchdown), P(KFA) (ROM of knee flexion angle from touchdown to initial peak of knee flexion), TD(HFA) (hip flexion angle at touchdown), P(KFA) (ROM of hip flexion angle from touchdown to initial peak of hip flexion), (significant in univariate tests)

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86 Table 4-5. Collapsed mean and SD values of touchdown angle and extension ROM of each spinal region for different genders and fatigue levels. SL1 SL2 Kinematic variables Mean (SD) Mean (SD) Gender: p Fatigue: p Fatigue Gender: p Touchdown angle of each spinal region TD(L/S) ( ) 15.1 (9.0) 14.0 (9.3) 0.473 0.343 0.558 Male 13.5 (10.0) 13.1 (9.8) Female 17.0 (7.7) 15.1 (9.1) TD(TL/L) ( ) 11.5 (9.0) 11.4 (9.9) 0.005* 0.988 0.545 Male 6.8 (6.4) 6.5 (7.0) Female 17.0 (8.7) 17.3 (9.8) TD(TH/TL) ( ) -22.9 (8.1) -22.9 (7.8) 0.068 0.924 0.993 Male -25.7 (8.2) -25.7 (7.0) Female -19.6 (7.1) -19.6 (7.6) TD(LC/TH) ( ) 0.3 (13.6) 3.2 (13.4) 0.819 0.056 0.184 Male 0.7 (15.5) 1.6 (15.2) Female -0.1 (11.6) 5.0 (11.5) Extension ROM of each spinal region P(L/S) ( ) 0.0 (0.1) 0.4 (1.2) 0.074 0.072 0.087 Male 0.0 (0.0) 0.0 (0.0) Female 0.0 (0.1) 0.9 (1.7) P(TL/L) ( ) 2.1 (2.4) 1.5 (1.5) 0.085 0.083 0.212 Male 1.3 (1.8) 1.1 (1.2) Female 3.1 (2.7) 1.9 (1.6) P(TH/TL) ( ) 7.8 (5.5) 8.1 (5.5) 0.013* 0.624 0.624 Male 5.2 (4.0) 5.7 (4.4) Female 11.0 (5.5) 11.0 (5.5) P(LC/TH) ( ) 16.6 (10.6) 19.2 (12.3) 0.776 0.355 0.042 Male 15.0 (11.2) 22.0 (14.0) Female 18.6 (10.1) 15.9 (9.5) Note: SL1 (soft landing before the fatigue procedur es), SL2 (soft landing after the fatigue procedures), TD(L/S) (lumbar regional angle at touchdown), TD(TL/L) (thoracolumbar regional angle at touchdown), TD(TH/TL) (thoracic regional angle at touchdown), TD(LC/TH) (lower cervical regional angle at touchdown), P(L/S) (extension ROM of lumbar region from touc hdown to initial peak during landing phase), P(TL/L) (extension ROM of thoracolumbar region from t ouchdown to initial peak during landing phase), P(TH/TL) (extension ROM of thoracic region from touc hdown to initial peak during landing phase), P(LC/TH) (extension ROM of lower cervical region from touchdown to initial peak during landing phase), (significant in univariate tests)

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87 Table 4-6. Collapsed mean and SD values of kinetic variables at L/S and C/T junctions for different genders a nd fatigue levels. SL1 SL2 Kinetic variables Mean (SD) Mean (SD) Gender: p Fatigue: p Fatigue Gender: p L/S junction AxF(L/S) (N kg-1) 5.79 (2.01) 5.76 (1.86) 0.063 0.9 0.019* Male 5.46 (1.82) 4.80 (1.38) Female 6.19 (2.26) 6.91 (1.73) ShF(L/S)ant (N kg-1) 0.84 (0.53) 0.85 (0.41) 0.763 0.877 0.185 Male 0.88 (0.53) 0.76 (0.40) Female 0.80 (0.54) 0.95 (0.43) ShF(L/S)post (N kg-1) 4.40 (2.48) 5.33 (2.45) 0.703 0.002 0.001* Male 4.71 (2.55) 4.66 (2.39) Female 4.02 (2.46) 6.13 (2.39) FlxM(L/S) (N m kg-1 BH-1) 0.93 (0.45) 0.98 (0.42) 0.709 0.576 0.961 Male 0.96 (0.37) 1.01 (0.40) Female 0.90 (0.55) 0.94 (0.47) ExtM(L/S) (N m kg-1 BH-1) 1.93 (0.73) 2.21 (0.68) 0.7 0.01 0.011* Male 2.02 (0.67) 2.02 (0.63) Female 1.82 (0.83) 2.44 (0.70) C/T junction AxF(C/T) (N kg-1) 1.74 (2.40) 2.21 (2.04) 0.541 0.096 0.002* Male 2.04 (2.33) 1.42 (1.28) Female 1.38 (2.56) 3.17 (2.42) ShF(C/T)ant (N kg-1) 2.71 (1.71) 2.66 (1.72) 0.398 0.907 0.715 Male 2.51 (1.64) 2.35 (1.48) Female 2.95 (1.84) 3.04 (1.98) ShF(C/T)post (N kg-1) 4.19 (0.86) 4.53 (0.90) 0.438 0.036 0.699 Male 4.28 (0.98) 4.69 (0.93) Female 4.07 (0.71) 4.35 (0.88) FlxM(C/T) (N m kg-1 BH-1) 1.84 (0.80) 1.94 (0.76) 0.782 0.474 0.787 Male 1.78 (0.73) 1.91 (0.63) Female 1.90 (0.92) 1.97 (0.92) ExtM(C/T) (N m kg-1 BH-1) 2.21 (0.72) 2.53 (0.70) 0.454 0.022 0.054 Male 2.24 (0.63) 2.30 (0.59) Female 2.17 (0.86) 2.79 (0.77) Note: SL1 (soft landing before the fatigue procedur es), SL2 (soft landing after the fatigue procedures), AxF (peak axial compressive force), ShFant(or post) (peak ant. or post. shear force), FlxM (peak flexor moment), ExtM (peak extensor moment), (L/S) (fo r lumbosacral junction), (C/T) (for cervicothoracic junction), (significant in univariate tests)

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88 Figure 4-1. Significant interactio ns of the touchdown angle and C/T kinetic variables between gender and landing technique: (A) TD(TH/TL) was decreased from NL to SL condition in males, but not decreased in females (t horacic regional angle becomes more flexed from NL to SL condition in males, but not in females), (B) ShF(C/T)post of males was greater than that of females during NL, but not different from each other during SL condition; ShF(C/T)ant of females was greater than th at of males during NL, but not different from each other dur ing SL condition; ShF(C/T)ant was decreased from NL to SL condition in females, but not decreased in males. Note: NL (self-selected normal landing), SL (soft landing).

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89 Figure 4-2. Representative kine matics of each spinal region ( L/S: lumbar regional angle relative to the sacral region, TL/L: thoracolumbar regional angle relative to the lumbar region, TH/TL: thoracic regional angle relativ e to the thoracolumbar region, LC/TH: lower cervical regional angle rela tive to the thoracic re gion), knee flexion angle ( KFA) and hip flexion angles ( HFA) in one male and one female subjects during NL and SL conditions. For the direction of motion, nega tive slope represents the flexion motion. Flexion of a spinal region means the forw ard rotation of a region relative to the adjacent lower region. Note: LP (landing phase), NL (self selected normal landing), SL (soft landing).

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90 Figure 4-3. Representative kine matics of each spinal region ( L/S: lumbar regional angle relative to the sacral region, TL/L: thoracolumbar regional angle relative to the lumbar region, TH/TL: thoracic regional angle relativ e to the thoracolumbar region, LC/TH: lower cervical regional angle rela tive to the thoracic re gion), knee flexion angle ( KFA) and hip flexion angles ( HFA) in one male and one female subjects during SL1 and SL2 conditions. For the direction of motion, nega tive slope represents the flexion motion. Flexion of a spinal region means the forw ard rotation of a region relative to the adjacent lower region. Note: LP (landing phase), SL1 (soft landing before the fatigue procedures), SL2 (soft landing af ter the fatigue procedures).

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91 Figure 4-4. Significant interactio ns of the kinetic variables at L/S and C/T junctions between gender and fatigue level: (A) AxF(L/S) wa s increased from SL1 to SL2 condition in females, but not increased in males; ShF(L/S)post was increased from SL1 to SL2 condition in females, but not increased in males; AxF(C/T) was increased from SL1 to SL2 in females, but not increased in ma les, (B) ExM(L/S) was increased from SL1 to SL2 condition in females, but not incr eased in males. Note: SL1 (soft landing before the fatigue procedures), SL2 (sof t landing after the fatigue procedures).

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92 CHAPTER 5 DISCUSSION Effects of Landing Technique Previous studies on landing mechanics focused mostly on biomechanical characteristics of lower extremities. Majority of these studies focused on defining injury mechanism (Fagenbaum & Darling, 2003; Wikstrom et al., 2006), ge nder differences (Kernozek et al., 2005) and conditional variability (Schot et al., 2002) associated with knee and ankle joints. To our best knowledge, no attempt has been made to evaluate the mechanical characteristics of upper body segments above the hip joints during drop landings. The landing variables used in the current study identified the overall la nding characteristics including knee and hip joints. Th e landing technique used in soft landing significantly decreased the PVGRF, extended the landing phase, exhib ited flexed landing postures of knee and hip joints, and exhibited more flexion motion of knee and hip joints. Additionally, female subjects demonstrated a greater degree of hip flexion than male subjects, and th is may indicate that females demonstrate absorbing type landi ng without any different instructions. Horita et al. (2002) proposed two types of la nding performances with regard to the prelanding motion of the knee joint. The proper pr e-landing movement was characterized by knee flexion just before touchdown, whic h is associated with a high init ial joint stiffness coupled with high joint power. This was called bouncing type in their study, which is similar to plyometrics. On the other hand, an inadequate pre-landing movement, associated with incomplete knee flexion induced subsequent deep-knee flexion after touchdown, was called the absorbing type of landing and was regarded as a poor strategy. Th e absorbing movement comes too late and demands longer contact time and lower takeoff ve locity. A recent study by Park et al. (2006a) revealed that the soft landing st rategy used by males was close to the bouncing type which is fit

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93 for the effective subsequent movements, and that of females was close to absorbing type of landing which is just pursuing the dissipati on of increased loads from ground impact. The lumbar region was in a flexion posture immediately before touchdown and reached its maximal flexion around the end of landing phase, a nd this was followed by a rapid extension to neutral after the landing phase. Fo r the purpose of discussion, the ne utral angle of a spinal region is defined as the spinal regiona l angle recorded during a standing posture at the end of a landing trial. Flexion motion of a spinal region refers to forward rotation of the spinal region relative to the lower adjacent spinal region. The thoracic region started an extension motion after touchdown followed by a flexion to neutral towa rd the end of landing phase. However, the thoracic spine moved more gradually compar ed to the movement of lumbar region. The thoracolumbar region had a short duration of sl ight extension around touchdown and reached its initial peak at the beginning of the landing pha se followed by a flexion motion similar to the lumbar region. The lower cervical region exhi bited an extension motion around touchdown, and reached its full extension during the landing phase followed by a rapid flexion to neutral. Most spinal regions demonstrated biphasic motions (i.e., extension followed by flexion or flexion followed by extension), but the thoracolumbar region demonstrated multiphasic motion during the landing phase of drop landing. The biphasic pattern of spinal motion has b een observed during walking. Crosbie et al. (1997) studied the patterns of spinal motion during walking using a model including upper and lower trunks, lumbar and pelvis segments. They used a surface marker in each spinal segment on the back surface of the subjects. The pattern of flexion/extension of each segment was generally biphasic throughout the gait cycle. The pelvis rotate d into negative pelvic tilt at heel strike, and this was followed by a counter-motion to a ma ximum positive pelvic tilt in the single support

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94 phase. The lumbar spine showed maximum flexi on at heel strike, and this was followed by a rapid extension to neutral until the single support phase. The lower thoracic segment extended maximally at heel strike, and returned to a neut ral at mid-stance, then extended through the late stance phase. Kinematics of the spinal column after t ouchdown of drop landing was addressed with touchdown angle and extension ROM variables. Re lative to research question Q1 (any gender differences in spinal kinematics?), the null hypot hesis 1a was rejected by the finding of gender differences in thoracolumbar touchdown angl e and extension ROM of the thoracic region. Females demonstrated a more extended posture in the thoracolumbar region at landing and more extension motion in the thoracic region duri ng the landing phase compared to males. The greater thoracic region extension may be related to the greater hip flexion in females (balance control). An extended lower ex tremity landing posture and more flexion motion of lower extremity joints are considered as specific landi ng characteristics in females (Decker et al., 2003). According to the work-energy relationship, the average vertical GRF experienced by the subject during a drop landing depends on the verti cal displacement of whole body CG (center of gravity) during the landing phase. Because an incr eased flexion motion of the hip joint will lower the body CG, the action of the thoracic region may compensate the lowered CG during the landing phase. As a result, the energy absorbing procedure in the landing phase may become less effective in females because of the movements of spinal column against the energy absorption such as the extension motion of the thoracic region. Relative to research question Q2 (any change s in spinal kinematics by the different landing technique?), the null hypothesis 2a was rejected by the finding of significant changes in P(TH/TL), P(LC/TH), and TD(TH/TL) across different landing techniques. When going from NL to SL

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95 condition, the thoracic and lowe r cervical regions exhibited more extension motion. Males exhibited more flexed landing posture in the t horacic region from NL to SL, while females did not or showed extended posture. Active movement of each spinal region in bot h landing conditions may indicate that the spinal column is actively i nvolved in energy absorption duri ng the landing phase. Without analyzing the flexion ROM in each spinal regi on (only extension ROMs were analyzed), the flexion ROM of each spinal region could not be apparently differentiated with each other. However, the very small initial peaks of extens ion observed in the lumbar region across different landing techniques indicate that the lumbar region exhibited flexion motion in most cases. Additionally, during a soft landing, the thoracic a nd lower cervical region s demonstrated active extension motions against the lumbar motion. This may suggest that the en tire spinal column is involved in energy absorbing pr ocedure during the NL condition, while only part of spinal column is involved during the SL condition. The part ial involvement is more apparent in females during the soft landing. Futhermore, the gender di fferences in thoracic landing posture during soft landings may suggest that the spinal column is less involved in energy absorbing procedure of females during soft la ndings (see Figure 4-4). The results also suggest that the thoracol umbar region could be highly stressed by the simultaneous motions of thoracic and lower cerv ical extension and lumbar flexion during the landing phase of soft landing and is more pr onounced in females. Schache et al. (2003) performed a kinematic study of the lumbo-pelvic-hip complex during running to define the gender differences. They found that females displa yed a shorter stance time, swing time, stride time and stride length, and a higher stride rate than males. Mean waveforms were different in the peak-to-peak oscillations and the offset of pelvis anterior/posterior tilt. Females displayed greater

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96 amplitudes of lumbar spine lateral bend and axial ro tation, pelvis anterior/pos terior tilt, obliquity and axial rotation, and hip adduction/abduction than their male counterparts. The mean positions of anterior pelvic tilt across the running cycle were 20.2 for females and 16.9 for males. The prevalence of pelvic-femoral stress fractures in female runners might be explained by these findings (Bennell et al., 1996; Pavl ov et al., 1982). The greater amplitude of lumbar spinal movements in females during drop landings is similar to the findings observed in running. The increased extension movement of the t horacic region relative to the thoracolumbar region from NL to SL condition suggests that the so ft landing procedure intended to decrease lower extremity loads may be a risk factor for developing thoracolu mbar or upper lumbar degeneration and spondylolysis in physically active individuals. Degenerative disc disease and spondylolysis are the most common structural abno rmalities associated with low-back pain in athletes. Disc degeneration appear s to be influenced by the type and intensity of the sport. Videman et al. (1995) demonstrated that weight lifter s have a higher rate of and more severe degenerative changes in the upper lumbar spine, whereas back problems in soccer players are almost exclusively in the L4 to S1 levels. Cappozzo et al. (1985) found that, when a person performed half-squat exercises with weights approximately 1 .6 times body weight, compressive loads across the L3/L4 motion segment were about 10 times body weight. The prevalence of spondylolysis in athletes is vari able, but some sports appear to be associated with a higher prevalence rate. Rossi and Dragoni (1990) reported a rate of 43% in divers, 30% in wrestlers, and 23% in weight lifters. A lthough the exact mechanism for the development of spondylolysis is not known, there is some suggestion that it may be a fatigue fracture following repeated hyperextension.

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97 Joint resultants at L/S and C/T junctions af ter touchdown of drop la nding were addressed with the L/S and C/T kinetic variables. Relative to research question Q1 (any gender differences in L/S or C/T kinetics?), the null hypothesis 1b wa s rejected by the finding of significant gender differences. Posteriorly directed shear for ce of the trunk segment at C/T junction was significantly greater in male subjects than in female subjects during NL, but no gender difference for the SL condition was found. Females demonstrated decreased ShF(C/T)ant from NL to SL condition, while males did not demonstrate any changes in ShF(C/T)ant across landing techniques (see Figure 4-1). ShF(C/T)ant of females was greater than that of males during NL, but no difference between genders was found for the SL condition. Increased posteriorly directed shear force in males and increased anteriorly directed shear force in females for the NL condition can place much stress on supporting anatomical structures. The cervical spine can be injured due to increas ed shear forces at the lower cervical region by a whiplash. During a whiplash, hyperextension of head and n eck is the basic mechanism for cervical spine injury and it commonly occurs in the rear-end impact in motor vehicle accidents (Luan et al., 2000). Studies of the natural hist ory of whiplash-associated disorders have suggested that chronic pain with continued sy mptoms develops in 6-33% of acutely injured patients (Hildingsson & Toolanen, 1990). Previous biomechanical studies have focused on injury mechanisms of the cervical facet joints and the in tervertebral discs as potential structures to develop whiplash-associated disorders. Follo wing a whiplash, the lower cervical spine experiences complex loading consisting of an extension moment, posterior shearing and compressive forces. This loading pattern ha s been hypothesized to injure the cervical intervertebral disc and facet jo ints (Panjabi et al., 2004).

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98 Facet joints of the lower cervi cal spine can be compressed by the increased posterior shear force, and a capsular ligament strain can be di stracted by the increased anterior shear force during the NL condition. Pearson et al. (2004) evaluated peak facet joint compression and capsular ligament strain using a whole cervical spine specimen with muscle force replication and a bench-top trauma sled to simulate whiplash of increasing severity. Peak facet joint compression was greatest at C4/C5 and reached over physiologic limits during the 3.5g simulation (low-level acceleration). Capsular ligament strains exceeded the physiologic strains at 6.5g and were largest at C6/C7 during the 8g simulation (high-level acceler ation). They concluded that peak facet joint compression occurred at maximum intervertebral extension, whereas peak capsular ligament strain was reached as the facet joint was returning to its neutral position after the maximum intervertebral extension. The great er degree of posteriorly directed shear force in males than in females may indicate that male subjects could be in risk of increased facet joint compression, while females could be in risk of increased capsular ligament strain due to the increased anteriorly directed shear force during NL condition. Panjabi and co-workers (2004) have studied a whole cervical spine model on a bench-top sled with muscle force replication and a surr ogate head to simulate rear-end impact. They underwent standard flexibility testing to determine the sagittal disc deformation with the various acceleration of the bench. They observed th e greatest strain in the posterior 150 fibers running posterosuperiorly at C5/C6. They also observed incr eased disc shear strain at the posterior region and increased axial deformation at the anterior re gion of the disc at C5/C6. They concluded that the cervical intervertebral disc is at risk for injury during a whiplash because of excessive 150 fiber strain, disc shear strain, and anterior axial deformation. Th e injury of the cervical spine by

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99 facet joint distraction may occur at the level lower than the cervical injury by the facet compression. Relative to research question Q2 (any cha nges in L/S or C/T kinetics by the landing technique?), the null hypothesis 2b was rejected by the finding of significant changes in all kinetic variables except for ShF(C/T)ant across different landing t echniques. During the SL, all kinetic variables except for ShF(C/T)ant decreased significantly comparing to NL condition. This may indicate that the soft landi ng technique used in the curren t study is an effective way to decrease the overall loads applied to L/S and C/T junctions However, ShF(C/T)ant was not decreased by the soft landing. The greatest loads at L/S juncti on were AxF(L/S) and ShF(L/S)post, while the ShF(C/T)post was the greatest load at C/T junction during drop landing. Calla ghan et al. (1999) conducted a biomechanical study using two models to estima te loads applied to the L4/L5 level during walking task: linked segment model with EMG t echnique and rigid segment model with inverse dynamic technique. The joint loading at L4/L5 ca lculated by the EMG model resulted in large increases in the maximum compressive forces (3 .5 times of body weight), compared with the joint reaction forces calculated using inverse dynamics (1.0 times of body weight). Including the muscular component resulted in a mo re than three-fold increase in joint load. Howe ver, the joint shear forces (anterior/posterior, lateral) obtained using the two t echniques were quite similar. The peak compressive axial forces at L/S j unction were 8.5 and 5.8 tim es of body mass for NL and SL conditions, respectively. Th e posteriorly directed shear fo rce at L/S junction from the current study (9.7 body mass in NL, 4.5 body mass in SL) was much higher than the values for their walking trials (EMG model: 0.18 BW, Inverse dynamics model: 0.19 BW). In the study of Callahan et al. (1999), there was a peak flexor moment at heel contact followed by a

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100 peak extensor moment around toe-off. During a fa ster speed gait, the flexion/extension moment at L4/L5 shifted to the extension side and demo nstrated a high extensor moment around toe-off. Callahan et al. concluded that the loads and motions for the lumbar spine during gait depended on the walking speed. Increasing walking speed in creases the lumbar spine ROM, activation of spinal and trunk muscles, and ante rior/posterior shear forces. Likewise, the loads and motions for the lumbar spine during drop landing depend on the landing technique which controls the involvement of body segments including lower extremity joints and spinal regions. One of the limitations of this study is that sagi ttal spinal kinematics re lative to the adjacent spinal region based on the spinal skin markers cannot precisely determine the vertebral motions in each spinal region. Additionally, there might be some errors in kinetic measures, because the locations of joint center and spinal junction (used for joint resultants computations) were estimated using surface markers instead of dete rmined using radiographic imaging techniques. Mechanical characteristics of the lower extremitie s were assumed to be symmetrical and only the GRF data collected from the left leg were used to determine joint resultants at the L/S and C/T junctions. Despite these limitations, the results fro m this study provide in sight into the spinal movement during two different landing techniques. Effects of Knee Joint Muscles Fatigue The purpose of the second part of this study was to determine whether spinal mechanics were affected by lower extremity fatigue dur ing drop landings. The overall landing mechanics did not change significantly by the presence of knee joint muscles fatigue in the current study. Instead, a significant gender difference in landing posture was found in the knee joint. Females landed with a more extended knee jo int posture than males which is consistent with the findings reported in a previous study (Decker et al., 2003). These authors found that females demonstrated a more erect landing posture and utilized greater hip a nd ankle joint range of

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101 motions and maximum joint angular velocities when compared to males. Females exhibited greater energy absorption and peak powers from the knee extensors and ankle plantar-flexors compared to males. Energy absorption contribut ions revealed that the knee extensor was the primary shock absorber for both genders. The ankle plantar flexor was the second largest contributor to energy absorption for the females and the hip extensor was for the males. The different shock absorption strategy used in fema les was proposed to prov ide a greater potential risk for non-contact ACL injury for females under certain landing conditions. Kinematics of the spinal column after t ouchdown of drop landing was addressed with touchdown angle and extension ROM variables. Re lative to research question Q3 (any gender differences in spinal kinematics?), the nu ll hypothesis 3a was rejected by the finding of significant gender differences in thoracolumbar touchdown angle and extension ROM of the thoracic region. Females demonstrated a more ex tended posture in the thoracolumbar region at landing and more extension motion in the thoraci c region during the landi ng phase compared to males. This is quite similar to the findings of lo wer extremity joints in th e study of Decker et al (2003). The greater extension motion of the thorac ic region is supposed to compensate the extended postures in the knee jo int and thoracolumbar region at landing in females (balance control). An extended lower ex tremity landing posture and more flexion motion of lower extremity joints are considered as specific landi ng characteristics in females (Decker et al., 2003). The extended postures of the knee and thoraco lumbar region will raise the CG location of the whole body and increase the stiffness of these regions comp ared to their male counterparts. The increased extension motion of the thoracic region will further raise the body CG and may

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102 contribute to an increase in s tiffness of the whole body during the landing phase for females during the fatigue landing condition. Kujala et al. (1997) studied the impact of lumb ar flexibility on low-back pain in a 3-year longitudinal study. They examined lumbar ROM in a group of adolescent athletes and nonathlete controls across both genders. Neither groups ha d previous low-back pa in, nor were lumbar measurements performed during episodes of pai n. While no differences were detected between athletes and nonathlete controls in males, female nonathletes e xhibited greater overall lumbar ROM and lower lumbar ROM than did female athl etes. They suggested that decreased ROM in the lower lumbar segments and decreased maximal lumbar extension as the predictive factors of low-back pain in women, because the girls w ithin the lowest quartile of maximal lumbar extension developed 3.4 times the chance of having pain lasting more than one week. Sward et al. (1990) studied lumbar mobility, in relation to back pain, in male at hletes, but no correlation was found between spinal flexibili ty and back pain. Both prev ious studies support that high mobility of the lumbar region is essential to main tain the lower-spinal integrity in females, and the active women or female athletes who were involved in vigorous activities demanding much spinal motion will have more chance to devel op limitation of spinal movement and low-back pain. Gender differences in kinematic characteri stics observed in the cu rrent study suggest that higher loads due to the hyperextension of the thoracic region could be placed on the thoracolumbar region of females during soft landi ng and higher incidence of low-back pain may be found in female athletes or active women who perform repeat ed jumps and landings. Without enough flexibility and strength of thoracolumbar to lu mbar region, females are likely to be at risk to develop spinal injury or back pa in due to repeated soft landings.

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103 The kinematic pattern of each spinal region in the current study is comparable to the findings observed in previous walking and running studies. Crosbie et al. (1997) evaluated spinal kinematics during walking tasks. The lumbar sp ine showed maximum flexion at heel strike followed by a rapid extension to neutral until the single support phase. The lower thoracic segment extended maximally at heel strike, and retu rned to a neutral posture at mid-stance, then extended through the late stance phase. Schache et al. (2002) obser ved the kinematics of pelvis and lumbar spine during running trials. The lumb ar spine flexed slig htly and the pelvis posteriorly tilted slightly dur ing the loading phase. The lumbar spine was extended while the pelvis was anteriorly tilted dur ing the swing phase. Both lumbar spine and pelvis displayed a biphasic movement pattern that corresponds to on e phase per each step. The biphasic patterns of pelvis, lumbar and thoracic spines in both prev ious studies resemble the movement pattern of each spinal region during soft landing trials. In the current study, the lumbar region star ted a flexion motion before the touchdown and reached its peak around the end of landing phase, and this was followed by a rapid extension to neutral after the landing phase. For the purpose of discussion, th e neutral angle of each spinal region in the current study is defined as the sp inal regional angle reco rded during the standing posture at the end of a land trial. The thor acic region started an extension motion around touchdown followed by a gradual flexion to neut ral toward the end of landing phase. The thoracolumbar region had a short duration of slight extension im mediately after touchdown and reached its initial peak extension at the begi nning of the landing phase followed by a flexion motion similar to the lumbar region. The lower cervical region started an extension motion around touchdown and reached its peak extension around the end of the landing phase followed by a rapid flexion to neutral. Generally, most spinal regions demonstrated biphasic motions

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104 during the landing phase of soft landing, but th e thoracolumbar region exhibited multiphasic motion. The kinematic pattern of spinal column obs erved in the soft landi ng trials suggests that the thoracolumbar region could be highly stress ed by simultaneous lumbar flexion and thoracic extension during a greatly loaded landing phase. Relative to research question Q4 (any change s in spinal kinematics during soft landing by the fatigue procedures?), the null hypothesis 4a was supported by a lack of significant kinematic differences in any spinal regions between SL 1 and SL2. Fatigue procedures applied to knee flexors and extensors in the current study was in troduced by repeated bouts of maximal work of target muscles, because muscular fatigue has b een defined as the development of less than the expected amount of force as a consequence of continuous voluntary muscle contractions (McCully et al., 2002). The expressi on of muscle fatigue as a per cent reduction in torque output has been used as a fatigue index (Katsiaras et al., 2005). The fatigue index is typically computed as the ratio of the average torque of the last five repetitions to th e average torque of the first five repetitions in a 30-repetition maximum effort tr ial. However, this quantification of muscle fatigue has been questioned in term s of its reliability. Because torque variability within the initial five contractions has a potentia tion effect, muscle fatigue appear ed to be underestimated when the first five repetitions were f actored into the fatigue index form ula (Neptune et al., 1997). Since torque output during maximal contractions at 3.14 rad s-1 has been shown to increase from repetitions four to as high as 10 (Wretling & Henriksson-Larsen, 1998), a more accurate estimate of muscle fatigue can be obtained with the highest consecutive five repetitions as the subjects best performance (Pincivero et al., 2003). Muscular fatigue is manifested as a pr ogressive decline in power output, and the magnitude of which is largely determined by th e duration of the interv ening recovery periods

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105 (Glaister, 2005). Thirty minutes of running was included during the fatigue procedures of the current study for the purpose of delaying the re covery period. In general, hypoxic condition is associated with the increased rates of fati gue. For example, under the hypoxic condition during the long-distance running, the decreased abi lity to perform cycled maximal sprints was associated with an increased accumulation of bl ood lactate, a reduced oxygen uptake, and an increased rate of muscular fatigue. Oxygen availa bility mediates its effect on multiple sprint performance by influencing the magnitude of the aerobic contribution to ATP (adenosin triphosphate) resynthesis duri ng work periods, and the rate of phosphocreatinine resynthesis during intervening rest peri ods (Balsom et al., 1994). Data collected from 3 female and 1 male s ubjects were not included in the statistical analyses because their fatigue indi ces did not increase due to the fa tigue procedure. It is possible that the intensity of fatigue procedure was not high e nough for these 4 individuals. More exclusion of female subjects than male subj ects from the current fatigue procedure may be explained by the fact that females are less sus ceptible to muscular fatigue. Hunter and Enoka (2001) observed that females sustained the maxi mal voluntary contraction for a much longer period of time (118% longer) than males and duri ng low-level contraction (20% of the maximal voluntary contraction). It has been suggested that muscle fibers in females possess a relatively greater oxidative capacity than males, which w ould enhance the respiratory capacity of the contracting muscles. On the other hand, males may possess an inherently greater ability to generate more force than females, which may be related to a significantly greater proportion of fast-twitch muscle fibers in the sk eletal muscle (Fulco et al., 2001). No significant changes in la nding posture in lower extremity joints due to the fatigue procedures may indicate that stretch reflex activ ity of the knee joint musc les was not affected by

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106 the fatigue procedures and the fatigue level wa s not high enough to cause any kinematic changes or to be detected by the current measuremen ts. The ground contact phases of running, jumping and hopping are examples of the stretch-shorte ning cycle (SSC) type of exercises for leg extensor muscles. Duing a SSC, the preactivated mu scle is first stretche d (eccentric action) and then followed by the shortening (concentric) ac tion. The stretched phase is mediated by the reflex activity of the extensor muscles before the ground impact of the foot during jumps and landings (Horita et al., 2002). As the SSC demand s a strong mechanical loads to the skeletal muscles, its influence on the reflex activation is essential to perform the fast and smooth SSC. Intensive SSC-type exercise results in reversib le muscle damage. This is associated with delayed-onset muscle soreness, and with propr ioceptive and neuromuscu lar impairments that may last for several days. These neuromuscular perturbations are typically associated with changes in muscle mechanics and activation th at result in major cons equences on joint and muscle stiffness regulation in SSC-type performan ce. This performance deterioration is called as neuromuscular fatigue and subsequent long-term recovery will take place in a bimodal fashion. In this bimodality, the acutely induced re duction in electromyographic activity (maximal voluntary contraction) is followe d by a short-term recovery (wit hin 2 hours), which is in turn followed by a secondary reduction with longer lasting recovery (1-2 days pos t-exercise) (Nicol et al., 2006). Joint resultants at L/S and C/T junctions af ter touchdown of drop la nding were addressed with the L/S and C/T kinetic variables. Relative to research question Q3 (any gender differences in L/S or C/T kinetics?), the null hypothesis 3b wa s rejected by the finding of gender differences in AxF(L/S), ShF(L/S)post, ExtM(L/S) and AxF(C/T). Males exhibited greater ShF(L/S)post,

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107 ExtM(L/S), and AxF(C/T) than females duri ng SL1, but females showed greater ShF(L/S)post, ExtM(L/S), and AxF(C/T) than male s during SL2 condition (see Figure 4-4). Significant kinematic gender differences found in the current study are extended landing posture of the knee joint and t horacolumbar junction and more extension motion of thoracic region during the landing phase in females. These findings may s uggest that the knee joints and lumbar to thoracolumbar region are stiffer duri ng the landing phase in females when comparing to males. This also suggests that males utilize th e spinal column in energy absorption in addition to lower extremities. However, despite the low fatigue level of the Knee joint muscles, females exhibited significant increases in joint resultants at L/S and C/T junctions. Because females rely more heavily on lower extremity action in en ergy absorption during the landing phase, their landing mechanics are likely to be more affected by lower extremity fatigue than their male counterparts. Just like the interaction between le g stiffness and reflex ac tivities (Horita et al., 1996), reflex activity of lumbar (to thoracolumbar region) spinal extensors also may play a role to maintain lumbar stiffness during the initial landing phase. Relative to research question Q4 (any changes in L/S or C/T kinetics during soft landing by the fatigue procedures?), the null hypothesis 4b was rejected by the finding of significant changes in AxF(L/S), ShF(L/S)post, ExtM(L/S) and AxF(C/T) from SL1 to SL2 condition in females. Females exhibited increased AxF(L/S), ShF(L/S)post, ExtM(L/S) and AxF(C/T) when going from SL1 to SL2 condition, while ma les did not (Figure 4-4 on p. 91). The significant increases in joint resultants at L/S and C/T junctions due to the fatigue procedure in females illustrate the significant role of spinal column in energy absorbing procedure during drop landing. However, the spinal column may play a different role in energy absorbing procedure in each gender. A recent st udy by Park et al. (2006b) also suggested the

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108 important role of lumbar region during the land ing phase. They studied the effects of limited lower back motion on soft landing mechanics of lo wer extremity joints with subjects wearing soft and hard low-back braces. Limited spinal motions by the brace caused alterations in knee and hip joint motions during the landing phase and an increase in impact force. Females demonstrated increased knee extensor moment and males showed increased axial force at the hip joint in the hard brace condition. Increased axial force at C/T junction in females during the landing with knee joint muscles fatigue can place much stress on the cervical spin e in some instances. Axial loading has been reported as a mechanism of catastr ophic cervical spine injuries in football players (Torg et al., 1990). Nightingale et al. ( 1996) reported that straightening of the cervical spine before injury may be another necessary element of the compressi ve injury mechanism. Cervical spine trauma accounts for about 25% of entire spin e injuries in sports (Leidholt, 1963). Cervical spine injury is more common in some sports than in others: football, gymnastics, r ugby, baseball, lacrosse, judo, skiing, jumping on trampolines, and diving (Torg et al., 2002). Trad itionally, hyperflexion and hyperextension have been implicated as the pr imary mechanisms of cervical spine injuries (Gehweiler et al., 1979; Macnab, 19 64; Paley & Gillespie, 1986). Torg (1987) provided the injury mechanism to the cervical spine in football player. A typical situation is that of a defensive back making a tackle involving contact with the other player with the top of the helmet. The injury mechanism involves axial loading with an element of buckling. With the neck in a ne utral position, the cervical spine is extended as a result of the normal cervical lordosis. When the neck is flexed to 30 the cervical spine becomes straight. When a force is applied to the vertex, the en ergy inputs are transmitted along the longitudinal axis of the cervical spine. The cervical spine loses its ability to di ssipate force and being

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109 compressed between the abruptly decelerated head and the force of the upcoming trunk. This is a typical example of cervical spine in jury with the axial loading as the injury mechanism, which is characterized as a clustering of in juries in the middle part (third and fourth) of the cervical spine. Although the axial force at C/T junction during dr op landings is much lower than the direct compression of vertex during cont act sports, repeated landings with fatigued knee joint muscles may increase the chance of injuries to the cervical spine in females. Luan et al. (2000) analyzed neck kinematics and loading patterns with high-speed X-ray video cameras during rear-end impact simula tion study using a cadaver body. They observed compression, tension, shear, flexion and extension at different cervical le vels during different stages of the whiplash event. They reported that compression of the neck is due to the straightening of the thor acic spine and possibly the upward rampi ng of the torso during the initial stage of the impact. Likewise, straightening of either cervical spine or thoracic spine will increase the axial compressive force more at the lower cervical region during the soft landing procedures. Accordingly, to avoid cervical inju ries, individuals with spinal degeneration or surgical fusion in the cervical spine may need to refrain from repeated landing type activities. In addition to the limitations mentioned previ ously, another limitation of this study is that the effects of fatigued muscle co uld not be precisely differentiate d between flexor and extensor of the knee joint.

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110 CHAPTER 6 CONCLUSION Effects of Landing Technique The spinal column is actively involved in energy absorbing procedure during the landing phase. However, males may use the the spinal column as the active weight bearing segment during the landing phase with less motion compari ng to females. Females may demand spinal column in their energy absorbing procedure in landing phase but they need more extension motion comparing to males. Vigorous extension motions of the thoracic and lower cervical regions during soft landing suggest that the spin al column is less involved in energy absorption, while the entire spinal column is involved during the normal landing condition. The thoracolumbar region is likely to be highly stressed by the simultaneous motions of thoracic extension and lumbar flexion in females during the landing phase of soft landing. The cervical spine can be injured by the great posterior shear force in males through the facet joint compression, and can be injured by the great anteri or shear force in females through the capsular ligament distraction during drop landings. Effects of Knee Joint Muscles Fatigue Despite the low level of the knee joint muscles fatigue, females demonstrated increased joint resultants at L/S and C/T juctions. Becau se extended landing postures of knee joint and thoracolumbar region were not compensated e ffectively by the increas ed motion of thoracic region during the landing phase of females, spinal mechanics are easily affected by the small changes in fatigue. It is also speculated that energy absorption in females could be mostly moderated by the extensor actitiv ities of knee joints. Additinall y, the thoracolumbar region could be stressed when simultaneous lumbar flexion and thoracic extension occurred during a highly loaded initial landing phase of soft landi ngs, and it was more prominent in females.

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111 Conclusively, knee joint muscle fa tigue during soft landing of female s can be a risk factor to the lower lumbar and cervical spine injuries because of increased joint resultants.

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112 CHAPTER 7 FUTURE WORK The landing task can be an idea l situation to eval uate spinal mechanics for the young and active population. The measurements us ed in the current study could be applied to individuals in different athlete and patient popul ations. In addition to dynamic X-rays, the movement of each spinal region may provide valuable informati on on roles played by spinal regions in dynamic tasks. The data collected from patient groups ma y provide insight into the spinal function for different populations, and could be useful for de cisions relative to conservartive and surgical treatments. The inverse dynamic approach used in the current study can be compared with other modeling approaches. The inverse dynamic appr oach can only be used to determine joint resultants. Other modeling approaches (e.g. EMG-op timization model) have to be used in order to estimate contact forces at a joint.

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113 APPENDIX A INFORMED CONSENT The informed consent approved by the Institutional Review Board of the University of Florida.

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118 APPENDIX B PRELIMINARY STUDY One healthy active male (age: 21 yrs, mass: 80.7 kg, height: 179.7 cm) and one female (21 yrs, 59.9 kg, 168.3 cm) were tested in this pilo t study. They were free fro m any musculoskeletal diseases or injuries which could influence spin al and lower extremity joints mechanics during physical activities, and signed informed consen t approved by the IRB of the University of Florida. The experimental setup, testing procedur es and data reduction were the same as what was described in Chapter 2. The peak extensor moments at the L/S j unction were used to determine sample size for th e current study (Table B-1). Table B-1. Kinetic characteris tics of L/S junction data fo r sample size justification. NL SL1 SL2 AxF(L/S) Mean (SD) 10.47 (0.25) 6.85 (0.57) 7.17 (2.47) (N kg-1) Male 10.65 7.25 5.42 Female 10.29 6.44 8.92 ShF(L/S)ant Mean (SD) 1.80 (0.88) 0.84 (0.35) 0.87 (0.53) (N kg-1) Male 2.42 0.59 0.49 Female 1.18 1.08 1.24 ShF(L/S)post Mean (SD) 7.35 (1.24) 4.37 (1.04) 6.11 (0.33) (N kg-1) Male 8.23 3.63 6.34 Female 6.47 5.10 5.88 FlxM(L/S) Mean (SD) 1.74 (0.48) 1.14 (0.09) 1.20 (0.08) (N m kg-1 BH-1) Male 1.40 1.20 1.14 Female 2.08 1.07 1.26 ExtM(L/S) Mean (SD) 3.05 (0.38) 2.27 (0.15) 2.59 (0.19) (N m kg-1 BH-1) Male 3.32 2.16 2.72 Female 2.78 2.37 2.45 Note: NL (self-selected normal landing), SL1 (soft land ing before fatigue proce dure), SL1 (soft landing after fatigue procedure), AxF (peak axial compressive force), ShFant(or post) (peak ant. or post. shear force), FlxM (peak flexor moment), ExtM (peak extens or moment), (L/S) (for lumbosacral junction).

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119 APPENDIX C MANOVA AND ANOVA TABLES Effects of Landing Technique Table C-1. MANOVA table for la nding variables, touchdown angl e variables, extension ROM variables, and L/S and C/T kinetic variables in the study of landing technique effects. Effect Roy's Largest Root F Hypothesis df Error df Sig. (p) Observed Power (a) Landing variables Gender* 2.824 8.942 6 19 <0.001 1.0 Landing* 8.148 25.8036 19 <0.001 1.0 Landing Gender 0.797 2.523 6 19 0.057 0.7 Touchdown angle variables Gender* 0.807 4.236 4 21 0.011 0.85 Landing 0.287 1.509 4 21 0.236 0.39 Landing Gender* 0.659 3.459 4 21 0.025 0.77 Extension ROM variables Gender* 0.788 4.136 4 21 0.013 0.84 Landing* 1.657 8.700 4 21 <0.001 0.99 Landing Gender 0.290 1.523 4 21 0.232 0.39 L/S Kinetic variables Gender 0.269 1.076 5 20 0.403 0.3 Landing* 3.801 15.2065 20 <0.001 1.0 Landing Gender 0.410 1.639 5 20 0.196 0.45 C/T Kinetic variables Gender* 0.773 3.094 5 20 0.031 0.76 Landing* 4.065 16.2595 20 <0.001 1.0 Landing Gender* 0.899 3.597 5 20 0.018 0.83 Significant main effect or interaction (p<0.05)

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120 Table C-2. Univariate tests for the different landing variables in the study of landing technique effects. Measure Source Type III Sum of Squares df Mean Square F Sig. (p) Observed Power(a) PVGRF Gender 0.383 1 0.383 0.025 0.877 0.05 Error 374.297 2415.596 Landing* 861.538 1 861.538 90.268 <0.001 1.0 Landing Gender 1.231 1 1.231 0.129 0.723 0.06 Error (Landing) 229.062 249.544 t(LP) Gender 0.038 1 0.038 2.231 0.148 0.3 Error 0.412 240.017 Landing* 0.230 1 0.230 26.355 <0.001 1.0 Landing Gender 0.003 1 0.003 0.395 0.536 0.09 Error (Landing) 0.209 240.009 TD(KFA) Gender 276.001 1 276.001 3.789 0.063 0.46 Error 1748.330 2472.847 Landing* 136.988 1 136.988 23.797 <0.001 1.0 Landing Gender 7.388 1 7.388 1.283 0.268 0.19 Error (Landing) 138.155 245.756 P(KFA) Gender 5.493 1 5.493 0.021 0.887 0.05 Error 6361.497 24265.062 Landing* 3711.240 1 3711.24071.329 <0.001 1.0 Landing Gender 127.109 1 127.109 2.443 0.131 0.32 Error (Landing) 1248.715 2452.030 TD(HFA) Gender 10.351 1 10.351 0.035 0.853 0.05 Error 7063.826 24294.326 Landing* 348.407 1 348.407 24.873 <0.001 1.0 Landing Gender 1.489 1 1.489 0.106 0.747 0.06 Error (Landing) 336.174 2414.007 P(HFA) Gender* 2089.157 1 2089.15715.422 0.001 0.96 Error 3251.093 24135.462 Landing* 4194.019 1 4194.01950.000 <0.001 1.0 Landing Gender 227.643 1 227.643 2.714 0.113 0.35 Error (Landing) 2013.138 2483.881 Significant main effect or interaction (p<0.05)

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121 Table C-3. Univariate tests for the touchdown angl e variables in the study of landing technique effects. Measure Source Type III SS df Mean Square F Sig. (p) Observed Power (a) Touchdown angle variables TD(L/S) Gender 241.231 1 241.231 1.625 0.215 0.23 Error 3562.197 24148.425 Landing 0.249 1 0.249 0.037 0.850 0.05 Landing Gender 4.327 1 4.327 0.635 0.433 0.12 Error (Landing) 163.634 246.818 TD(TL/L) Gender* 1053.000 1 1053.00011.098 0.003 0.89 Error 2277.088 2494.879 Landing 0.019 1 0.019 0.007 0.933 0.05 Landing Gender 3.250 1 3.250 1.219 0.280 0.19 Error (Landing) 63.971 242.665 TD(TH/TL) Gender 177.970 1 177.970 1.769 0.196 0.25 Error 2415.068 24100.628 Landing 6.797 1 6.797 1.593 0.219 0.23 Landing Gender* 22.231 1 22.231 5.209 0.032 0.59 Error (Landing) 102.432 244.268 TD(LC/TH) Gender 4.099 1 4.099 0.013 0.910 0.05 Error 7534.978 24313.957 Landing 20.188 1 20.188 0.587 0.451 0.11 Landing Gender 1.357 1 1.357 0.039 0.844 0.05 Error (Landing) 825.895 2434.412 Significant main effect or interaction (p<0.05)

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122 Table C-4. Univariate tests fo r the extension ROM variables in the study of landing technique effects. Measure Source Type III SS df Mean Square F Sig. (p) Observed Power (a) Extension ROM variables P(L/S) Gender 1.699 1 1.699 2.030 0.167 0.28 Error 20.093 240.837 Landing 2.862 1 2.862 3.854 0.061 0.47 Landing Gender 1.422 1 1.422 1.915 0.179 0.26 Error (Landing) 17.825 240.743 P(TL/L) Gender 24.647 1 24.647 3.605 0.070 0.45 Error 164.086 246.837 Landing 1.357 1 1.357 0.761 0.392 0.13 Landing Gender 0.739 1 0.739 0.414 0.526 0.10 Error (Landing) 42.814 241.784 P(TH/TL) Gender* 199.685 1 199.685 7.323 0.012 0.74 Error 654.477 2427.270 Landing* 147.236 1 147.236 30.860 < 0.001 1.0 Landing Gender 19.082 1 19.082 3.999 0.057 0.48 Error (Landing) 114.508 244.771 P(LC/TH) Gender 36.056 1 36.056 0.296 0.591 0.08 Error 2924.496 24121.854 Landing* 465.005 1 465.005 9.101 0.006 0.83 Landing Gender 18.125 1 18.125 0.355 0.557 0.09 Error (Landing) 1226.265 2451.094 Significant main effect or interaction (p<0.05)

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123 Table C-5. Univariate tests for the L/S kinetic variables in the study of landing technique effects. Measure Source Type III SS df Mean Square F Sig. (p) Observed Power (a) L/S Kinetic variables AxF(L/S) Landing* 98.533 1 98.533 28.603 < 0.001 0.999 Landing Gender 4.431 1 4.431 1.286 0.268 0.193 Error (Landing) 82.677 243.445 ShF(L/S)ant Landing* 10.593 1 10.593 25.316 < 0.001 0.998 Landing Gender 2.495 1 2.495 5.962 0.022 0.649 Error (Landing) 10.042 240.418 ShF(L/S)post Landing* 375.363 1 375.363 74.279 < 0.001 1.000 Landing Gender 2.646 1 2.646 0.524 0.476 0.107 Error (Landing) 121.282 245.053 FlxM(L/S) Landing* 2.696 1 2.696 23.927 < 0.001 0.997 Landing Gender 0.000 1 0.000 0.002 0.961 0.050 Error (Landing) 2.704 240.113 ExtM(L/S) Landing* 29.280 1 29.280 59.699 < 0.001 1.000 Landing Gender 0.333 1 0.333 0.679 0.418 0.124 Error (Landing) 11.771 240.490 Significant main effect or interaction (p<0.05)

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124 Table C-6. Univariate tests fo r the C/T kinetic variables in the study of landing technique effects. Measure Source Type III SS df Mean Square F Sig. (p) Observed Power (a) C/T Kinetic variables AxF(C/T) Gender 1.265 1 1.265 0.105 0.749 0.061 Error 289.164 2412.048 Landing* 143.989 1 143.989 14.269 0.001 0.952 Landing Gender 11.197 1 11.197 1.110 0.303 0.173 Error (Landing) 242.184 2410.091 ShF(C/T)ant Gender 9.883 1 9.883 1.761 0.197 0.247 Error 134.708 245.613 Landing 5.617 1 5.617 2.641 0.117 0.345 Landing Gender* 9.762 1 9.762 4.591 0.042 0.538 Error (Landing) 51.034 242.126 ShF(C/T)post Gender* 8.360 1 8.360 5.609 0.026 0.623 Error 35.774 241.491 Landing* 21.543 1 21.543 62.879 <0.001 1.000 Landing Gender* 5.066 1 5.066 14.785 0.001 0.958 Error (Landing) 8.223 240.343 FlxM(C/T) Gender 0.099 1 0.099 0.097 0.758 0.06 Error 24.521 241.022 Landing* 8.074 1 8.074 16.232 <0.001 0.971 Landing Gender 0.467 1 0.467 0.940 0.342 0.154 Error (Landing) 11.938 240.497 ExtM(C/T) Gender 1.688 1 1.688 1.113 0.302 0.173 Error 36.404 241.517 Landing* 33.681 1 33.681 47.772 <0.001 1.0 Landing Gender 1.596 1 1.596 2.264 0.145 0.303 Error (Landing) 16.921 240.705 Significant main effect or interaction (p<0.05)

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125 Effects of Knee Joint Muscles Fatigue Table C-7. MANOVA table for the fatigue indices in the study of knee joint muscles fatigue. Effect Roy's Largest Root F Hypothesis df Error df Sig.(p) Observed Power (a) Gender 0.057 0.657 2 23 0.528 0.147 Fatigue* 0.875 10.067 2 23 0.001 0.971 Fatigue Gender 0.047 0.542 2 23 0.589 0.129 Significant main effect or interaction (p<0.05) Table C-8. Univariate tests of within-subjects ef fects for the fatigue indi ces in the study of knee joint muscles fatigue. Measure Source Type III Sum of Squares df Mean Square F Sig. (p) Observed Power (a) Knee Extensors Fatigue* 362.525 1 362.525 5.120 0.033 0.584 Fatigue Gender 6.582 1 6.582 0.093 0.763 0.060 Error (Fatigue) 1699.338 2470.806 Knee Flexors Fatigue* 736.509 1 736.509 20.596 < 0.001 0.992 Fatigue Gender 40.163 1 40.163 1.123 0.300 0.174 Error (Fatigue) 858.252 2435.761 Significant main effect or interaction (p<0.05)

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126 Table C-9. MANOVA table for la nding variables, touchdown angl e variables, extension ROM variables, and L/S and C/T kinetic vari ables in the study of knee joint muscles fatigue. Effect Roy's Largest Root F Hypothesis df Error df Sig. (p) Observed Power (a) Landing variables Gender* 1.642 4.104 6 15 0.012 0.88 Fatigue 0.335 0.838 6 15 0.559 0.24 Fatigue Gender 0.870 2.174 6 15 0.104 0.58 Touchdown angle variables Gender* 0.773 3.287 4 17 0.036 0.72 Fatigue 0.216 0.920 4 17 0.475 0.23 Fatigue Gender 0.110 0.467 4 17 0.759 0.13 Extension ROM variables Gender* 1.259 5.351 4 17 0.006 0.92 Fatigue 0.462 1.965 4 17 0.146 0.47 Fatigue Gender 0.556 2.363 4 17 0.094 0.55 L/S Kinetic variables Gender 0.644 2.062 5 16 0.124 0.529 Fatigue 0.748 2.395 5 16 0.084 0.602 Fatigue Gender* 1.010 3.234 5 16 0.033 0.752 C/T Kinetic variables Gender 0.336 1.074 5 16 0.411 0.285 Fatigue 0.476 1.522 5 16 0.238 0.399 Fatigue Gender* 0.936 2.995 5 16 0.043 0.714 Significant main effect or interaction (p<0.05)

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127 Table C-10. Univariate tests of between-subjects effects for the landing variables in the study of knee joint muscles fatigue. Measure Source Type III Sum of Squares df Mean Square F Sig. (p) Observed Power(a) PVGRF Gender 18.567 1 18.567 1.077 0.312 0.17 Error 344.679 2017.234 t(LP) Gender 0.051 1 0.051 1.111 0.304 0.17 Error 0.912 200.046 TD(KFA) Gender* 580.417 1 580.417 8.755 0.008* 0.80 Error 1325.875 2066.294 P(KFA) Gender 582.142 1 582.142 1.279 0.271 0.19 Error 9100.907 20455.045 TD(HFA) Gender 743.700 1 743.700 3.087 0.094 0.39 Error 4817.829 20240.891 P(HFA) Gender 20.202 1 20.202 0.062 0.807 0.06 Error 6561.283 20328.064 Significant main effect or interaction (p<0.05) Table C-11. Univariate tests of between-subjects effects for the touchdown angle variables in the study of knee joint muscles fatigue. Measure Source Type III Sum of Squares df Mean Square F Sig. (p) Observed Power (a) TD(L/S) Gender 83.277 1 83.277 0.535 0.473 0.11 Error 3111.027 20155.551 P(TL/L) Gender* 1201.868 1 1201.8689.729 0.005 0.84 Error 2470.578 20123.529 P(TH/TL) Gender 405.650 1 405.650 3.730 0.068 0.45 Error 2175.019 20108.751 P(LC/TH) Gender 19.032 1 19.032 0.054 0.819 0.06 Error 7090.599 20354.530 Significant main effect or interaction (p<0.05)

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128 Table C-12. Univariate tests of between-subjects effects for the extension ROM variables in the study of knee joint muscles fatigue. Measure Source Type III Sum of Squares df Mean Square F Sig. (p) Observed Power (a) P(L/S) Gender 2.401 1 2.401 3.567 0.074 0.44 Error 13.463 200.673 P(TL/L) Gender 18.888 1 18.888 3.295 0.085 0.41 Error 114.659 205.733 P(TH/TL) Gender* 328.901 1 328.901 7.437 0.013 0.74 Error 884.457 2044.223 P(LC/TH) Gender 17.388 1 17.388 0.084 0.776 0.06 Error 4162.524 20208.126 Significant main effect or interaction (p<0.05) Table C-13. Univariate tests of within-subjects ef fects for the kinetic vari ables of L/S junction in the study of knee joint muscles fatigue. Measure Source Type III Sum of Squares df Mean Square F Sig. (p) Observed Power (a) AxF(L/S) Fatigue 0.013 1 0.013 0.016 0.900 0.052 Fatigue Gender* 5.186 1 5.186 6.529 0.019 0.681 Error (Fatigue) 15.887 200.794 ShF(L/S)ant Fatigue 0.003 1 0.003 0.024 0.877 0.053 Fatigue Gender 0.209 1 0.209 1.884 0.185 0.258 Error (Fatigue) 2.213 200.111 ShF(L/S)post Fatigue 11.637 1 11.637 12.694 0.002 0.923 Fatigue Gender* 12.752 1 12.752 13.910 0.001 0.944 Error (Fatigue) 18.335 200.917 FlxM(L/S) Fatigue 0.025 1 0.025 0.323 0.576 0.084 Fatigue Gender 0.000 1 0.000 0.003 0.961 0.050 Error (Fatigue) 1.571 200.079 ExtM(L/S) Fatigue 1.047 1 1.047 8.029 0.01 0.769 Fatigue Gender* 1.036 1 1.036 7.943 0.011 0.765 Error (Fatigue) 2.609 200.130 Significant main effect or interaction (p<0.05)

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129 Table C-14. Univariate tests of between-subject s and within-subjects effects for the kinetic variables of C/T junction. Measure Source Type III Sum of Squares df Mean Square F Sig. (p) Observed Power (a) AxF(C/T) Fatigue 3.734 1 3.734 3.048 0.096 0.383 Fatigue Gender* 15.864 1 15.864 12.947 0.002 0.928 Error (Fatigue) 24.507 201.225 ShF(C/T)ant Fatigue 0.018 1 0.018 0.014 0.907 0.051 Fatigue Gender 0.174 1 0.174 0.138 0.715 0.064 Error (Fatigue) 25.305 201.265 ShF(C/T)post Fatigue 1.275 1 1.275 5.030 0.036 0.569 Fatigue Gender 0.039 1 0.039 0.154 0.699 0.066 Error (Fatigue) 5.068 200.253 FlxM(C/T) Fatigue 0.104 1 0.104 0.533 0.474 0.107 Fatigue Gender 0.015 1 0.015 0.075 0.787 0.058 Error (Fatigue) 3.898 200.195 ExtM(C/T) Fatigue 1.283 1 1.283 6.133 0.022 0.654 Fatigue Gender 0.874 1 0.874 4.175 0.054 0.494 Error (Fatigue) 4.185 200.209 Significant main effect or interaction (p<0.05)

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145 Wretling, M. L. & Henriksson-Larsen, K. ( 1998). Mechanical output and electromyographic parameters in males and female s during fatiguing knee-extensions. Int J Sports Med, 19 (6), 401-407. Wrigley, A. T., Albert, W. J., Deluzio, K. J., & Stevenson, J. M. (2005). Differentiating lifting technique between those who develop low back pain and those who do not. Clin Biomech (Bristol, Avon), 20 (3), 254-263. Yahia, L. H., Garzon, S., Strykowski, H., & Riva rd, C. H. (1990). Ultrastructure of the human interspinous ligament and ligamentum flavum. A preliminary study. Spine, 15 (4), 262268. Yang, K. H. & King, A. I. (1984). Mechanism of facet load transmission as a hypothesis for lowback pain. Spine, 9 (6), 557-565. Zhang, S. N., Bates, B. T., & Dufek, J. S. (2000) Contributions of lower extremity joints to energy dissipation during landings. Med Sci Sports Exerc, 32 (4), 812-819.

PAGE 146

146 BIOGRAPHICAL SKETCH Soo-An Park was born on January 19, 1965 to his parents Seung-Jae Park and Jung-Nim Kang in Seoul, Korea. He is married to H yunhee Kwon, and has one son, David Joonsuh. He also has one younger sister, Eun-Hee and one younge r brother, Soo-Min. Soo-An lived in Seoul, Korea throughout his childhood until he became an orthopaedic surgeon. He enrolled in the Medical College in the Catholic University of Kor ea, and worked as an intern and a resident of orthopaedic surgery at Kangnam St. Marys Hospital after graduation. During his residency, Soo-An was interested in the study of spine and sports medicine. His interest in the field of spinal surgery led hi m to pursue a Spine Research Fellowship under the guidance of Hansen A. Yuan, M.D. in S UNY-Upstate Medical University. A quest for interesting spine research and an interest in human performance in the sport led him to the Biomechanics Laboratory at the University of Florida and to doctoral work. New academic information in biomechanics and his clinical e xperience in orthopaedic spine surgery provided the foundation for his dissertation. Soo-An will continue his research work as he has accepted a 1-year postdoctoral fellowship in the Department of Orthopaedic Su rgery at SUNY-Upstate Me dical University. He will resume his clinical work in 2008 when he re turns to the Department of Orthopaedic Surgery at Asan Medical Center of Ulsan Medical University in Seoul, Korea, where he will specialize in spinal surgery.


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Title: Spinal Mechanics during Drop Landing: Effects of Gender, Fatigue and Landing Technique
Physical Description: Mixed Material
Copyright Date: 2008

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Table of Contents
    Title Page
        Page 1
        Page 2
    Dedication
        Page 3
    Acknowledgement
        Page 4
    Table of Contents
        Page 5
        Page 6
    List of Tables
        Page 7
        Page 8
    List of Figures
        Page 9
    Abstract
        Page 10
        Page 11
    Introduction
        Page 12
        Page 13
        Page 14
        Page 15
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    Materials and methods
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    Literature review
        Page 34
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    Results
        Page 77
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    Discussion
        Page 92
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    Conclusion
        Page 110
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    Future work
        Page 112
    Appendices
        Page 113
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    References
        Page 130
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    Biographical sketch
        Page 146
Full Text





SPINAL MECHANICS DURING DROP LANDING:
EFFECTS OF GENDER, FATIGUE AND LANDING TECHNIQUE





















By

SOO-AN PARK


A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL
OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT
OF THE REQUIREMENTS FOR THE DEGREE OF
DOCTOR OF PHILOSOPHY

UNIVERSITY OF FLORIDA

2006

































Copyright 2006

by

Soo-An Park



































To my parents









ACKNOWLEDGMENTS

Many people have earned my gratitude for their guidance and support during my doctoral

education and the completion of my dissertation. First I would like to thank my parents, Seung-

Jae Park and Jung-Nim Kang. They have supported me in every path with love I have taken and

they have provided me with the work ethic, value, and encouragement necessary to achieve my

goals. I would also like to express my gratitude to my wife, Hyunhee Kwon and my son,

Joonsuh, for their love and patience they have provided over the past several years. The

completion of my graduate education has been a joint endeavor and a shared achievement.

I would like to express thanks to my committee members who have challenged me to

become a better scientist and person. Dr. John Chow has exponentially strengthened my research

and dissertation and my abilities as a biomechanist. Dr. Mark Tillman has provided great support

and assistance in my research. Dr. Ronald Siders has provided support in my teaching during

graduation education, and Dr. Falsetti has assisted me in setting up the research hypotheses in

my dissertation.









TABLE OF CONTENTS



A C K N O W L E D G M E N T S ..............................................................................................................4

L IST O F T A B L E S ......................................................................................................... ........ .. 7

LIST OF FIGURES ............................................. .. .......... ............ ...............9

A B S T R A C T .......................................................................................................... ..................... 10

CHAPTER

1 INTRODUCTION .................................. .. ........... ..................................... 12

Purpose of the Study .................................. .. .......... ............................. 16
Significance of the Study ............. .. .................. .................. ............ ........ .... ............... 16
H y p oth eses ............................................................................................................. ....... .. 17
L im itatio n s ............................................................................................................. ........ .. 18

2 MATERIALS AND METHODS ...................................................................................20

Subjects .......................................................................................................... 20
Sam ple Size Justification ............. .. .................. .................. ............ ........ .... ............... 20
Experimental Setup ............................... .. ........... ............................... 21
T testing P protocol .................................................................................................... ........ .. 2 1
P re-F atigue L ending T rials ............................................................................ ................ 22
F atig u e P ro cedu re ............................................................................................................ 2 2
P ost-F atigue L ending Trials .................................................................... ................ 23
D ata R edu action .................................................................................................... ....... .. 23
D ata A n aly sis ......................................................................................................... ........ .. 2 5

3 L ITE R A TU R E R E V IE W .............. ..................................................................... 34

Biomechanical Properties of Spinal Structures .................................................................35
Intervertebral D isc ................................................................... ...... .. ... ...............3 5
V erteb ra ......................................................................................................... ....... .. 3 9
S p in al L ig am ents ................ .. ................................. ................................................ 4 0
Biomechanical Properties of Spinal Segments..................................................................41
Multisegmental Mechanics of the Spine ....................................................................41
R regional M echanics of the Spine ...................................... ...................... ................ 43
Biomechanical Performance of Spine In Vivo..................................................................47
T ru n k P o stu rin g ............................................................................................................... 4 7
W eig h t L iftin g ................................................................................................................ 4 8
S ittin g an d Stan din g ........................................................................................................ 50
W alk in g ......................................................................................................... ....... .. 5 1
R u n n in g ........................................................................................................ ....... .. 5 3









B iom echanical Etiology of Spinal Pain ............................................................. ................ 56
V ib ratio n ............................................................................... ............. .................. 5 7
L o rd o sis ................................................................................................. ..................... 5 9
T orsion ................................................... ..... .................... .......................6 0
Biomechanical Performance of Painful Spine..................... ...................................61
Landing B iom echanics .........................................................................................................63
Biomechanical Performance of Lower Extremity Joints during Landing.......................64
G ender D difference .................................................................................................... 66
L ending Stiffness........................................................................ ........................... ......... 67
Performance of Adapted Landing Biomechanics to Various Conditions .......................70

4 R E S U L T S ..................................................................................................... ..................... 7 7

Effects of L ending Technique .............................................................................................77
Effects of K nee Joint M uscles Fatigue ................................................................................79

5 D IS C U S S IO N .........................................................................................................................9 2

Effects of L ending Technique .............................................................................................92
Effects of Knee Joint M muscles Fatigue ..................... .........................................100

6 CONCLUSION .............................................................................. 110

E effects of L ending T technique ......................................................................... ...............110
Effects of Knee Joint M muscles Fatigue ..................... .........................................110

7 F U T U R E W O R K ..................................................................................................................1 12

APPENDIX

A IN F O R M E D C O N SE N T ......................................................................................................113

B P R E L IM IN A R Y ST U D Y .....................................................................................................118

C M AN OVA and AN OV A TABLES ....................... .......................................... ...............119

E effects of L ending T technique ......................................................................... ...............119
Effects of Knee Joint M muscles Fatigue ..................... .........................................125

L IS T O F R E F E R E N C E S .............................................................................................................130

B IO G R A P H IC A L SK E T C H .......................................................................................................146










6









LIST OF TABLES


Table page

2-1 Geisser-Greenhouse Correction D etail Report ............................................. ................ 27

2-2 M AN OV A table for the fatigue indices........................................................ ................ 27

2-3 Collapsed mean and SD values of fatigue indices before/after the fatigue procedure ......27

3-1 Average neutral zones for a functional spinal units in different regions of the spine
(0). .................................. ................ 73

3-2 Representative ranges of motion of CO-C 1-C2 complex (0)................ ............... 73

3-3 Representative ranges and limits of motion of the middle and lower cervical spines
(0). .................................. ................ 73

3-4 Normal active cervical ranges of motion (in vivo) reported in the literatures (0)..............73

3-5 Representative ranges and limits of motion of the thoracic spine (0).............................74

3-6 Representative ranges and limits of motion of the lumbar spine (0). ............................74

3-7 Comparison of lumbar compression loads in various trunk postures without external
lo ad in g ............................................................................................................. ....... .. 7 4

3-8 Average ranges of motion of the lumbar spine in normal walking and running in
different stu dies (0). ................................................................. ................................ ..... .. 7 5

3-9 Peak compression loads to the lower lumbar level during walking (x BW). ....................75

3-10 Intradiscal pressure of low lumbar level during various activities ...............................75

4-1 Collapsed mean and SD values of different landing variables for different genders
and landing techniques ................... ............. .............................. 82

4-2 Collapsed mean and SD values of touchdown angle and extension ROM of each
spinal region for different genders and landing techniques..........................................83

4-3 Collapsed mean and SD values of kinetic variables at L/S and C/T junctions for
different genders and landing techniques. ....................... ......................................... 84

4-4 Collapsed mean and SD values of different landing variables for different genders
an d fatig u e lev els. .............................................................................................................. 8 5

4-5 Collapsed mean and SD values of touchdown angle and extension ROM of each
spinal region for different genders and fatigue levels ............................. ..................... 86









4-6 Collapsed mean and SD values of kinetic variables at L/S and C/T junctions for
different genders and fatigue levels ..................................... ...................... ................ 87

B-1 Kinetic characteristics of L/S junction data for sample size justification......................118

C-1 MANOVA table for landing variables, touchdown angle variables, extension ROM
variables, and L/S and C/T kinetic variables in the study of landing technique effects. .119

C-2 Univariate tests for the different landing variables in the study of landing technique
effects ...................................................................................................... 120

C-3 Univariate tests for the touchdown angle variables in the study of landing technique
effects ........................................................................................... ........ .. 12 1

C-4 Univariate tests for the extension ROM variables in the study of landing technique
effects ...................................................................................................... 122

C-5 Univariate tests for the L/S kinetic variables in the study of landing technique
effects ...................................................................................................... 123

C-6 Univariate tests for the C/T kinetic variables in the study of landing technique
effects ...................................................................................................... 124

C-7 MANOVA table for the fatigue indices in the study of knee joint muscles fatigue........125

C-8 Univariate tests of within-subjects effects for the fatigue indices in the study of knee
joint m u scles fatigu e ............................................................................................... 12 5

C-9 MANOVA table for landing variables, touchdown angle variables, extension ROM
variables, and L/S and C/T kinetic variables in the study of knee joint muscles
fatig u e ................................................................................................... .................... 12 6

C-10 Univariate tests of between-subjects effects for the landing variables in the study of
knee joint m uscles fatigue ........................................................................... ............... 127

C-11 Univariate tests of between-subjects effects for the touchdown angle variables in the
study of knee joint m muscles fatigue....................................................... ............... 127

C-12 Univariate tests of between-subjects effects for the extension ROM variables in the
study of knee joint m muscles fatigue....................................................... ............... 128

C-13 Univariate tests of within-subjects effects for the kinetic variables of L/S junction in
the study of knee joint muscles fatigue ...... .......... ....... ...................... 128

C-14 Univariate tests of between-subjects and within-subjects effects for the kinetic
variables of C /T junction. .................... ................................................................ 129









LIST OF FIGURES


Figure page

2-1 Correlations of power and sample size for each combination of variables ....................28

2-2 E xperim ental setup .................................................... ................................................ 29

2-3 A subject w ith m arkers on. ................................................. .............. ................ 30

2-4 Overview of the experim ental procedures ................................................... ................ 31

2-5 Marker placement (left) and definition of regional angles of the spine (right). ..............32

2-6 Kinematic variables defined by the critical instants identified from kinematic and
fo rc e p late d a ta ................................................................................................................. ... 3 3

3-1 The load-displacement curve of a functional spinal unit (FSU) is generally nonlinear
and biphasic [neutral zone (NZ) and elastic zone (EZ)] ..............................................76

4-1 Significant interactions of the touchdown angle and C/T kinetic variables between
gender and landing technique ................................................................. ................ 88

4-2 Representative kinem atics of each spinal region ......................................... ................ 89

4-3 Representative kinem atics of each spinal region ......................................... ................ 90

4-4 Significant interactions of the kinetic variables at L/S and C/T junctions between
gender and fatigue level ..................... ............ .....................................91









Abstract of Dissertation Presented to the Graduate School
of the University of Florida in Partial Fulfillment of the
Requirements for the Degree of Doctor of Philosophy

SPINAL MECHANICS DURING DROP LANDING:
EFFECTS OF GENDER, FATIGUE AND LANDING TECHNIQUE

By

Soo-An Park

December 2006

Chair: John W. Chow
Major: Health and Human Performance

Objective: To investigate the kinematics of the spinal column and the kinetics of

lumbosacral (L/S) and cervicothoracic (C/T) junctions during the drop landing, and to evaluate

the effects of gender, landing technique, and fatigue.

Methods: Thirteen male and 13 female healthy young volunteers were tested. To track the

kinematics of different spinal regions, surface markers were placed on skin over selected spinous

processes. Data were collected using a 3-D motion capture system and a forceplate. The subject

performed 3 drop landings using his/her own landing technique (NL) and 3 soft landings with

instruction (SL; SL1). During each trial, the subject descended from a 50-cm height platform and

landed on a forceplate with the left foot at the center of forceplate. After completing isokinetic

knee flexion/extension exercises and a 30-minute run on a motorized treadmill, the subject

performed 3 more soft landings in a knee joint muscle fatigue state (SL2). Kinematic variables

included touchdown angle and initial extension range of motion of different spinal regions.

Kinetic variables included joint resultants at L/S and C/T junctions computed using an inverse

dynamics approach. Multivariate analyses of variance (MANOVA) and follow-up univariate

analyses of variance (ANOVA) were used to examine the effects of gender, landing technique,

and fatigue on different kinematic and kinetic variables.









Results: Females exhibited a significantly extended thoracolumbar landing posture and

greater thoracic motion than males in all 3 experimental conditions. Thoracic and lower cervical

extension motions increased and most joint resultants decreased significantly when going from

NL to SL. Posterior shear force in males and anterior shear force in females were significantly

greater than their counterparts at C/T junction during NL. Females exhibited significantly

increased joint resultants from SL1 to SL2, while males did not.

Conclusion: The spinal column is more actively involved in energy absorption during

drop landings in males, and the thoracolumbar region could be more loaded by hyperextension

during soft landing in females comparing to males. Repeated drop landings may cause injuries to

the cervical spine by different mechanisms in each gender. For females, soft landings under

fatigue condition can be a risk factor of spinal injury.









CHAPTER 1
INTRODUCTION

Spinal function in vivo has been generally determined by radiographs, showing the

kinematics of each spinal segment and spinal region (Wong et al., 2006) or by inclinometer

measurements (Ng et al., 2001). Most previous investigations about spinal kinematics were

based on postures of the spinal column in static conditions and static spinal kinematics. Active

spinal motion can be identified as the end range of motion (ROM) achieved by the subjects, and

the passive motion is the end ROM obtained by applying external forces to a fully motioned

spine. Both active and passive ROM of the spine have been used to interpret the functional status

of the spine in clinical and laboratory studies (Dvorak et al., 1988).

Several researches have evaluated the kinematics of the spine in dynamic situations like

walking (Callaghan et al., 1999; Crosbie et al., 1997) and running (Schache et al., 2002).

Although results from both walking and running analyses may have clinical implications, these

two locomotive tasks may not be the most adequate tasks to reveal spinal function in dynamic

situation.

The age of patients who develop spinal degeneration and undergo spinal surgeries is

getting younger and this population is getting larger (Kjaer et al., 2005). The prevalence of spinal

pain and degeneration is higher in active individuals (Bono, 2004). The biomechanics of the

spine in a static condition are quite different from those in a dynamic condition. Demands for

developing a functional evaluation of the spine in vivo have increased, because conventional

radiographic study of spinal ROM is not good enough to estimate the spinal function of every

subject. Also with the development of the surgical technique of spinal arthroplasty, the

restoration of the original function of each patient has become a primary goal of the surgery.









Therefore, normal spinal mechanics in vivo in various tasks and the mechanical factors

contributing to spinal pain and degeneration are useful information to medical practitioners.

It has been suspected that most disabling and chronic back pain arises from intervertebral

disc degeneration. With disc degeneration, biologic as well as biomechanical changes follow.

Disc degeneration occurs most commonly in the third to fifth decades of life. Aging causes

definite changes in the morphology and composition of spinal tissues that are mostly unrelated to

pain. However, many studies reported that aging weakens the intervertebral disc tissues due to

decreased cell number, high apoptosis rate, and the different response to biologic environments.

Impaired function of the intervertebral discs may make people more vulnerable to mechanical

injuries, which can initiate further structural and symptomatic disc degeneration.

Nonphysiologic loading to spinal structures can contribute to intervertebral disc

degeneration. With the increased loading to the trunk, spinal shrinkage was found double to the

unloaded condition (Fowler et al., 1994). A flexion posture significantly increases extensor

muscle activity when compared with a standing neutral posture (Arjmand & Shirazi-Adl, 2006).

Forward leaning of the trunk causes the vertebrae in anterior translation, and disc loads and

stresses were significantly increased most markedly at the L5/S1 level (Harrison et al., 2005).

When the torso is fully flexed during repetitive lifting tasks, fatigue failure of spinal tissues can

occur rapidly (Wrigley et al., 2005). On the other hand, abnormally low loading causes atrophy

in muscle, cartilage and bone, leaving them less able to resist high loads (White & Panjabi,

1990).

Abnormalities in the lower extremities can affect spinal mechanics. For example, patients

who have leg length discrepancy due to lower extremity disorders demonstrate different spinal

kinematics during gait when compared with normal subjects suggesting greater risk of









developing spinal disorders (Kakushima et al., 2003). However, a walking task may not be

sensitive enough to detect the effects of altered mechanics in the lower extremity upon spinal

mechanics.

Gender differences in spinal kinematics during trunk posturing hinted that there may be

different mechanical causes of spinal degeneration in different genders. During a prolonged

sitting, males exhibit more flexion of lumbar spine and trunk than females. Males and females

may be exposed to different loading patterns during certain prolonged postures and can develop

different injuries or degeneration mechanisms of the spine (Dunk & Callaghan, 2005).

Various in vivo and in vitro biomechanical techniques have been developed to investigate

spinal mechanics, but they all have limitations. Physical properties of the human spine may be

obtained from studies of living subjects, whole cadavers, isolated whole cadaveric spines, and

isolated spinal segments. A living subject provides realistic but less accurate measurements. An

isolated spinal segment lacks muscles, but can provide accurate data and allows the possibility of

studying the effects due to trauma and surgical stabilizations (White & Panjabi, 1990).

The force applied to the spine depends on body weight (Rodacki et al., 2005), external

loads (Lawrence et al., 2005), and internal muscle forces (Arjmand & Shirazi-Adl, 2006) which

can be varied during dynamic activities (Chow et al., 2003; Tully et al., 2005). Direct

measurement of spinal compression force pioneered by Nachemson (1966; 1964), obtained by

inserting a pressure needle into the lumbar intervertebral discs of living subjects, has been

benchmarked and compared to many other studies that measured forces applied to the spine in

vivo (Ledet et al., 2005; Sato et al., 1999). However, most researchers who have investigated

spinal mechanics utilized indirect techniques on living subjects and cadaveric spinal segments

because of the invasiveness and limited localization of direct measurements.









Normal spinal mechanics during activities of living subjects and each spinal segment as

functional spinal unit (FSU) and spinal column are usually studied using indirect measurements.

Spinal mechanics during various movements walking, forward/backward bending, sit-to-

stand/stand-to-sit, etc. have been investigated using image-based motion analysis systems.

Using surface electromyography (EMG) techniques, spinal muscle activity during different

spinal movements has been investigated. McGill (1992) estimated the moments generated by

trunk muscles using an EMG-driven musculoskeletal model during trunk posturing movements.

By applying inverse dynamic techniques to a rigid segment model, joint resultants at different

lower extremity joints and spinal motion segments can also be calculated. However, many spinal

vertebrae included in one trunk segment may mimic the joint resultant values at the trunk

segment, and the kinematics of the spinal region in vivo could not be accessed in details (Khoo et

al., 1995).

Using forceplate and kinematic data, inverse dynamic techniques are commonly used to

calculate joint resultants at lower extremity joints during various activities (Kernozek et al.,

2005). Jumping and landing were commonly adopted for measuring joint resultants at lower

extremity joints simulating active and vigorous movements. A recent study completed by the

author indicated that limiting trunk movements caused changes in the mechanics of lower

extremity joints during drop landings (Park et al., 2006b). This finding suggests that body

segments proximal to the hip joints could be involved in regulating the force transmitted to the

lower extremity joints during the landing phase. Conversely, different mechanical configurations

of lower extremity joints may affect spinal mechanics during landing and the findings from

spinal mechanics for each specific configuration may provide insights into the spinal function in

dynamic situations.









Purpose of the Study

The purpose of this study was to investigate the kinematics of different regions of the

spinal column and the kinetics at the lumbosacral (L/S) and cervicothoracic (C/T) junctions

during drop landings. Specifically, this study evaluated the effects of gender and landing

techniques (self-selected landing and instructed soft landing techniques) on spinal mechanics

during drop landing in the first part of this study. In the second part of this study, the effects of

gender and muscular fatigue of knee flexor/extensor on spinal mechanics during drop landings

were examined.

Significance of the Study

Previous researchers have evaluated the mechanical causes of spinal degeneration in vivo

and in vitro while simulating various activities using direct and indirect measurement techniques.

Most activities employed in these biomechanical studies were simple, everyday activities that do

not demand much spinal movement. However, the populations who have spinal degeneration and

undergo spinal surgeries are getting larger and younger. These individuals want to be physically

active and participate in activities that may demand vigorous spinal movements. Drop landings

have been widely used to examine coordination and mechanical stress at different joints under

dynamic situations. However, most landing studies were confined to the biomechanics of lower

extremity joints and very few studies evaluated mechanics of upper body movements during

landings.

In addition to lower extremity joints, this study attempted to identify mechanical

characteristics of different spinal regions during drop landings. The results might provide insight

into the effects of gender, landing technique and fatigue of knee joint muscles on spinal

mechanics during drop landings.









Hypotheses

In the absence of extensive pilot data and relevant data reported in the literature, the

following questions were raised and null hypotheses were tested to address the aims of this

study:

* To determine the effects of gender and landing technique on spinal column kinematics and

loads to the lumbosacral (L/S) and cervicothoracic (C/T) junctions during drop landings.

QI: Would there be significant differences in spinal column kinematics and loads to the L/S

and C/T junctions during drop landings between males and females regardless of landing

technique?

la. There would be no significant differences in motion characteristics of the spinal

column between males and females.

lb. There would be no significant differences in peak resultant forces and moments

transmitted to the L/S and C/T junctions between males and females.

Q2: Would the self-selected and soft landing techniques cause significant differences in

spinal column kinematics and loads to the L/S and C/T junctions during drop landings?

2a. The conditions for landing technique would not cause significant differences in

motion characteristics of the spinal column.

2b. The conditions for landing technique would not cause significant differences in peak

resultant forces and moments transmitted to the L/S and C/T junctions.

* To determine the effects of gender and fatigue of knee joint muscles on spinal column

kinematics and loads to the L/S and C/T junctions during drop landings using a soft landing

technique.









Q3: Would there be significant differences in spinal column kinematics and loads to the L/S

and C/T junctions during drop landings using a soft landing technique between males

and females regardless of fatigue condition?

3a. There would be no significant differences in motion characteristics of the spinal

column between males and females.

3b. There would be no significant differences in peak resultant forces and moments

transmitted to the L/S and C/T junctions between males and females.

Q4: Would fatigue of knee joint muscles cause significant differences in spinal column

kinematics and loads to the L/S and C/T junctions during drop landings using a soft

landing technique?

4a. Fatigue of knee joint muscles would not cause differences in motion characteristics of

the spinal column.

4b. Fatigue of knee joint muscles would not cause differences in peak resultant forces and

moments transmitted to the L/S and C/T junctions.

Limitations

* Measurement errors of forceplate and digital video cameras are always present but they are

considered acceptable within the specifications of the manufacturers.

* Marker placement was controlled cautiously to minimize errors.

* Sagittal spinal kinematics relative to the adjacent spinal region based on spinal marker

locations would have some errors due to skin movement, but will be considered

acceptable.









* The center of L/S junction was assumed to be located at the midpoint between both

posterior iliac crest markers at L5/S1 level. The center of the C/T junction was assumed to

be located at the midpoint between the two acromial process markers.

* Mechanical characteristics of the lower extremities were assumed to be symmetrical and

only the data collected from the left leg were used to calculate joint resultants at the L/S

and C/T junctions.









CHAPTER 2
MATERIALS AND METHODS

Subjects

Thirteen male (age: 21.4 1.3 yrs, mass: 74.2 10.2 kg, height: 174.8 5.2 cm) and 13

female (age: 21.1 1.3 yrs, mass: 58.6 6.8 kg, height: 165.5 5.3 cm) healthy and physically

active individuals participated in this study. They were free from any cardio-respiratory diseases

that would prevent them from completing the fatigue procedures and musculoskeletal diseases or

injuries that could influence spinal and lower extremity mechanics. Before testing, each subject

carefully read and signed a written informed consent approved by the Institutional Review Board

of the University of Florida (Appendix A).

Sample Size Justification

To simplify the calculation of sample size determination, data of a dependent variable

(peak extensor moment at the lumbosacral junction) from a pilot study were used to set up a 2 x

2 [gender x landing technique: normal landing and soft landing] ANOVA with repeated

measures on the last factor. Based on the preliminary data collected from one male and one

female, the means of male and female data were 2.74 and 2.58 N.m.kg-'.BH-1, respectively, and

those for NL and SL were 3.05 and 2.27 N.m.kg-'.BH-1, respectively (Note. BH: body height).

There was a 39% difference between male and female, and a 26% difference between normal

and soft landing conditions. The mean standard deviation value was 0.3 (Table B-1 in Appendix

B).

Based on these values, the current study was designed to detect at least 12% changes in

peak extensor moment in between-subject group and 10% difference in within-subjects groups

with alpha = 0.05 and beta = 0.2 (80% power). A Geisser-Greenhouse correction report indicated

that power values were over 80% for all the terms with a sample size of n= 13 (Table 2-1, Figure









2-1). Therefore, at least 12 subjects were required for each within-subject group to determine the

effects of the treatment. All the calculations were performed using PASS 2005 (Number

Cruncher Statistical Systems, Kaysville, Utah).

Experimental Setup

A forceplate (Type 4060-10, Bertec Corporation, Columbus, OH) operating at 1,200 Hz

was set up at the center of the Biomechanics Research Laboratory (Figure 2-2). A 50-cm height

platform was placed behind and slightly to the right of the forceplate. Seven Hawk digital

cameras (Motion Analysis Corp., Santa Rosa, CA) were stationed around the forceplate to collect

kinematic data and were 3-4 m from the landing area. Kinematic data were captured at a

frequency of 100 Hz. The system was calibrated prior to each testing session according to the

procedures specified by the manufacturer.

Testing Protocol

To expose the lower extremity and the back of the trunk, subjects were asked to wear only

short pants (both males and females) and sports bra (for females only) (Figure 2-3). They wore

their own sports shoes during testing. Upon completing the measurements of body weight and

height, each subject jogged on a treadmill and stretched with self-selected exercises for 10

minutes as a warm-up (Figure 2-4).

Reflective markers (1.0 cm in diameter) were placed on the left second metatarsal head,

dorsal navicular surface, heel, lateral malleolus, lower-shank, mid-shank, lateral tibial

epicondyle, lower-thigh, mid-thigh, greater trochanter, anterior superior iliac spines, 2nd spinous

process on median sacral crest of sacrum to track the locations of the left lower extremity and

pelvis (Kadaba et al., 1990; Kadaba et al., 1989). Another set of reflective markers was applied

over the subject's spinous processes for measuring the kinematics of the spinal column (C4, C6,

Tl, T3, T6, T10, T12, L2, L4, both posterior iliac crests at L5/S1 level) (Figure 2-5).









After marker placement, drop landing practice trials were provided to ensure consistent

landing during experimental trials. In each trial, the subject descended from a 50 cm-height

platform and landed on a forceplate with his/her left foot at the center of the forceplate and right

foot on a wooden platform of the same height.

Pre-Fatigue Landing Trials

After practice drop landings, the subject completed 3 trials of drop landing using his/her

own landing technique (normal landing: NL). The subject was then instructed on how to perform

a drop landing using the soft landing technique (soft landing: SL). He/She was instructed to try

soft landing by touching the balls of feet on the ground at initial impact, delaying heel contact,

and using more knee flexion after the landing. Three trials of soft landing were collected (SL1).

The averages over the three trials for each landing technique were used in subsequent analyses.

In each trial, kinematic and GRF data were sampled for 4 s. If the subject did not maintain

balance after landing, that trial was discarded and repeated.

Fatigue Procedure

The subject was asked to sit on the chair of a KinCom dynamometer (Chattanooga Group,

Inc., Hixson, TN) and perform 30 repetitions of isokinetic reciprocal knee flexion/extension with

full ROM and maximal effort at 600/s to induce muscular fatigue of knee joint flexor/extensor.

The exercise was repeated at the speed of 1800/s to measure the isokinetic strength for the

purpose of quantifying the fatigue level. After the isokinetic exercises, the subject ran on a

motorized treadmill at 4-6 mph for 30 minutes. The running intensity was lower than the typical

daily exercise for developing and maintaining fitness (Fletcher et al., 2001; Pollock et al., 1998).

Also, the speed and duration of running used in this study was known not to elicit cardio-

respiratory fatigue (Hardin et al., 2004). If the subject became exhausted before the end of the 30









minutes run, the running speed was reduced to 4.0 mph so that he/she could complete the run.

Isokinetic knee flexion/extension exercises were repeated immediately after running to quantify

the fatigue state of knee joint muscles.

Post-Fatigue Landing Trials

Immediately after the fatigue procedure, the subject performed 3 trials of drop landing

using the soft landing technique from the same platform (soft landing performed under fatigue

condition: SL2). The averages over the 3 trials were used in subsequent analyses.

Data Reduction

Kinematic data were processed using EVaRT 4.6 software (Motion Analysis Corp., Santa

Rosa, CA). The animation of reflective markers in each trial was examined qualitatively by the

investigator. Positional data were smoothed using a Butterworth low pass filter with a cutoff

frequency of 10 Hz. Three-dimensional kinematic and kinetic data for the left ankle, knee, and

hip joints, pelvis and trunk, and the kinematic data of spinal column markers were calculated

using Kintrak 6.2 software (Motion Analysis Corp., Santa Rosa, CA). Locations of spinal

column markers were used to define 5 spinal regions: lower cervical (LC), thoracic (TH),

thoracolumbar (TL), lumbar (L) and sacral (S) regions (Figure 2-5).

Joint resultants at the left knee and hip joints and L/S and C/T junctions were computed

based on the kinematic and forceplate data using an inverse dynamics approach. Assuming

symmetry in lower extremity mechanics, mechanical characteristics of the right leg were the

same at the left leg for the purpose of computing joint resultants at the L/S and C/T junctions. To

minimize the variation due to individual differences in physique, force variables were

normalized to the subject's body mass, and moment variables were normalized to body mass and

body height.









Muscle fatigue was quantified by the fatigue index. Fatigue index was determined by the

decline in peak torque in 30 repetitions, and calculated by the following formula to yield a

percent decrease for each isokinetic torque value:

Fatigue index = 100 [(last 5 repetitions/highest consecutive 5 repetitions) x 100]

For each subject, the highest consecutive five repetitions were determined by the values

attained from the two repetitions immediately prior to, and following, the single highest

repetition value. If the single highest repetition value was observed within the first 3 repetitions,

the first 5 repetitions were used to calculate the fatigue index (Pincivero et al., 2003).

Fatigue levels of knee joint muscles before/after the fatigue procedure were evaluated with

the fatigue indices of knee flexor and extensor muscles. Repeated measures MANOVA revealed

a significant main effect of knee joint muscles fatigue for the fatigue index variables (p=0.001),

but did not reveal any significant main effect of gender (p=0.528) and interaction between

gender and fatigue level (p=0.589) (Table 2-2). The univariate contrast procedures indicated that

the fatigue indices of both knee flexors and extensors increased significantly by the fatigue

procedure (Table 2-3). Only those subjects who demonstrated increased fatigue indices in both

knee flexors and extensors (12 males and 10 females) were included in subsequent analyses (SL1

vls. SL2).

The landing phase was defined using the critical instants identified from the kinematic and

GRF data (Figure 2-6), and the critical instants are as follows:

* The instant when the vertical ground reaction force (VGRF) starts to increase, the initial
touchdown (the beginning of landing phase), was identified from GRF data.

* Maximal knee joint flexion after initial touchdown was identified from the kinematic data
(the end of landing phase).

At the completion of data reduction, the dependent variables were divided into five groups:









* Landing variables: peak vertical GRF (PVGRF), time for landing phase (t(LP)), knee flexion
angle at touchdown (OTD(KFA)), ROM of knee flexion from touchdown to initial peak of
knee flexion (OP(KFA)), hip flexion angle at touchdown (OTD(HFA)), ROM of hip flexion from
touchdown to initial peak of hip flexion (OP(HFA)).

* Touchdown angles: lumbar regional angle at touchdown (YTD(L/S)), thoracolumbar regional
angle at touchdown (YTD(TI i ,), thoracic regional angle at touchdown (YTD(TH/TL)), lower
cervical regional angle at touchdown (YTD(LC/TH)).

* Extension ROMs: extension ROM of lumbar region from touchdown to initial peak during
landing phase (YP(L/S)), extension ROM of thoracolumbar region from touchdown to initial
peak during landing phase (YP(T I ,), extension ROM of thoracic region from touchdown to
initial peak during landing phase (YP(TH/TL)), extension ROM of lower cervical region from
touchdown to initial peak during landing phase (YP(LC/TH)).

* Kinetic variables at L/S junction: peak axial compressive force [AxF(L/S)], peak anterior
shear force [ShF(L/S)ant], peak posterior shear force [ShF(L/S)post], peak flexor moment
[FlxM(L/S)], peak extensor moment [ExtM(L/S)] after touchdown.

* Kinetic variables at C/T junction: peak axial compressive force [AxF(C/T)], peak anterior
shear force [ShF(C/T)ant], peak posterior shear force [ShF(C/T)post], peak flexor moment
[FlxM(C/T)], peak extensor moment [ExtM(C/T)] after touchdown.

OKFA was defined as the angle between the line of shank axis and thigh axis. OHFA was

defined as the angle between the line of thigh axis and pelvis axis. The negative angles of

OTD(KFA) and OTD(HFA) mean the flexions of knee and hip joints at the touchdown. For OP(KFA) and

OP(HFA), absolute values were used.

A regional angle of the spine was defined as the angle between the lines representing a

spinal region and its lower adjacent region. A positive touchdown angle indicates the spinal

region is in an extended state or extension motion relative to the lower adjacent region and a

negative angle indicates the spinal region is in a flexed state or flexion motion (Figure 2-5).

Data Analysis

For the non-fatigued data (NL and SL data), the 5 groups of dependent variables were

submitted to five separate 2 x 2 (Gender x Landing type) MANOVA with repeated measures on

the last factor. For the soft landing data (SL1 and SL2 data), the 5 groups of dependent variables









were submitted to 5 separate 2 x 2 (Gender x Fatigue level) MANOVA with repeated measures

on the last factor. Follow-up univariate analyses were conducted when appropriate. Bonferroni

adjustments were used during follow-up testing. A priori alpha level was set at 0.05 for all

statistical procedures. All statistical tests were performed using SPSS 13.0 for Windows (SPSS

Inc., Chicago, IL).









Table 2-1. Geisser-Greenhouse Correction Detail Report.
Term (levels) Power Alpha F Lambda dfl|df2 Epsilon E (Epsilon) GI
n= 13 N =26 Means x l
Gender (B: 2) 0.8310 0.05 4.26 9.25 1124 1 1 0
Landing (W: 2) 1 0.05 4.26 43.33 1124 1 1 0
BW 1 0.05 4.26 43.33 1124 1 1 0

Table 2-2. MANOVA table for the fatigue indices.
Roy's Largest Hypothesis Error Observed
Effect Root F df df Sig.(p) Power (a)
Gender 0.057 0.657 2 23 0.528 0.147
Fatigue* 0.875 10.067 2 23 0.001 0.971
Fatigue x Gender 0.047 0.542 2 23 0.589 0.129
Significant main effect or interaction (p<0.05)

Table 2-3. Collapsed mean and SD values of fatigue indices before/after the fatigue procedure.
Before After Fatigue x
Fatigue indices Mean (SD) Mean (SD) Gender: p Fatigue: p Gender: p
Knee extensors
(%) 14.9 (9.0) 20.2 (11.2) 0.253 (0.2) 0.033 (0.58)* 0.763 (0.06)
M 16.4 (9.3) 22.4 (8.2)
F 13.4(8.8) 17.9(13.4)
Knee flexors
(%) 10.6 (8.8) 18.1 (10.7) 0.887 (0.05) <0.001 (1.0)* 0.3 (0.17)
M 9.5 (9.0) 18.8 (13.1)
F 11.7(8.7) 17.5 (8.1)
* Significant main effect or interaction (p<0.05)











Power vs n by Terms with K=1.00 GG F Test


B B


0.9+-


0.8--

0.7-


0.6--

0.4--

0.3--

0.2--

0.1--


9 11 12 13 15 17 19


Figure 2-1. Correlations of power and sample size for each combination of variables. Note: B
(between-subject variable), W (within-subject variable), BW (interaction between B
and W).


A BW


1.0-F r, [ n R M









































Figure 2-2. Experimental setup.















































Figure 2-3. A subject with markers on.








Subject Consent


Pre-fatigue


Fatigue Protocol


Post-fatigue


Demographic Data


Warm-up
*-- Subject Preparation

Practice Drop Landings


Drop landing: Normal Landing (NL) Trials


Instructions of Soft Landings


Drop landing: Soft Landing (SL; SL1) Trials


Isokinetic Exercise


30-min Run


Isokinetic Exercise


Drop landing: Soft Landing (SL2) Trials


Figure 2-4. Overview of the experimental procedures.












Surface Back Markers
M,: C4
M,: C6
M3: T1
M4:T3
M,: T6
Me: T10
M7: T12
NAI'), "


Mg: L4
J M: Lt. Post. Iliac Crest at L5/S1
M,,: Rt. Post. Iliac Crest at L5/S1
M12: S2
P,: Midpoint between MIo and M1l


Definition for motion of
spinal region relative to
the lower adjacent region


'Y(LCTrH)


S Mel
' M.
M P1
8 f


M 1


Marker placement (left) and definition of regional angles of the spine (right).


M '
M3
M4
M


Spinal region
Lower cervical region (LC): M2MI
Thoracic region (TH): M5M3
Thoracolumbar region (TL): M7Me
Lumbar region (L): MOM8
Sacral region (S): M12P1

Regional angle of the spine
Lower cervical angle (Y(LCTH): M2M1 / MM3
Thoracic angle ((THTL): M5M3 L M7M6
Thoracolumbar angle (Y(TUL): M7M6 Z MOPs
Lower lumbar angle (v(us)): MOMS M12P1


Figure 2-5.














Angle ()


Forceplate











Critical instants


Kinematic variables


Touchdown of the foot to
the ground
Extension peak of spinal region
I during landing phase (LP)


C Spinal regional angle


UJ
a ----^.A -yT


r '*-'*


LP ,


Maximal
SKnee flexior,'


'YTD(LC/TH) "fP(LCfT


I L IIT ILI
'YTD(TLL)
YTD(L/S)
Touchdown angle of
spinal region


Knee flexion angle


YP(THITL)
TP(TLAL)
YP(L/S)
Extension ROM of
spinal region


Figure 2-6. Kinematic variables defined by the critical instants identified from kinematic and
forceplate data.


Time (s)









CHAPTER 3
LITERATURE REVIEW

The purpose of this study was to investigate the mechanical characteristics of the spine

during drop landings using different landing techniques and fatigue status of lower extremity

muscles. A biomechanical model of the lower extremity and spine was employed to study the

kinetics and kinematics of the spine and lower extremity joints. Research investigating spinal

mechanics using in vitro and in vivo techniques is quite extensive while studies of spinal

mechanics in jumping and landing are quite limited. Previous in vitro research focused on range

of motion of spinal segments in various conditions. Previous in vivo research focused on the

spinal mechanics during common, everyday activities without considering lower extremity

activities. More specific knowledge about spinal mechanics during vigorous physical activities is

necessary to have a better understanding of spinal mechanics when the spine is under dynamic

loadings. A description of spinal mechanics during drop landings is not currently available in the

literature.

Early studies to investigate spine biomechanics were mostly focused on defining

mechanical properties of spinal structures and segments in various conditions using different

instrumentations. Recent in vitro research focuses on the cause-and-effect relationship and

mathematical modeling of specific situations with cadaveric spinal segments using

computational programming and statistical procedure to explain the complicated situation and to

develop the spinal instrumentation by minimizing and simplifying the mechanical conditions.

Recent in vivo research focuses on estimating and verifying the results from in vitro and clinical

studies using optoelectrical systems, EMG, forceplate and so on.

To assist the reader in understanding the current issues of research on spinal mechanics,

the literature review will be presented under the following headings: biomechanical properties of









spinal structures, biomechanical properties of spinal segments, biomechanical performance of

spine in vivo, biomechanical etiology of spinal pain, biomechanical performance of painful

spine, and landing biomechanics.

Biomechanical Properties of Spinal Structures

The mechanical properties of the bony structures (including facet joints), ligaments, spinal

muscles and intervertebral disc of each spinal level have been thoroughly identified in previous

studies. The intervertebral disc has received more attention than the other spinal structures due to

its specific anatomical and biomechanical features. Defining disc characteristics is the first step

in reviewing the mechanical properties of the spine.

Intervertebral Disc

An intervertebral disc consists of three components: the nucleus pulposus, the annulus

fibrosus, and the cartilaginous end-plate. However, there is no clear landmark to differentiate the

nucleus pulposus and annulus fibrosus because the peripheral region of the nucleus pulposus

merge with the inner region of the annulus fibrosus. The nucleus pulposus is a centrally located

mucoid material in semi-fluid state, and its water content of which ranges from 70-90%

(Panagiotacopulos et al., 1987). The annulus fibrosus forms the outer region of the disc, and

consists of collagen fibers in a highly ordered pattern. The collagen fibers are layered in 10-20

sheets called lamellae and arranged in a helicoid manner. They run in the same direction in a

given lamella but in opposite directions in adjacent lamella. The lamellae are thick in anterior

and lateral regions of the annulus, and thin posteriorly (Inoue, 1981). The end-plate is composed

of hyaline cartilage about 0.6-1.0 mm thick which separates the other two components of the disc

from the vertebral body (Roberts et al., 1989). The end-plate covers the entire nucleus pulposus,

but does not cover the entire annulus fibrosus peripherally. Instead, the ring apophysis, which is

part of a vertebral body, covers the peripheral region of the annulus fibrosus. Because of the









attachment of the annulus fibrosus to the end-plate, the end-plate is strongly bound to the disc,

but weakly attached to the vertebral body (Inoue, 1981).

The basic functions of the disc are to transmit loads from one vertebral body to the next

and to allow movement between vertebral bodies. All components of the disc are involved in

weight-bearing. When an axial compressive load is applied to a nucleus, the nucleus tends to

reduce the height and expand radially towards the annulus fibrosus. This radial expansion exerts

a pressure on the annulus which tends to stretch its collagen lamellae outwards. However, the

tensile properties of the collagen resist this stretch, and the lamellae oppose the outward pressure

exerted by the nucleus. Application of a 400 N load to an intervertebral disc causes only 1 mm of

vertical compression and only 0.5 mm of radial expansion of the disc (Hirsch & Nachemson,

1954). The nucleus pressure is also towards the end-plates, and constrained by the end-plates and

vertebral bodies. The pressure on the end-plates serves to transmit the part of applied load from

one vertebra to the next, and the radial pressure on the annulus fibrosus braces it and prevents the

annulus from buckling (Roaf, 1960b).

Brown and colleagues (1957) conducted static tests to compare the relative strength of

the disc with that of the vertebral body in compressive loads, without the posterior elements.

They found the first structure to fail in such a construct was the vertebra, instead of the

intervertebral disc, because of the fracture of the end-plates. They also observed no difference

between the vertebrae with normal discs and those with degenerated discs. The mode of failure

by the pure compressive loads seemed to be mostly dependent on the condition of the vertebral

body (osteoporosis of the vertebrae), not on the condition of the disc.

During distraction, all points on one vertebral body move an equal distance

perpendicularly from the upper surface of the other vertebral body. Consequently, every collagen









fiber in the annulus fibrosus is equally strained, and resists distraction. However, the disc is not

often subjected to the tensile loads under normal physiologic activities. Also, even under the

distraction of the spine, the discs are under the compression load due to spinal muscle activity.

However, the annulus fibrosus is subjected to tensile stresses in various physiologic activities. In

addition to compression, any direction of the bending (flexion, extension, and lateral bending)

moves the instantaneous axis of rotation to lie outside of the disc and the disc is subjected to the

tensile stress at the opposite side of the bending (White & Panjabi, 1990).

Bending involves lowering one end of the vertebral body and raising the opposite end.

This causes distortion of the annulus fibrosus and the nucleus pulposus. In forward bending, the

anterior annulus is compressed and the disc tends to bulge anteriorly. The nucleus pulposus is

also compressed anteriorly, but the elevation of the posterior end of the vertebral body relieves

the pressure on the nucleus posteriorly (Brown et al., 1957; Shah et al., 1978). However, Roaf

(1960a) did not find any changes in shape or position of nucleus pulposus on the nucleographs of

the disc during flexion/extension. This supports the relevance of maintaining a slightly flexed

lumbar spine posture as a treatment and prophylaxis for the patients with low back pain. The

increase in disc pressure observed in vivo during bending of the lumbar spine may not be just

from bending, but from the result of compressive loads applied to the discs by the action of the

spinal muscles which are involved in bending motion (Ortengren et al., 1981).

During torsional movement of the inter-body joint, only the collagen fibers in the annulus

in the direction of movement have their points of attachment separated. Thus the annulus

fibrosus resists torsional movements with only half number of lamellae. Farfan et al. (1970)

conducted experiments using cadaveric vertebra-disc-vertebra construct including posterior

structures to examine the effect of torsional load. They found that the failures occurred at the









annulus in the final phase of the loading. The average failure torque for the normal discs was

found to be 25% higher than that for the degenerated discs, and the average torsional angles at

failure were 160 and 14.50 for normal and degenerated discs, respectively. Torsion of inter-body

joints seems be the most likely mechanical factor to injure the annulus.

In pure shear movements of the inter-body joint, only half of the fibers in the annulus are

strained, and the shear stress is raised mostly at the side of the loading. The shear stiffness in the

horizontal plane was found to be about 260 N/mm, and this could be the large force to cause an

abnormal horizontal displacement in the normal disc (White & Panjabi, 1990). This means that it

is rare for the annulus to fail clinically due to pure shear loading.

In addition to the load characteristics, the intervertebral disc demonstrates the time

dependent behavior which is called the viscoelastic property: creep and hysteresis. If a constant

force is applied to a viscoelastic structure for a prolonged time, further movement will be

detectable after the end of physiologic motion. If this movement is small in amplitude, occurs

slowly, almost imperceptible, then it is called as creep. Kazarian (1975) performed compression

creep study on functional spinal units (FSUs) and differentiated the disc specimens into four

grades according to the degree of degeneration. He found the creep and degeneration grade of the

disc are related. The non-degenerated discs creep slowly and reach their final deformation after a

long time, compared with the degenerated discs. This means that degenerated disc loses the

capability to attenuate shock and to distribute the load uniformly over the entire end-plate.

Viscoelastic structures also show differences in mechanical behavior during loading and

unloading. Restoration of the initial length of a structure from unloading occurs at a slower rate

and to a lesser extent than did the deformation from loading. This difference in mechanical

behavior is referred to as hysteresis, and reflects the amount of energy lost compared with









structure at initial loading. Virgin (1951) observed that hysteresis is largest in young people and

smallest in the middle-aged ones, and the lower thoracic and upper lumbar discs show less

hysteresis than the lower lumbar discs. He also found hysteresis decreased when the same disc

loaded repeatedly. This means that the disc is less protected against repetitive loads.

When forces are repeatedly applied to a material, it does not behave the same way each

time. Each application produces a certain amount of hysteresis, and the material is altered

slightly. Following many repetitions, small weaknesses accumulate and weakness in the material

becomes apparent. After several frequent repetitions of a stress, the material may fail at a certain

stress which is less than that required to damage the material following a single application of a

force. This is referred as the fatigue failure, and the fatigue tests of the disc were developed to

identify the number of load cycles that can be tolerated before disc failure develop. Brown et al.

(1957) conducted a fatigue test on the disc with a small constant axial load and a repetitive

forward bending of 5. The disc failure started to occur after 200 cycles of bending, and

complete failure occurred after 1000 cycles. However, the fatigue tolerance of the disc in vivo is

not known.

Vertebra

The basic morphology of the vertebrae in various regions of the spine from C3 to L5 is

approximately the same. The size and mass of the vertebrae increase from Cl to L5 vertebra.

This is the mechanical adaptation of the vertebrae to the progressively increasing compression

loads. Chalmers et al. (1966) observed that the strengths of vertebral cancellous bone of each

lumbar segment are almost the same. This means the variation of the vertebral strength according

to the spinal level is mostly due to the size of vertebrae. However, vertebral strength decreases

with age. Bell et al. (1967) found that a small loss of osseous tissue produces much loss of









vertebral strength: a 25% loss of the osseous tissue results in a more than 50% decrease in the

vertebral strength. The bone mineral density of female vertebrae is less than that of male at any

corresponding age, but the rate of decrease is not different between males and females (Hansson

& Roos, 1980).

The vertebral body is designed for load-bearing of large compression. First, the vertebral

body is not a solid bone block, but a shell of cortical bone and cancellous core. Second, the space

between the trabeculae in the cancellous core can be used as the channels for the blood supply

and venous drainage of the vertebral body. Likewise, the presence of bone marrow in the

intertrabecular spaces acts as a useful element for transmitting the loads and absorbing the force

(White & Panjabi, 1990). Rockoff et al. (1969) conducted compression strength tests on two

groups of vertebrae without posterior elements: the vertebral bodies with central hollow and the

other ones without outer shell. They found the loss of strength in both specimens, and the sharing

of the compressive load by the cortical shell was about 65%.

Facet joints are important stabilizing structures and carry about 18% of total compressive

load in the lumbar spine region (Nachemson, 1960). However, King et al. (1975) pointed out that

the load-sharing between facet joints and disc is much complex in their dynamics study with

whole cadavers. They observed that the load sharing carried by the facets could be 0-33% and

the value depended on the spine posture. Facet joints also contribute to torsional strength of FSU.

Farfan (1970) observed that the vertebral body-disc-body with longitudinal ligaments share the

torsional strength equally with the two facets and capsular ligaments, about 45% each. The

remaining 10% was carried by the interspinous ligaments.

Spinal Ligaments

The ligaments from C2 to the sacrum are similar, and seven ligaments are generally

referred to as the spinal ligaments in this region: the anterior and posterior longitudinal









ligaments, the intertransverse ligament, the capsular ligament, the ligamentum flavum, and the

interspinous and supraspinous ligaments. The anterior and posterior longitudinal ligaments attach

to the anterior and posterior edges of vertebral bodies and discs from basiooccipital to the sacrum

and coccyx. These ligaments become deformed by the relative separation between adjacent

vertebrae and by the bulging of the disc. Tkaczuk (1968) observed that the anterior longitudinal

ligament was twice stronger than the posterior longitudinal ligament. The intertransverse

ligament is known to have no mechanical significance in the lumbar region because of its small

cross-sectional size (Chazal et al., 1985). The capsular ligaments contribute to the flexion

stability in the cervical spine region (Panjabi et al., 1975). The ligamentum flavum has a high

percentage of elastin (80%) when compared to the other ligaments. This allows a large extension

without permanent deformation (Yahia et al., 1990). Nachemson and Evans (1968) found that the

ligamentum flavum has pre-tension, and this produces resting compression of the disc.

Consequently, the high elasticity and pre-tension of the ligamentum flavum minimizes the

chances of any impingement to the spinal canal during sudden spine motions.

Biomechanical Properties of Spinal Segments

Multisegmental Mechanics of the Spine

The basic motion segment of the spine is referred to as the functional spinal unit (FSU),

which consists of two adjacent vertebrae, intervertebral disc, and the connecting ligaments

without the spinal musculatures. Generally, a FSU exhibits similar biomechanics to those of the

entire spine, and can be used as a common testing specimen in vitro. The range of motion

(ROM) of a FSU is represented by the sum of two distinct phases: neutral zone and elastic zone.

The neutral zone is defined as the low-load response of FSU near the neutral position, and the

elastic zone is defined as the spinal behavior beyond the neutral zone up to the end of the

physiologic limit (Figure 3-1) (White & Panjabi, 1990).









Generally, the neutral zone is referred to as a quantitative measurement of the laxity

around the neutral position of a FSU. It is known to increase with degeneration, surgical injury,

repetitive cyclic loads, and high-speed trauma. In flexion/extension and lateral bending, the

neutral zone is the largest in the lower cervical region. In axial rotation, the neutral zone is the

largest in the C1-C2 region (Table 3-1).

By defining the neutral and elastic zones in the load-displacement curve of an FSU, the

coefficient of flexibility and stiffness can be calculated. The flexibility coefficient is defined as

the ratio of the displacement produced to the load applied. The stiffness coefficient is defined as

the ratio of the resistance offered to the displacement imposed. However, the load characteristics

of the spine is quite complex (nonlinear, biphasic, and viscoelastic) and cannot be demonstrated

by a single number. Previous studies showed that FSUs are more flexible in tension than in

compression in all regions of the spine. The shear flexibility is not quite different in each

direction (anterior, posterior, or lateral). The spine is more flexible in flexion than in extension in

all regions except the sacroiliac joint. Flexibility values for lateral bending are in between the

values of flexion and extension (Panjabi et al., 1988; Panjabi et al., 1976).

Axial rotation is generally known to be more harmful to the disc than the other motions,

except for a combination of axial rotation and lateral bending (Farfan et al., 1970). In the cervical

region, the spine is about 37% as flexible in torsion as compared with flexion. In the upper

thoracic region, the torsional stiffness is about the same as in flexion. The torsional flexibility of

the lumbar region is about 27% of flexion flexibility, which is the lowest value in all regions.

However, the torsional flexibility is much greater at the lumbosacral joint (55% of flexion) and

sacroiliac joint (250% of flexion) (White & Panjabi, 1990).









Individual physiologic motions of FSUs are inherently connected, and which is called the

coupling. Coupling of spinal motions is due to the geometry of individual vertebrae, connecting

ligaments, discs, and the curvature of the spine. The motion produced by an external load is

defined as the main motion, and the accompanying motions are called the coupled motion. In the

thoracic region, there is a strong coupling between all the motions in the sagittal plane

(translation and rotation) (Panjabi et al., 1976). The coupling of axial rotation with lateral

bending is a very common physiologic motion in the cervical and lumbar regions (Moroney et

al., 1988; Panjabi et al., 1977).

Regional Mechanics of the Spine

The entire spine is divided into cervical (upper CO-C1-C2, middle C2 C5, lower C5 -

Tl), thoracic (upper Tl T4, middle T4 T8, lower T8 LI), lumbar (LI L5), lumbosacral

(L5 Sl) and sacroiliac regions, based on the kinematic, kinetic, and clinical characteristics.

The upper cervical region is composed of occipital-atlanto-axial joints (CO-C 1-C2) and is

the most complex region of the spine, anatomically and kinematically. Most of the axial rotation

and some of the flexion-extension and lateral bending of the head come from the upper cervical

movements. The dominant atlantooccipital (CO-C1) motion is mostly flexion/extension, some

lateral bending, and tiny axial rotation (Table 3-2). The atlantoaxial (C 1-C2) articulation consists

of four joints: two atlantoaxial lateral joints, the atlantoaxial median joint (between anterior arch

of the atlas and dens axis), and a joint between the posterior surface of the dens and the

transverse ligament. The lateral atlantoaxial joint capsule is loose and allows a great deal of axial

rotation, in which the vertical axis of dens acts as a pivot about the atlas rotation. The possible

atlantoaxial motions are also summarized in Table 3-2.

The anatomical structures and the function of the middle and lower cervical regions are

quite different to those of the upper cervical region. The dominant motion in the lower cervical









spine is flexion/extension. Lysell (1969) observed the routes of each cervical vertebra in the

sagittal plane from flexion to extension or vice versa. The movement is a combination of

translation and rotation, and he called that movement the 'top angle', which indicated the arch

steepness of the route, generated during the flexion/extension. The arches were flat at C2,

steepest at C6 and followed by C7. The average ROM in healthy adults in the middle and lower

cervical regions are summarized in Table 3-3.

Dvorak et al. (1992) performed an in vivo test to measure the cervical ROM, based on the

modified inclinometer technique, to define the age- and gender- related differences. They found

a general tendency of decreasing cervical ROM as the age increased. The most drastic decrease

in cervical motion occurred at the age of 30-39 and 40-49 years. The cervical motions that did

not decrease with age were the rotation out of maximum flexion and the upper cervical rotation.

Females showed greater ROMs in all planes of cervical motion. However, there were no

significant differences between genders for the group over 60 years. The typical cervical ROM

values are presented in Table 3-4.

The human thoracic spine is a unique spinal region to be adapted to an erect posture and

load-bearing. The predominant posture of the thoracic spine is a kyphotic curve while the last

region (T11 L2) is almost straight in the sagittal plane. Thoracic kyphosis may arise from

postural and structural factors. Postural factor come from positioning of the spine due to the

ligamentous tension and muscle tone, as well as the disc configuration. The shorter ventral height

of thoracic vertebral bodies than dorsal one contributes to the structural kyphotic curve in the

thoracic region. Due to the kyphotic curve of the thoracic spine, the axial load applied to this

region generates a bending moment to cause further flexion. As a stability of this region, dorsal

tension-band capacity by the posterior ligaments and ventral weight-bearing by the vertebral









bodies are the vital combination to prevent spinal deformities in this region (Benzel &

Stillerman, 1999).

There are two transitional regions in the thoracic spine: cervicothoracic and

thoracolumbar junctions. The biomechanical characteristics of these junctional regions are

described as a blending of two adjacent regions. The upper thoracic region has a very limited

flexion/extension, whereas the caudal region from T10 has a larger range of flexion/extension.

The sagittal orientation of the facets in the lower thoracic region severely limits axial rotation

and, to a lesser extent, laterals bending. The facets of the upper thoracic spine are similar in

orientation to those of the cervical spine, and similar motion characteristics occurred at the

cervical and upper thoracic regions. Likewise, the facets of the lower thoracic spine are similar to

those of the lumbar spine and similar motions are seen in both lower thoracic and lumbar

regions. Representative ROMs of different motion segments of the thoracic spine are

summarized in Table 3-5.

The pattern of motion in the sagittal plane for the thoracic spine is similar to that in the

cervical spine. To describe the motion of the thoracic vertebra in the frontal and sagittal planes,

the 'top angle' was also employed. In the sagittal plane motion, the arch is quite small, and there

is no variation according to the level. The arch in the frontal plane is also flat, but greater than

that in the sagittal plane. Also in the frontal plane, the arch tends to increase from TI to T12.

There is also coupling of lateral bending and axial rotation in the thoracic region. The pattern of

coupling in the thoracic region is similar to the one in the cervical region. However, the coupling

pattern in the lower thoracic region is not as strong as that in the cervical region (White, 1969).

Movements in the thoracic spine are greatly limited by the facet orientation and by the rib

cage. The costotransverse and costovertebral articulations provide strong, stable attachment of









the thoracic vertebrae to the ribs. The costosternal articulations also contribute to the stability of

the thoracic spine (Pal & Routal, 1987). Oda et al. (1996) found a significant increase of

flexion/extension in thoracic motions after resection of the posterior elements. With the removal

of the costovertebral joints bilaterally, large increases in lateral bending and axial rotation were

observed. They concluded that the integrity of the costovertebral joints and the rib cage

significantly contribute to the spinal stability of the thoracic region.

In the sagittal motion of the lumbar spine, there is a cephalocaudal increase in

flexion/extension. The lumbosacral joint provide more sagittal plane motion than the other

lumbar motion segments do. For the coronal motion, the ROM for each lumbar level is about the

same, except for the lumbosacral region, which demonstrates limited lateral bending. Limitation

of lumbosacral motion is about the same for the axial rotation. Representative ROMs of the

lumbar spine are summarized in Table 3-6. Another important kinematic component in the

lumbar spine is the sagittal plane translation. In general, 2.0 to 2.8 mm is referred to as the

normal limit of anterior translation ofr a lumbar spinal vertebra, and 4.5 mm is an evidence of

clinical instability for a lumbar motion segment.

There are several patterns of coupling motion in the lumbar spine. Pearcy et al. (1984)

observed coupling of slight axial rotation and lateral bending with flexion/extension, based on

their stereoradiographic study. Another coupling pattern is that of lateral bending and axial

rotation. The direction of lateral bending with axial rotation in the lumbar region is that the

spinous processes point to the same direction as the lateral bending. It is opposite to the pattern

occurred in the cervical and thoracic regions. However, the coupling pattern of lateral bending

with axial rotation at the lumbosacral joint is the opposite of that in the lumbar region and similar

to the cervical and thoracic regions (Pearcy & Tibrewal, 1984).









The sacroiliac joint is partly synovial and partly syndesmotic. It is known to be completely

ankylosed in 76% of the subjects over 50 years of age (White & Panjabi, 1990). However,

studies of the motion about the sacroiliac joint have produced a wide range of results. Miller et

al. (1987) performed a kinematic study of the sacroiliac joints in cadaveric specimens. They

measured the displacements of the sacrum in relation to the ilium with each plane of loading.

They observed that lateral bending of one side was 1.40, anterior translation 2.74 mm, and axial

rotation of one side 6.21 in their study. Walheim et al. (1984) observed 2-3 mm of vertical

translation and 30 of rotation of the pubis at the symphysis pubis with one leg standing.

Sturesson et al. (1989) also observed tiny motions of sacroiliac joints in their stereoradiographic

study.

Biomechanical Performance of Spine In Vivo

Trunk Posturing

Trunk postures during various activities are related to the risk of developing a low back

disorder (Granata & Wilson, 2001). Spinal compression below 3,400 N may be considered as a

safe margin to prevent low back disorder in occupational population (Konz, 1982). However, a

spinal injury associated with the instability can occur at a low compressive load (Granata &

Marras, 1999). This means that the appropriate recruitment of the spinal and trunk muscles

provide the stable support for the large load applied to the spine, but some postures may limit the

ability of the muscles to maintain spinal stability (Wilke et al., 1995).

Nachemson (1966; 1981) conducted studies to verify the compression load applied to the

disc in vivo by measuring the intradiscal pressure at L3/L4 level in various postures. The lowest

compression was observed in the supine position (300 N), which is about 50% of that in standing

(700 N) without external loads. During sitting without a back support, the compression went up









to 1,000 N. Forward bending of the trunk generated the largest compression force to the spine up

to 1200 N. Ledet et al. (2005) measured intradiscal loads in baboons in vivo and found similar

results to Nachemson's. Takahashi et al. (2006) also performed the same testing at L4/L5 level

with young, healthy subjects, and found similar results. However, their values in specific

postures are different from those reported by others. The values of compression loads in different

studies are summarized in Table 3-7.

The data from Takahashi et al. (2006) were greater than those from Nachemson (1966;

1981). Takahashi et al. explained the differences with several reasons. Their subjects were all

young and normal subjects. Their techniques seemed to be much more developed than the old

one used in Nachemson studies. The specific level used in intradiscal measurement was lower

than that in Nachemson studies. Takahashi et al. concluded that the risk of intervertebral disc

injuries or degeneration could be induced by a simple repetitive forward bending of the trunk in

everyday movements. However, slightly flexed, relatively straight or nonlordotic position of the

lumbar spine during standing is used more frequently in the populations who complain less often

about back pain. Also, the flexion of the hip reduces the tension of the psoas muscles and the

lordosis of the lumbar spine, resulting in reduced load on the lumbar spine.

In addition to the forward bending of the trunk, twisting, lateral bending and asymmetric

posture combined with lifting are also known to increase the risk of low back problems (Kelsey

et al., 1984; Marras et al., 1993). Some EMG studies explained that the high loads to the spine in

the unstable and asymmetric postures are from the co-contraction of agonistic and antagonistic

spinal muscles (Cholewicki et al., 1997).

Weight Lifting

Lifting and bending episodes accounted for 33% of all work-related causes of back pain

(Damkot et al., 1984). Increasing weights anterior to the spinal column greatly increases the









forces which are exerted on the lumbar spine. This is due to the forces developed in the spinal

muscles in order to maintain the equilibrium. The resultant forces applied to the fulcrum, which

is the lower lumbar region, are very high.

The distance of the weight from the body is directly related to the joint reaction force

(high disc pressure) to the lower back, the greater force required by the erector spinae muscles

(high electromyographic activity), and a need of greater truncal support to protect the spine (high

truncal pressure). This indicates the significance of closer distance of the object to the body as

the proper lifting technique. Another factor to determine a proper lifting is the back posture. For

the optimal lifting methods, the squat lift (knee bent and back straight) is generally considered to

be safer than the stoop lift (knee straight and back bent) in bringing the load closer to the body,

and reducing the back muscle demands to counterbalance the additional moments. These

techniques consider the posture of the back in addition to the distance of the objects. However,

many workers prefer the stoop lift over the squat lift. There is an increased physiologic cost and

more rapid fatigue development in a squat lift. And the squat lift is not always possible because

of the lift setup and load size. Likewise, the risk of developing low back pain by lifting tasks

depends rather on the lumbar posture than the choice of lifting techniques (van Dieen et al.,

1999).

There is a conflict about the favorable lumbar postures during lifting tasks. Some

advocate lordotic and straightening lumbar posture because they believe increased erector spinae

activity is beneficial in augmenting spinal stability and decreasing anterior shear force on the

spine (Hart et al., 1987). However, others favored the kyphotic lift (flexed lumbar spine),

because they believe passive ligaments of the lumbar spine can relieve the active extensor

muscles (Gracovetsky et al., 1981). Cholewicki et al. (1992) tested professional power lifters to









evaluate the kinematics of the lumbar spine and resultant posterior ligament lengths during

lifting tasks. They observed significantly smaller lumbar flexion and increased lengths of the

ligaments during the lifting when compared with the full flexion of the trunk. They concluded

that the back muscles were substantially responsible for resisting trunk flexion moments during

heavy lifting.

Arjmand and Sirazi-Adl (2005) tested the kinematics of the lumbar spine and activity of

selected spinal and trunk muscles in healthy subjects during a static lifting. They examined the

lumbar spine postures lordosiss, kyphosis) during the lifting procedures and found the lordotic

posture increased extensor muscle forces, axial compression and shear forces at L5/Sl. They

recommended the moderate flexion posture of the lumbar spine as a posture of choice in static

lifting tasks. Lifting capacity is generally used to determine the degree of spinal impairment

state, and the back strength and aerobic capacity are known to be the contributing factors

(Matheson et al., 2002).

Sitting and Standing

The back rest and lumbar support is known to decrease the loads applied to the spine

during sitting. Andersson et al. (1977) performed a study to estimate intradiscal pressure of

L3/L4 under different backrest inclinations and lumbar support conditions. They found the

highest intradiscal pressure in the sitting with no lumbar support and a 900 backrest inclination,

and the lowest disc pressure and the least electromyographic activity of the paraspinal muscles in

the sitting of 1200-inclination and 5-cm width of lumbar support. More specifically, the lumbar

support has the greater influence on lumbar lordosis and the backrest inclination has more

influence in reducing the loads on the lumbar disc (Andersson et al., 1979). The arm rest is also

known to reduce intradiscal pressure (Kelsey & Hardy, 1975).









The ability to stand up from sitting is a prerequisite for walking and other independent

function. Sit to stand demands coordinated movements of linked body segments to transport the

center of body mass in a horizontal then vertical direction while maintaining balance over a

small base of support, the feet. Previous studies reported these kinematic characteristics of

standing from sitting: flexion of the trunk and hips bring the center of mass forward, followed by

bilateral extension of the lower limb joints, and trunk extension to raise the body in a vertical

direction over the feet (Doorenbosch et al., 1994). Tully et al. (2005) studied the kinematics of

the body segments including the thoraco-lumbar region during standing from sitting. They

divided the standing movement into phases before and after the lift-off, based on the relation of

position between the buttock and the sitting object. Before the lift-off, they observed a forward

leaning of the trunk, accomplished by concurrent lumbar and hip flexion (1:3). As the lumbar

spine flexed the thoracic spine extended, resulting in a trunk angle of 45.70 at lift-off with respect

to the horizontal plane. Following the lift-off, the hip and lumbar spine extended and the thoracic

spine flexed, with the standing thoracic angle approximating the initial thoracic posture in sitting.

Walking

The biomechanical function of the trunk during walking has been investigated extensively.

Earlier studies examining the trunk kinematics during walking considered only the entire trunk

motions with respect to the pelvis. Later studies examined the movements of the lumbar and

thoracic spines or pelvis. Crosbie et al. (1997) studied the patterns of spinal motion during

walking using a model including upper and lower trunk, lumbar and pelvis segments. They used

three spatial surface markers in each spinal segment on the back surface of the subjects. The

pattern of flexion/extension of each segment was generally biphasic throughout the gait cycle.

The pelvis rotated into negative pelvic tilt at heel strike. This was followed by a counter-motion









to a maximum positive pelvic tilt in the single support phase. The lumbar spine reached

maximum flexion at heel strike. This was followed by a rapid extension to neutral until the single

support phase. The lower thoracic segment extended maximally at heel strike, and returned to a

neutral at mid-stance, then extended again through the late stance. The counter-motions occurred

between the lumbar and lower thoracic segments at heel strike. They concluded that spinal

segments demonstrate complementary movements to the motion of the pelvis, and pelvic motion

responds to the need of advancing the lower limb and transferring the body weight from one

supporting side to the other. Lumbar ROMs during walking and running are summarized in

Table 3-8.

Thousands of repetitive low level loadings are applied to the spine in everyday activities.

During normal walking, activation of the spinal muscles, acceleration of the trunk, combined

with the external loadings result in cyclic spinal loads. Some studies investigated the magnitude

of these loads, in conjunction with the spinal motion and muscular activities during walking.

Callaghan et al. (1999) conducted a biomechanical study using two models to estimate loads

applied to the L4/L5 level: linked segment model with EMG technique and rigid segment model

with inverse dynamic technique. The joint loading (at L4/L5) calculated by the EMG model

resulted in large increases in the maximum compressive forces, compared with the joint reaction

forces calculated using inverse dynamics. Including the muscular component resulted in a more

than three-fold increase in joint load. The compression loads applied to the lower lumbar level

during walking from different studies are summarized in Table 3-9.

Unlike the differences in compressive load, the joint shear forces (anterior/posterior,

lateral) obtained using the two techniques were similar. The flexion/extension moment had two

peaks throughout the gait cycle. At heel contact there was a peak flexor moment followed by a









peak extensor moment around toe-off. During a faster speed gait, the flexion/extension moment

at L4/L5 shifted to the extension side and demonstrated a high extensor moment around toe-off.

The lumbar spine motion with respect to the pelvis was quite consistent within and between

subjects. The sagittal motion of the lumbar spine showed several dominant phases. Following

heel contact a flexion phase was present until the relative spine motion reached maximum

flexion just following toe off. And the spine remained in a constant posture and additional

extension phase during the single stance. They concluded that the loads and motions for the

lumbar spine during gait depended on the walking speed. Increasing walking speed increases the

lumbar spine ROM, activation of spinal and trunk muscles, and anterior/posterior shear forces.

Running

Running typically requires the spine moving through only a limited range of motions.

Acute injuries to the spine directly from running activities are relatively infrequent and amounted

approximately 11 13% of all injuries sustained (Walter et al., 1989). The frequent and

significant spine injuries related to running are largely due to the repetitive axial compressive

loading of the spine which occurs during the foot strike in each stride. A typical distance runner

who runs 130 km per week in training might subject to about 40,000 foot strikes per week

(Cavanagh & Lafortune, 1980). Several case studies have highlighted the overuse injuries of the

lumbar spine and pelvis in running (Guten, 1981; Koch & Jackson, 1981).

Schache et al. (2002) conducted a kinematic study of lumbar spine and pelvis during

running. The average ROMs of the lumbar spine and pelvis on each plane are summarized in

Table 3-8. They found significant inverse correlations between flexion/extension of the lumbar

spine and anterior/posterior tilt of the pelvis and lateral bend of the lumbar spine with obliquity

of the pelvis. Essentially, as anterior tilt of the pelvis increased during the terminal stance,

extension of the lumbar spine also increased. When the lumbar spine was laterally bent to the









same side of the foot contacted to the ground, the pelvis was lowered on the opposite side. When

the lumbar spine began to laterally bend towards the opposite side of the foot contacted, the

pelvis began to elevate on the opposite side of the foot contacted. They also found a significant

positive correlation between the axial rotations of lumbar spine and pelvis. However,

coordination between the axial rotations of the lumbar spine and pelvis was out of phase by 21%

of the running cycle.

The kinematic pattern of axial rotation of the pelvis during running is different to that

during walking. At the initial contact of one foot during walking, the pelvis showed maximal

rotation to the opposite side (Whittle & Levine, 1999). This movement helps in augmenting the

stride length at that time. With the loss of the double support phase during running, the pelvis

along with lower extremities are no longer required to be engaged in a stride lengthening

mechanism. At the initial contact during running, the pelvis was rotated to the same side of the

foot contacted. This movement was suggested as minimizing the horizontal braking force at the

initial contact to avoid potential loss of running speed (Novacheck, 1998; Schache et al., 1999).

Schache et al. (2003) performed another kinematic study of lumbo-pelvic-hip complex

during running to define the gender differences. They found that females displayed a shorter

stance time, swing time, stride time and stride length, and a higher stride rate than males. Mean

waveforms were different in the peak-to-peak oscillations and the offset of pelvis

anterior/posterior tilt. Females displayed greater amplitudes of lumbar spine lateral bend and

axial rotation, pelvis anterior/posterior tilt, obliquity and axial rotation, and hip

adduction/abduction than their male counterparts. The mean positions of anterior pelvic tilt

across the running cycle were 20.20 for females and 16.90 for males. The prevalence of pelvic-









femoral stress fractures in female runners might be explained by these findings (Bennell et al.,

1996; Pavlov et al., 1982).

Wilke et al. (1999) studied the intradiscal pressure of non-degenerated disc at L4/L5 in a

45-years old man during various activities. They observed intradiscal pressure of 0.5 MPa during

relaxed standing and 0.35-0.85 MPa during jogging with tennis shoes. Rohlmann et al. (2001)

also recorded a peak intradiscal pressure 0.85 MPa while jogging on the treadmill of, which was

170% of the pressure noted in standing. The intradiscal pressures recorded in different studies

are summarized in Table 3-10.

There are several factors suggested in previous studies that affect the spinal posture and

loading during running. Exhaustive running, as fatigue occurs, has shown biomechanical changes

in the legs that lead to a lower effective body mass during heel strike (Derrick et al., 2002). This

resulted in increased peak leg impacts and increased shock attenuation, which may change the

spine loading forces with fatigue. Therefore, the loads applied to the spine may vary during a

run, depending on the lower extremity activity and the level of fatigue (Lennard & Crabtree,

2005).

Another factor that affects the spinal loading during running is the type of shoes and insole

materials utilized. An impulsive shock wave is generated at heel strike that is transmitted from

the lower extremities through the spine. The use of shock absorbing insoles has been used to

treat low back pain patients to lower the shock wave at low back level. On the other hand, the

development of external force and the transmissibility of impact forces through the human body

are increased by wearing soft soles. Ogon et al. (2001) performed a study to investigate the

behavior of low back muscles during jogging barefoot and wearing identical athletic shoes. They

observed that wearing shoes and insert materials decreased the rate of shock transmission to the









lower back and reduced the time interval between peak acceleration of lower back and peak

spinal muscle response in jogging. It is from the increased latency between heel strike and peak

acceleration at the lower back by wearing shoes. They suggested wearing shoes decreases the

time interval between maximum external (peak acceleration at lower back) and maximum

internal force (generated by the spinal muscles) in the lower back during running. Wen and

associates (1997) found a significant correlation between leg length discrepancy and onset of

lumbar pain within 12 months of running in a marathon training program. However, the clinical

importance of leg length discrepancy in short distance running is not clear.

Biomechanical Etiology of Spinal Pain

Spinal pain can be caused by trauma, infection, tumor, and systemic diseases. However,

the term 'spinal pain' is generally used to refer to the cervical, thoracic, and lumbar pain that is

not related to these injuries and diseases. The common neck pains with or without arm pain and

the back pain with or without leg pain which are frequently encountered in daily livings and

cause the spinal degeneration are called spinal pain. Although there are specific considerations

associated with spinal pain in different regions of the spine, there is much similarity in different

spinal regions. Spinal pain usually occurs in the more mobile and lordotic portions of the spine

and onsets at 30-50 years of age. Spinal pain occurs most often in the lumbar region, followed by

the cervical region, with the lowest incidence in the thoracic region. Spinal pain can come from

direct irritation of the nerves that innervate most of spinal structures. The posterior annulus

fibrosus, the posterior vertebral body and the posterior longitudinal ligament are innervated by

the sinu-vertebral nerve, which is considered the most common origin of spinal pain (Bogduk &

Twomey, 1997). Another type of spinal pain is the indirect referred pain which is not fully

understood and explained at this moment. Spinal pain is a major socioeconomic problem and

approximately 80% of all back problems are of unknown origin (Vogt et al., 2001).









Specific motion, force, and high-quantity repetitive loading, or any combination of these

may serve as mechanical stimulus to the spine, and which can be referred to as the abnormal

mechanical causes of spinal pain and degeneration either quantitatively or qualitatively. There

are many biomechanical factors known to contribute to spinal pain: vibration, lordosis, torsion,

driving motor vehicles, material handling, leg length discrepancy, etc. However, there are still

controversies on the roles of mechanical factors relative to spinal pain.

Vibration

Epidemiologic studies reported increased spinal pain and/or disc disease in those who

drive more than 3 hours per day or operate vibrating equipment (Frymoyer et al., 1983).

Vibration, particularly in the frequency domain of 5 to 15 Hz in which resonance of the spine can

occur (Goel et al., 1994), is considered a key etiologic factor in low back pain (Magnusson et al.,

1996), neurovestibular disorders (Seidel et al., 1988), and a causal factor in circulatory disorders

such as Raynaud's syndrome (Dandanell & Engstrom, 1986). Thus, the industries such as the

transportation and construction, as well as the military are working toward minimizing

occupational exposure to potentially noxious mechanical stimuli (Bongers et al., 1988; de

Oliveira et al., 2001).

Brumagne et al. (1999) performed a study to test the proprioceptive changes in response

to vibration in human. They applied a vibration (70 Hz, 0.5 mm amplitude) to the multifidus

muscle for 5 seconds and measured the lumbosacral repositioning accuracy. They found a

significant increase in directional error during vibration of the paraspinal muscles. The subjects

had the illusion during vibration that their pelvis was more posteriorly tilted, and accordingly,

they undershot the target position. They explained their results with the reflex inhibition of the

muscle spindle, which play an important role in proprioception. They concluded that vibration on

the spine and trunk muscle can result in damage to or dysfunction of the muscle spindles.









Consequently, a decreased muscle spindle input could jeopardize spinal proprioception and

segmental stability, and likely to make the spine more vulnerable to injuries and low back pain.

On the other hand, whole-body vibration has been utilized as an exercise therapy for

musculoskeletal problems in sports, geriatrics, and rehabilitation laboratories (Bosco et al., 1999;

Rittweger et al., 2000). Vibration is thought to elicit muscle activity via stretch reflexes (Clark et

al., 1981). Rittweger et al. (2002) observed that metabolic power increased during the whole-

body vibration from a ground platform (amplitude of 2.4 mm, 5.0 mm, 7.5 mm, frequency of 18

Hz, 26 Hz, 34 Hz, duration of 4 minutes) and that this metabolic power is augmented by the

application of additional axial loads. Ritweger et al. (2002) performed another study on patients

of chronic lower back pain devoid of specific spine diseases to verify the therapeutic effects of

vibration exercise (maximum amplitude of 6 mm, frequency of 18 Hz, duration of 4 to 7

minutes). They compared whole-body vibration exercise with lumbar extension exercise, and

observed a significant reduction in pain sensation and pain-related disability in both groups.

Lumbar extension torque increased in the vibration exercise group, but significantly more in the

lumbar extension exercise group. They concluded that a well-controlled vibration can be the cure

rather than the cause of lower back pain.

Some animal studies indicated that brief (<20 min) daily durations of extremely low-level

(0.5g), high-frequency (15-90 Hz) vibration can be strongly anabolic to the trabecular bone,

increasing bone mineral density, trabecular width and number in the weight-bearing skeleton

(Rubin et al., 2001; Rubin et al., 2002). These studies suggested the osteogenic potential of

extremely low-level mechanical stimuli as a treatment for osteoporosis. Rubin et al. (2003)

studied transmissibility of high-frequency (15-35 Hz) ground-based, whole-body vibration to the

proximal femur and lumbar vertebrae. They observed 30-130% of transmissibility of the loading









vibration, regardless of the target region, frequencies, and posture of the subjects. In addition,

transmissibility at the hip was different to the spine, mostly at the lowest frequencies in this

study. For the loading frequency less than 20 Hz, the resonance at the hip exceeded 100% in the

erect standing and relaxed standing postures. However, the resonance at the lumbar vertebrae

was lower than that at the hip, and it was constantly maintained at the high frequency of loading

vibrations and the other standing postures (relaxed standing, bend knee). They mentioned that

slight changes in posture can have significant influence on the degree of vibration to be delivered

to the spine or hip regions. They also emphasized the possible undesirable side effects of using

whole-body vibration, against the prevention strategy for the osteoporosis. They suggested that

vibration which approaches 1 g (9.8 ms-2), even at beneficial high frequencies, should be avoided

considering the risks to many physiologic systems.

Lordosis

Reduced or flattened lumbar lordosis has signified lumbar spine problems in previous

studies (Adams & Hutton, 1985; Evcik & Yucel, 2003; Farfan et al., 1972). On the other hand,

the cultural groups who spend considerable time in lumbar flexed position are known to suffer

less low back pain and disc degeneration (Fahrni & Trueman, 1965). Although there are

controversies raised in previous studies, the biomechanics of lumbar lordosis and back pain were

found to be closely related in some instances. Frymoyer et al. (1984) performed a radiologic

study to investigate the lumbar lordosis in low back pain patients. They found no association of

lumbar lordosis with low back pain. Murrie et al. (2003) conducted MRI assessments of lumbar

lordosis in between patients of low back pain and normal controls, and did not find any

significant difference in both groups. Instead, they observed that lumbar lordosis is more

prominent in women and those with a higher body mass index.









However, there are some cogent observations on the significant changes in the facet joints

and discs associated with spine extension and hyperlordosis. Dunlop et al. (1984) reported that

each degree of increased extension of the spine leads to a 4% increase in peak articular pressure

in the facet joints. Yang and King (1984) reported that arthritic facet joints may bear up to 47%

of the load transmitted. Thus, these increased loads to the facet joints lead to abnormal motion to

the inferior facet articulation and damage to the joint structures, which can cause low back pain.

The available evidence does not support strong conclusions, but there seem to be disadvantages

to hyperextension of the lumbar spine if there is facet joint arthritis or disc pathology.

Roussouly et al. (2005) performed a radiologic study to classify the normal variation of

sagittal lumbar spine and pelvis. They found that sagittal alignment of the human spine and

pelvis was highly variable in different individuals in a standing position. The angle of the

superior end plate of S with respect to the horizontal axis varied between 200 and 650, the angle

of global lumbar lordosis varied between 410 and 820, and the number of vertebral bodies in a

lordotic orientation varied from 1 to 8. The characteristics of the lumbar lordosis were most

dependent on the orientation of the sacral slope and the pelvis. The upper arc of lumbar lordosis

remained relatively constant, with an average value of approximately 200. In contrast, the lower

arc of lordosis was the most important determinant of the global lordosis. They concluded that

changes in the specific sagittal alignment and lumbar lordosis are potentially responsible to

indicate degenerative changes and symptomatic back pain, instead of utilizing the vague term of

lordosis.

Torsion

Axial rotation of the spine involves torsion of the intervertebral discs and impaction of

the facet joints. It is considered to be a possible mechanism of spinal damage and pain, especially









to the lumbar spine (Hadjipavlou et al., 1999). The assumed injury mechanism is that shear

loading of the annulus and the damage of the facet joints and ligament structures may lead to the

segmental damage. However, the axial torque was not considered a significant factor to

contribute to the disc degeneration in previous studies. Because the facet joints limit the torsion

to small range (1-2), it does not appear to allow enough stress to lead to the disc damage

(Adams & Hutton, 1981). Likewise, in an intact disc, the facet joints and posterior ligaments are

known to protect the intervertebral disc from torsional loading. Because the axis of rotation of a

lumbar vertebra passes through the posterior part of the vertebral body, all the posterior elements

of the moving vertebra swing around the axis during the axial rotation (Cossette et al., 1971).

Quantitative analysis revealed that the disc contributes 35% of the resistance to torsion, and the

remaining 65% are being exerted by the posterior elements (Farfan et al., 1970). Another

experimental study showed that the facet joints contribute 42 to 54% of the torsional stiffness in

an FSU (Asano et al., 1992).

In contrast, a study by Liu et al. (1985) supported that cyclic torsional loads can lead to the

failures in the disc, facet, laminae, and capsular ligaments. The anterior and posterior

components are probably damaged or irritated by the axial torque of the lumbar spine. The initial

torsional loading on the vertical axis are borne by the posterior elements (White & Panjabi,

1990).

Biomechanical Performance of Painful Spine

Patients with low back pain demonstrate a change in the mobility of the spine (Pearcy et

al., 1985), deficits in reaction time, coordination (Luoto et al., 1996), and postural control with

reduced velocity (Marras & Wongsam, 1986) when compared with healthy subjects.

Differentiating the mechanical performance of pathologic spine with the normal variation









associated with gender, aging, and physical function will be the basic step to determine the

treatment methods for patients with spinal pain. However, previous studies have been limited to

identifying the behaviors of the spinal kinematics of low back pain patients, instead of

approaching the various biomechanical performance of the spine.

Vogt et al. (2001) performed a kinematic study of lumbar spine during the treadmill

walking in patients of chronic low back pain. They found significantly shorter stride times and

stride-to-stride variability in all anatomic planes in these patients, as compared with healthy

subjects. Decreased stride time, suggesting smaller steps of the patients, could be interpreted as a

rigid or more cautious walking pattern and a protective way to reduce or avoid the pain. The

increased between-subject variability could be interpreted as patients' individual adaptations and

adjustments in walking behavior. These findings were interpreted as changed neuromuscular

strategies to maintain an effective manner of locomotion which could be mediated by the altered

proprioception in the lower back region. Nevertheless, the overall pattern of angular spinal

displacements in patient group was shown to be within normal limits. It was suggested that pain

of musculoskeletal origin had no significant effect on the magnitude of lumbar angular

displacements.

Shum et al. (2005) performed a kinematic study of the lumbar spine and hip during sitting

and standing movements in patients of low back pain. They found that the mobility and velocity

of the spine and hip were significantly limited and the contribution of the lumbar spine relative to

that of the hip was reduced in back pain subjects. The patients with low back pain, in particular

those with positive straight leg raising (SLR) sign, had altered hip-spine coordination. Shum et

al. (2006) performed another spine kinematic study for a picking up activity, and found the same

reduced mobility of lumbar spine and hip in low back pain subjects.









Al-Eisa et al. (2006) completed a kinematic study of the trunk in patients of unilateral

nonspecific low back pain during sitting. They found a significant correlation between pelvic

asymmetry and asymmetric trunk motion in the patient group, and suggested that people with

low back pain may have a distinct compensatory mechanism, secondary to the pelvic asymmetry

from the unilateral low back pain, which put the lumbar spine under higher stresses. They

concluded that movement asymmetry, rather than ROM, may be a better indicator of disturbed

function for people with low back pain.

McGill et al. (1999) studied the motions of the spine and trunk muscle activity in normal

elderly subjects (without back pain) during trunk posturing movements. They found the elderly

exhibited slower motion, reduced ROM in full flexion and lateral bending. Furthermore, there

was more coupled motion in the twisting efforts and abdominal muscles appeared to become

more active earlier in the lateral bending movement. The earlier activation and increased co-

contraction suggested that elderly people might be seeking greater stabilization either for general

balance or for actual spine stabilization.

Landing Biomechanics

Previous studies on landing mechanics were mostly descriptive in biomechanical

properties of various landing conditions. In majority of previous studies focused on defining

injury mechanism (Fagenbaum & Darling, 2003; Wikstrom et al., 2006), gender differences

(Kernozek et al., 2005), and conditional variability (Schot et al., 2002) of landing mechanics,

associated with knee and ankle joints. There were several different techniques and models

developed for the landing biomechanics of lower extremity joints (Baca, 1999; Nagano et al.,

1998; Spagele et al., 1999). Additionally, there were several landing techniques tested to

compare the efficiencies and injury rates during landing activities (McNair et al., 2000; Onate et

al., 2005). Because landing from a variety of jumping is a good example of active movements,









the ability of a subject to control kinetic and kinematic changes applied to his/her body segments

from landing can be a good measure to determine the functional status. However, most studies

focused on the performance of lower extremity joints, and to our best knowledge none were

developed to evaluate the performance of upper body segments above the hip joints during

landing so far.

Biomechanical Performance of Lower Extremity Joints during Landing

Landing involved in many activities are vigorous and violent in nature. There are many

reports about landing injuries from various activities (Ford et al., 2003; Salci et al., 2004). Many

studies have been conducted to define the biomechanical etiology of lower extremity injuries

during landing procedures (James et al., 2003; McNitt-Gray, 1993; Spagele et al., 1999). Vertical

ground reaction force (GRF) from the ground impact of landing ranged from 1.0 to 14.0 times of

body weight for normal subjects (Caulfield & Garrett, 2004; Decker et al., 2002; Fritz &

Peikenkamp, 2001; James et al., 2003). Landing conditions investigated include landing height

(McNitt-Gray, 1993; Zhang et al., 2000), landing posture and technique (Eloranta, 1996; Kovacs

et al., 1999), joint stiffness of lower extremity joints (Devita & Skelly, 1992; Horita et al., 2002),

and the number of legs involved in landing (Caulfield & Garrett, 2004; Hass et al., 2003).

Devita and Skelly (1992) studied biomechanical variables applied to the lower extremity

joints from drop landing with a fall-height of 59 cm, comparing soft and stiff landing conditions.

Soft and stiff landings were defined with knee flexion angle after ground impact as greater and

less than 900. The shapes of GRF, moment, and power curves were identical between both

landings. Larger GRF, hip extensor, knee flexor, and ankle plantar flexor moments were

observed during descent in the stiff landing, which produced a more erect body posture and a

flexed knee position at impact. Also, the stiff landing exhibited larger power in ankle muscles,









while the soft landing showed larger powers in knee and hip muscles. They concluded that ankle

plantar flexors absorbed more energy in the stiff landing, whereas the hip and knee extensors

absorbed more energy in the soft landing.

To compare gymnasts and recreational athletes, McNitt-Gray (1993) evaluated lower

extremity joints kinetics during drop landing with different landing heights. They found

gymnasts chose to dissipate the impact loads by using the larger ankle and hip extensor moments

at higher impact velocities than recreational athletes who chose to adjust their strategy by using

greater degrees of hip flexion and longer landing phase durations than the gymnasts. The greater

demands placed on the ankle and hip extensors by the gymnasts, as compared to the recreational

athletes, was explained by the need to maintain balance during competitive gymnastics landings

or, by the inability of recreational athletes to produce larger extensor moments at the ankle or hip

during landings from the great heights.

Zhang et al. (2000) studied lower extremity joints kinetics with different landing heights

(0.32 m, 0.62 m, and 1.03 m) and different landing techniques (soft, normal, and stiff landings).

They found increases in peak GRF, peak joint moments, and powers with increases in landing

height and stiffness. The mean eccentric work were 0.52, 0.74, and 0.87 J/kg by the ankle plantar

flexors, 1.21, 1.63, and 2.26 J/kg by the knee extensors, and 0.94, 1.31, and 2.15 J/kg by the hip

extensors, for heights of 0.32, 0.62, and 1.03 m, respectively. They concluded that knee

extensors were consistent contributors to energy dissipation, while the ankle plantar flexors

contributed more in the stiff landings and the hip extensors did more in the soft landings. This

shift from ankle to hip strategy was observed as landing height increased. They explained that

larger volume of proximal muscles (knee and hip joint muscles) of the lower extremity was more

capable of energy absorption compared with the ankle muscle group. Another study suggested









that biarticular muscles are used effectively for power transportation during locomotion (Bobbert

& van Ingen Schenau, 1988).

Gender Difference

The majority of knee injuries occurred during landing from various activities are caused

by the non-contact mechanism. Numerous studies have found females have a higher rate of non-

contact anterior cruciate ligament (ACL) injuries compared to males (Decker et al., 2003;

Kernozek et al., 2005). Many studies had been conducted to identify different mechanical

properties of landing in each gender and etiology of this gender disparity. Anatomically or

intrinsically, small cross-sectional area of ACL (Feagin & Lambert, 1985), narrower

intercondylar notch (Souryal et al., 1988), greater quadriceps angle, and greater knee laxity

(Malinzak et al., 2001) of females have been suggested to contribute to the higher ACL injuries

in females than in males. Except for the intrinsic anatomical factors, the differences in level of

conditioning, level of muscle strength, and motor control strategies in females were suggested as

the extrinsic factors, related with gender disparity (Delfico & Garrett, 1998).

Chappell et al. (2002) compared knee kinetics of male and female recreational athletes

performing forward, vertical, and backward stop-jump tasks. They observed females exhibited

greater proximal anterior shear force, greater knee extension and valgus moments than males did

during the landing phase. During the takeoff phase, males showed greater proximal tibia anterior

shear force than females. They concluded that females might have altered motor control

strategies that result in knee positions in which ACL injuries might occur.

Decker et al. (2003) compared the biomechanical variables of lower extremity joints

between male and female subjects performing a 60 cm drop landing. They found females

demonstrated a more erect landing posture and utilized greater hip and ankle joint range of

motions and maximum joint angular velocities when compared to males. Females exhibited









greater energy absorption and peak powers from the knee extensors and ankle plantar-flexors

compared to males. Energy absorption contributions revealed that the knee extensor was the

primary shock absorber for both genders. The ankle plantar flexor was the second largest

contributor to energy absorption for the females and the hip extensor was for the males. The

different shock absorption strategy used in females was proposed to provide a greater potential

risk for non-contact ACL injury for females under certain landing conditions.

Kernozek et al. (2005) compared gender differences in frontal and sagittal plane

biomechanics during a 60-cm drop landing. They observed that females exhibited greater peak

hip and knee flexion, and ankle dorsiflexion angles in the sagittal plane, and greater peak knee

valgus and ankle pronation angles in the frontal plane. Females exhibited greater peak vertical

and posterior GRF, and reduced varus moment than males. They noted the importance of gender

differences in the frontal plane variables in addition to those in sagittal plane.

Landing Stiffness

During landing from various activities, the actions of different musculoskeletal structures,

including muscles, tendons, and ligaments, are integrated together so that the overall skeletons

behave like a spring. As a result, landing from a jump can be modeled by using a simple spring-

mass system (Asmussen & Bonde-Petersen, 1974). Stiffness of the leg spring represents the

overall stiffness of the integrated musculoskeletal system during the landing phase, which is

referred to as leg stiffness. Leg stiffness influences the kinetics and kinematics applied to the

whole body during landing. Leg stiffness is greatly determined by the knee joint stiffness, which

is mostly affected by knee extensor muscles. Joint stiffness is calculated by a linear regression of

the knee joint moment/angle relationship. The interaction between leg stiffness and reflex

activities plays a major role in regulating muscle power and performance in pre- and post-









landing to absorb or dissipate the large energy developed from the ground impact (Horita et al.,

1996).

Horita et al. (2002) studied the interaction between pre-landing activities and stiffness

regulation of the knee joint during a drop landing followed by a countermovement jump by

examining landing motions, GRF, and EMG activity of the vastus lateralis during the pre- and

post-landing. They divided the contact phase into three phases from initial contact to takeoff of

the countermovement jump: (1) initial impact to initial peak of knee joint moment, (2) initial

peak to onset of pushoff, and (3) concentric phase until takeoff. Drop landing performance was

evaluated using the takeoff velocity, average contact time, and knee joint moment. A positive

correlation was found between positive peak power of knee joint and the knee joint moment.

However, they did not find any significant relationships between any drop landing performance

parameters and ankle measures. They explained leg stiffness with a combination of pre-

contraction of the vastus lateralis muscle and knee joint angular velocity at touchdown. They

proposed two types of landing motions with regard to the pre-landing motion of the knee joint.

The proper pre-landing movement could be characterized by the knee flexion just before

touchdown, which is associated with a high initial joint stiffness coupled with the high joint

power. This was called bouncing type in their study, which is close to the plyometrics. On the

other hand, an inadequate pre-landing movement, associated with incomplete knee flexion

induced subsequent deep-knee flexion after touchdown was called the absorbing type of landing

and was regarded as the poor type. The absorbing movement comes too late and demand longer

contact time and lower takeoff velocity.

The pre-activation of landing is initiated by the centrally pre-programmed motor

commands of the required landing task (Dyhre-Poulsen et al., 1991), and increase the sensitivity









of the muscle spindle to enhance the stretch reflex (Gottlieb et al., 1981). The stretch reflex

enhances muscle stiffness, and thus, improved stiffness regulation could be attained by proper

pre-landing muscle activation (Allum & Mauritz, 1984).

GRF acting on the body during landing has been implicated in injuries to the lower

extremity, and controlling peak GRF at impact is directly associated with lowering the landing

stiffness. Apparently, the movement of lower extremity joints can influence the magnitude of the

impact forces. It is generally known that subjects who land on the balls of their feet and flex their

knee and ankle joints more have lower peak GRF. It was suggested that more joint movements

allowed the body mass to decelerate over a longer period of time thus the impact force and time

to peak force were decreased.

McNair et al. (2000) reported that GRF could be decreased immediately after instruction

of landing technique about the lower extremity joints kinematics. Additionally, they commented

that instruction could be more effective if the subjects' attention was drawn to distinct cues

(sound of soft landing impact) in the environment.

Cowling et al. (2003) assessed the efficacy of verbal instruction about landing techniques

to change impact force. They used instructions to increase knee flexion, to recruit hamstring

muscles earlier, and muscle bursts immediately before landing. Only the instruction to increase

knee flexion resulted in significantly greater knee flexion at initial ground contact and lower

GRF, compared with the other instruction conditions.

Park et al. (2006a) compared mechanical characteristics of soft landing between male and

female subjects performing vertical leaps and drop landings. They found significant increases in

ankle plantar flexion at touchdown, knee flexion motion during the landing phase, times for

maximal flexion of all joints, and decreases in peak vertical GRF, axial hip joint force, and knee









extensor moments in soft landing. Peak vertical GRF of males was significantly greater than that

of females in soft landing. Pre-landing extension of the distal joints in vertical leap and that of

proximal and distal joints in drop landing were suggested to activate the soft landing, with pre-

contracting the muscles to prepare the proper landing. They noted that the soft landing strategy

for males was fit for plyometrics and that of females was for absorbing type of landing.

Performance of Adapted Landing Biomechanics to Various Conditions

Landing is ideally suited for a performance study of weight-bearing segments of body, as

it requires large eccentric muscle forces during the control of joint flexion and mimics the

muscular stresses experienced during the landing phase. Therefore, numerous landing studies

have been conducted to test the performance of normal subjects in various conditions, as well as

to test subjects of ACL deficiency, ankle instability, fatigue of knee joint muscles, and so on.

A lesion of ACL is a major trauma of the knee joint, and mostly treated with a ligament

reconstruction. ACL reconstruction demands sophisticated rehabilitation program to regain the

original function before the injury. Decker et al. (2002) studied the biomechanics of lower

extremity joints in fully rehabilitated ACL reconstructed and healthy subjects performing drop

landing from a 60-cm height. At initial touchdown, the ACL group demonstrated greater hip

extension and ankle plantar flexion, compared to the healthy group. The peak vertical GRF was

not different between groups, but the ACL group delayed the time to its occurrence. The knee

extensors provided the major energy absorption function for both groups; however, the ACL

group performed 37% more ankle plantar flexor work and 39% less hip extensor work compared

with the healthy group. They concluded that the ACL group utilized a different landing strategy

adapted to the ACL reconstruction which employed the hip extensor muscles less and the ankle

plantar flexors more.









Doorenbosch et al. (2003) studied EMG activity of the quadriceps and hamstrings in

patients with ACL deficiency and healthy subjects during the vertical jump and landing. They

observed significantly higher co-contraction index of quadriceps/hamstring muscles in ACL

deficiency subjects. This was suggested that higher levels of co-contraction of quadriceps and

hamstrings during movements in ACL deficient subject help to compensate for the loss of the

passive constraining structure.

Caulfield et al. (2004) studied jump and landing performance in subjects with functional

instability of the ankle joint and normal control. The subjects performed five single leg jumps

onto a forceplate. They observed lateral and anterior force peaks occurred significantly earlier in

subjects with ankle instability. These changes occurred immediately post-impact and too early

for the reflex correction.

Madigan et al. (2003) studied the effect of lower extremity fatigue on the performance of

lower extremity joints during drop landing. EMG data were used to confirm fatigue in the

quadriceps muscles. They observed a decrease in peak vertical GRF, an increase in maximum

flexion of knee and ankle joints occurred early in a fatigue landing, while significant changes in

vertical GRF impulse and time to maximum knee flexion occurred during the middle or late

stages of a fatigue landing. For the first half of a fatigue landing, hip extensor impulse increased,

knee extensor impulse did not change, and ankle plantar flexor impulse decreased. These

changes were explained with a distal-to-proximal redistribution of extensor moments, which

allowed the larger proximal muscles to contribute more to resisting collapse during landing.

They also suggested active insufficiency of gastrocnemius, since it crosses the knee and shortens

as knee flexion increases. The shortening of this muscle diminishes the ability to produce plantar

flexor moment at the ankle. The increase in knee flexion upon landing during the first half of the









fatigue landing may have reduced the contribution of the gastrocnemius muscle to plantar flexor

impulse and resulted in an overall decrease in plantar flexor impulse. This decrease also has

dictated an increase in hip extensor impulse to generate the necessary support moment for

landing deceleration.

Park et al. (2006b) studied the effects of limited lower back motion on soft landing

mechanics of lower extremity joints with subjects wearing various low back braces, simulating

different low back stiffness conditions. They found that knee and hip joint flexions were

decreased and peak vertical GRF and axial hip force were increased in stiff brace condition,

compared with no brace and soft brace conditions. Additionally, typical sequential joint flexion

from distal to proximal was disrupted in females wearing the stiff brace, comparing to male

counterpart. They concluded that limited spinal motions by the brace caused alterations in knee

and hip joint motions during the landing phase and an increase in impact force. They emphasized

that lower back motion is one of the factors in determining landing mechanics, and a stiff lower

back is associated with a stiff landing. With limited trunk motion, more decelerating torque

might be concentrated on the knee extensors for females, and more axial loads be transmitted to

the proximal body segments of males during the soft landing.









Table 3-1. Average neutral zones for a functional spinal units in different regions of the spine
(0).
Flexion/Extension Lateral bending Axial rotation
Region (total) (one side) (one side)
CO C1 1.1 1.6 1.5
C1 C2 3.2 1.2 29.6
C3 C6 4.9 4 3.8
C7 T1/T11 T1 1.5 2.2 1.2
L1 L2 /L3 L4 1.5 1.6 0.7
L5 S1 3 1.8 0.4
Note: Cited from the work of White and Panjabi (1990).

Table 3-2. Representative ranges of motion of CO-C1-C2 complex (0).
Flexion/Extension Lateral bending Axial rotation
Level Reference (total) (one side) (one side)
CO-C1 Penning (1978) 35 10 0
Goel et al. (1988) 23 3.4 2.4
Panjabi et al. (1988) 24.5 5.5 7.2
C1-C2 Penning (1978) 30 10 70
Goel et al. (1988) 10.1 42 23.3
Panjabi et al. (1988) 22.4 6.7 38.9

Table 3-3. Representative ranges and limits of motion of the middle and lower cervical spines
(0).
Lateral bending Axial rotation
Region Flexion/Extention (one side) (one side)
Middle
C2-C3 10(5-16) 10(11-20) 3 (0-10)
C3-C4 15 (7-26) 11(9-15) 7(3- 10)
C4-C5 20(13-29) 11(0-16) 7(1-12)
Lower
C5-C6 20 (13- 29) 8 (0- 16) 7 (2- 12)
C6-C7 17 (6- 26) 7 (0- 17) 6 (2- 10)
C7-T1 9 (4- 7) 4 (0- 17) 2 (0- 7)
Note: Cited from the work of White and Panjabi (1990).

Table 3-4. Normal active cervical ranges of motion (in vivo) reported in the literatures (0).
Motion Dvorak (1992) Lanz (1999) AMA (2001)
Flexion/Extension 141.3 116.1 110
Lateral bending 91.4 84.1 90
Axial rotation 175 144.2 160
Rotation from flexion 81.4
Rotation from extension 165










Table 3-5. Representative ranges and limits of motion of the thoracic spine ().


Flexion/Extension
(total)
4 (3 5)
4(3 -5)
4(2-5)
4(2-5)
4(3 -5)
5 (2 -7)
6 (3 8)
6 (3 8)
6 (3 8)
9 (4 14)


T11- T12 12(6-20) 9
T12-L1 12(6- 20) 8
Note: Cited from the work of White and Panjabi (1990).


Lateral bending


(one side)
5 (5)
6(5 -7)
5 (3 -7)
6(5 -6)
6(5 -6)
6 (6)
6 (3 8)
6(4 -7)
6(4 -7)
7(3 10)


'(4- 13)
(5- 10)


Axial rotation
(one side)
9(14)
8(4- 12)
8(5- 11)
8(5- 11)
8(5- 11)
7(4- 11)
7(4- 11)
6 (6 7)
4 (3 5)
2 (2 3)
2(2- 3)
2(2- 3)


Table 3-6. Representative ranges and limits of motion of the lumbar spine ().
Level Flexion/Extension Lateral bending Axial rotation
L1-L2 12(5-16) 6(3-8) 2(1-3)
L2-L3 14(8-18) 6(3-10) 2(1-3)
L3- L4 15 (6-17) 8(4-12) 2(1-3)
L4-L5 16(9-21) 6(3-9) 2(1-3)
L5- S1 17(10-24) 3 (2- 6) 1 (0- 2)
Note: Cited from the work of White and Panjabi (1990).

Table 3-7. Comparison of lumbar compression loads in various trunk postures without external
loading.
Nachemson Takahashi et al. Ledet et al.
(1966; 1981) (2006) (2005)
% of % of x body % of
Trunk posture N standing N standing weight standing
Supine 300 25 1.9 95
Standing 700 100 645 100 2 100
Sitting 1000 140 2.5 125
Standing flexed 2.6* 130*
100 1277 198
200 1200 150 1922 298
300 2305 357
Sitting flexed 185 2.8 140
The trunk flexion angles were not specified in Ledet et al. (2005).


Level
T1 T2
T2 T3
T3 T4
T4 T5
T5 T6
T6 T7
T7 T8
T8 T9
T9 T10
T10-T11









Table 3-8. Average ranges of motion of the lumbar spine in normal walking and running in
different studies ().
Source Flexion/Extension Lateral bending Axial rotation
Walking
Crosbie et al. (1997) 3.5 9 4.5
(Lower thoracic) 2.5 7.0 4.0
(Pelvis) 3.5 6.0 4.0
Callaghan et al. (1999) 6.2 6.7 7.1
Van Herp et al. (2000) 2.3 4 6.6
Running
Schache et al. (2002) 13.3 18.5 23.0
(Pelvis) 7.6 10.6 13.9

Table 3-9. Peak compression loads to the lower lumbar level during walking (x BW).
Source Model Compression force
Cappozzo (1983) Single muscle equivalent model 1.0 2.5
Cromwell et al. (1989) EMG model 1.0
Khoo et al. (1995) Single muscle equivalent model 1.5 -2.1
Callaghan et al. (1999) EMG model 0.9 3.5
Inverse dynamic model 0.2 1.0
Nachemson (1964) Direct intradiscal pressure measurement 850 (N)

Table 3-10. Intradiscal pressure of low lumbar level during various activities.
Wilke et al. Rohlmann et al. Nachemson
(1999) (2001) (1966; 1981; 1987)
Intradiscal pressure Percentage of Percentage of
Position (MPa) standing (%) standing (%)
Lying supine 0.1 20 25
Relaxed standing 0.5 100 100
Standing with forward bending 1.1 216 150
Sitting without backrest 0.46 90 140
Sitting with maximum flexion 0.83 185
Standing up from a chair 1.1
Walking 0.53 0.65 130 121
Jogging 0.35 0.65 170
Jumping 240-380 157
Lifting 20 kg, squat lift 1.7 300
Lifting 20 kg, stoop lift 2.3 460 486



















NZ 4

-'4----Load---------v -

Maximum physiologic load Load
Neutral point

Figure 3-1. The load-displacement curve of a functional spinal unit (FSU) is generally nonlinear
and biphasic [neutral zone (NZ) and elastic zone (EZ)]. Range of motion (ROM) is
the sum of the neutral and elastic zones. Average flexibility coefficient (FC) is the
elastic zone divided by the maximum physiological load.









CHAPTER 4
RESULTS

Effects of Landing Technique

Overall landing characteristics were evaluated using different landing variables (Table 4-

1). Repeated measures MANOVA revealed significant main effects of gender (p<0.001) and

landing technique (p<0.001) for the landing variables. However, no significant interaction was

found between gender and landing technique (p=0.057) (Table C-1 in Appendix C).

Follow-up univariate contrast procedures revealed that t(LP), OP(KFA), and OP(HFA) increased

significantly and PVGRF decreased significantly when going from NL to SL condition. For

touchdown angles, both OTD(KFA) and OTD(HFA) were significantly more flexed when going from

NL to SL condition. Female subjects exhibited significantly greater OP(HFA) than male subjects

across both landing conditions (Table 4-1).

Kinematic characteristics of the spinal column after touchdown of drop landing were

evaluated with touchdown angle and extension ROM variables of each spinal region. Relative to

research questions Q1 and Q2 and associated hypotheses (la and 2a), repeated measures

MANOVA revealed a significant main effect of gender (p=0.011) and a significant interaction

between gender and landing technique (p=0.025) for touchdown angle variables, and significant

main effects of gender (p=0.013) and landing technique (p<0.001) for extension ROM variables.

However, no significant main effect of landing technique was found for touchdown angle

variables (p=0.236), and no significant interaction was found between gender and landing

technique for extension ROM variables (p=0.232) (Table C-1).

Follow-up univariate contrast procedures for touchdown angle variables revealed that

YTD(TL/L) of females was significantly greater than that of males (i.e., females demonstrated

significantly more extended thoracolumbar regional angle at touchdown than males), and a









significant interaction was found in YTD(TH/TL) (Table 4-2). The significant interaction of YTD(TH/TL)

between gender and landing technique means that males demonstrated more flexed thoracic

regional angle at touchdown from NL to SL condition, while females did not demonstrate

differences across different landing techniques (Figure 4-1).

Follow-up univariate tests for extension ROM variables identified a significant main effect

of gender for tP(TH/TL) and significant main effects of landing technique for YP(TH/TL) and YP(LC/TH).

The contrast procedures indicated that females exhibited significantly greater extension motion

in the thoracic region than males did, and extension motions in the thoracic and lower cervical

regions increased significantly from NL to SL condition (Table 4-2).

The overall kinematic characteristics of each spinal region reveal that the lumbar and

thoracolumbar regions exhibit flexion and the thoracic and lower cervical regions show

extension during the landing phase. However, the thoracolumbar region undergoes a short period

of extension followed by flexion during the landing phase in selected subjects (Figure 4-2).

Kinetic characteristics of L/S and C/T junctions after touchdown during drop landing were

evaluated with the kinetic variables at L/S and C/T junctions. Relative to research questions Q

and Q2 and associated hypotheses (lb and 2b), repeated measures MANOVA revealed a

significant main effect of landing technique for the kinetic variables at L/S junction (p<0.001),

and significant main effects of gender (p=0.031) and landing technique (p<0.001) for the kinetic

variables at C/T junction. There was also a significant interaction between gender and landing

technique for the kinetic variables at C/T junction (p=0.018). However, no significant gender

effect (p=0.403) or interaction (p=0.196) between gender and landing technique was found for

the kinetic variables at L/S junction (Table C-1). The univariate contrast procedures revealed that









all the kinetic variables at L/S junction decreased significantly from NL to SL condition (Table

4-3).

For the kinetic variables at C/T junction, all but ShF(C/T)ant decreased significantly from

NL to SL condition and ShF(C/T)ant and ShF(C/T)post demonstrated significant interactions

between gender and landing technique (Table 4-3). The significant interactions between gender

and landing technique for ShF(C/T)ant and ShF(C/T)post indicate that ShF(C/T)post of males was

greater than that of females during NL, but no gender difference was observed during SL

condition. On the other hand, ShF(C/T)ant of females was greater than that of males during NL,

but not different from each other during SL condition. Females demonstrated decreased

ShF(C/T)ant from NL to SL condition, but not for males (Figure 4-1).

Effects of Knee Joint Muscles Fatigue

Overall landing characteristics were evaluated using different landing variables. Repeated

measures MANOVA identified a significant main effect of gender (p=0.012) for the landing

variables. However, no significant fatigue effect (p=0.559) or interaction (p=0.104) between

gender and fatigue level was found for the landing variables (Table C-9). Follow-up univariate

tests revealed that female subjects exhibited a significantly greater OTD(KFA) than male subjects

across both landing conditions (i.e., females demonstrated more extended posture of knee joint at

touchdown than males) (Table 4-4).

Kinematic characteristics of the spinal column after touchdown of drop landing were

evaluated with touchdown angle and extension ROM variables. Relative to research questions

Q3 and Q4 and associated hypotheses (3a and 4a), repeated measures MANOVA revealed

significant main effects of gender for the touchdown angle (p=0.036) and extension ROM









(p=0.006) variables. However, there was not any significant fatigue effect or interaction for both

touchdown angle and extension ROM variables (Table C-9).

Univariate contrast procedures revealed that females demonstrated a significantly extended

posture of thoracolumbar region at touchdown (greater YTD(TL/L) in females than in males) and a

greater extension motion (greater YP(TH/TL) in females than in males) during landing phase than

males did (Table 4-5).

The overall kinematic characteristics of each spinal region reveal that the lumbar region

exhibits flexion and thoracic and lower cervical regions show extension during the landing

phase. However, the thoracolumbar region undergoes a short period of extension followed with

flexion during the landing phase in selected subjects (Figure 4-3).

Kinetic characteristics of L/S and C/T junctions after touchdown of drop landing were

evaluated with the kinetic variables of L/S and C/T junctions. Relative to research questions Q3

and Q4 and associated hypotheses (3b and 4b), repeated measures MANOVA revealed

significant interactions between gender and fatigue level for the kinetic variables at L/S junction

(p=0.033) and C/T junction (p=0.043) (Table C-9).

Follow-up univariate tests identified significant interactions of gender x fatigue level for

AxF(L/S), ShF(L/S)post, ExtM(L/S), and AxF(C/T) (Table 4-6). The significant interaction for

AxF(L/S) indicated that females exhibited increased AxF(L/S) from SL1 to SL2 condition, while

males did not (Figure 4-4). The significant interaction of ShF(L/S)post revealed that females

exhibited increased ShF(L/S)post from SL1 to SL2 condition, while males did not. Also,

ShF(L/S)post was greater in males than in females for SL1, but was greater in females than in

males for SL2 condition. The significant interaction of AxF(C/T) indicated that females

exhibited increased AxF(C/T) from SL1 to SL2 condition, while males did not. Also, AxF(C/T)









was greater in males than in females for SL1, but was greater in females than in males for SL2

condition. The significant interaction of ExtM(L/S) revealed that females exhibited increased

ExtM(L/S) from SL1 to SL2 condition, while males did not. Lastly, ExtM(L/S) was greater in

males than in females for SL1, but was greater in females than in males for SL2 condition.









Table 4-1. Collapsed mean and SD values of different landing variables for different genders
and landing techniques.
NL SL Landing x
Landing variables Mean (SD) Mean (SD) Gender: p Landing: p Gender: p
PVGRF (N.kg-1) 23.91 (3.87) 15.77 (3.03) 0.877 <0.001* 0.723
Male 23.67 (3.22) 15.83 (2.91)
Female 24.15 (4.55) 15.70 (3.27)
t(LP) (S) 0.222 (0.084) 0.355 (0.139) 0.148 <0.001* 0.536
Male 0.203 (0.071) 0.320 (0.075)
Female 0.241 (0.095) 0.390 (0.179)
OTD(KFA) (0) -7.3 (6.8) -10.5 (6.4) 0.063 <0.001* 0.268
Male -9.2(4.6) -13.2(5.5)
Female -5.3 (8.2) -7.8 (6.2)
9p(KFA) (0) 53.0 (10.4) 69.9 (14.2) 0.087 <0.001* 0.131
Male 51.7 (12.6) 71.8 (16.0)
Female 54.2 (7.9) 68.0 (12.6)
OTD(HFA) (0) -61.5 (12.0) -66.6 (12.3) 0.853 <0.001* 0.747
Male -60.9 (10.7) -66.4 (11.7)
Female -62.1 (13.6) -66.9 (13.4)
9p(HFA) (0) 33.3 (13.7) 51.2 (10.8) 0.001* <0.001* 0.113
Male 24.8 (11.5) 47.0 (7.2)
Female 41.7 (10.2) 55.5 (12.3)
Note: NL (self-selected normal landing), SL (soft landing), PVGRF (peak vertical GRF), t(LP) (time for
landing phase), OTD(KFA) (knee flexion angle at touchdown),Op(KFA) (ROM of knee flexion angle from
touchdown to initial peak of knee flexion), OTD(HFA) (hip flexion angle at touchdown), OP(KFA) (ROM of hip
flexion angle from touchdown to initial peak of hip flexion), (significant in univariate tests)










Table 4-2. Collapsed mean and SD values of touchdown angle and extension ROM of each
spinal region for different genders and landing techniques.
Kinematic NL SL Landing x
variables Mean (SD) Mean (SD) Gender: p Landing: p Gender: p
Touchdown angle of each spinal region


YTD(L/S) (0o)
Male
Female

YTD(TL L) (0)
Male
Female

YTD(TH/TL) (0o)
Male
Female

YTD(LC/TH) (0o)
Male
Female


15.1 (8.8)
13.3 (9.0)
17.0 (8.5)
11.9 (7.8)
7.6 (5.1)
16.1 (7.8)
-21.9 (6.9)
-23.1 (6.2)
-20.7 (7.6)
-1.6 (12.4)
-1.4 (10.6)
-1.7 (14.5)


Extension ROM of each spinal region
YP(L S) (0) 0.5 (1.3)
Male 0.2 (0.3)
Female 0.9(1.7)


YP(TL/L) (0o)
Male
Female

YP(TH/TL) (0o)
Male
Female

YP(LC/TH) ()
Male
Female


2.3 (2.0)
1.8 (1.9)
2.9 (2.0)
4.6 (3.7)
3.3 (2.8)
6.0 (4.2)
10.3 (7.8)
10.0 (9.7)
10.5 (5.7)


15.0 (9.0)
12.6 (9.5)
17.4 (8.2)
11.9 (8.7)
7.2 (6.1)
16.7 (8.4)
-22.6 (7.8)
-25.1 (8.2)
-20.1 (6.9)
-0.3 (13.4)
0.1 (14.6)
-0.8 (12.8)


0.0 (0.1)
0.0 (0.1)
0.1 (0.1)
2.0 (2.3)
1.2 (1.7)
2.8 (2.6)
8.0 (5.1)
5.4 (3.9)
10.6 (4.9)
16.3 (10.4)
14.8 (10.8)
17.7(10.1)


0.215


0.003*


0.196



0.91




0.167



0.070



0.012*



0.591


0.85


0.933


0.219



0.451




0.061



0.392



< 0.001*



0.006*


0.433


0.28


0.032*



0.844




0.179



0.526



0.057



0.557


Note: NL (self-selected normal landing), SL (soft landing), yTD(L/S) (lumbar regional angle at touchdown),
YTD(TL/L) (thoracolumbar regional angle at touchdown), YTD(TH/TL) (thoracic regional angle at touchdown),
YTD(LC/TH) (lower cervical regional angle at touchdown), YP(L/ S) (extension ROM of lumbar region from
touchdown to initial peak during landing phase), YP(TL/L) (extension ROM of thoracolumbar region from
touchdown to initial peak during landing phase), YP(TH/TL) (extension ROM of thoracic region from
touchdown to initial peak during landing phase), YP(LC/TH) (extension ROM of lower cervical region from
touchdown to initial peak during landing phase), (significant in univariate tests)










Table 4-3. Collapsed mean and SD values of kinetic variables at L/S and C/T junctions for
different genders and landing techniques.
NL SL Landing x
Kinetic variables Mean (SD) Mean (SD) Gender: p Landing: p Gender: p


L/S junction
AxF(L/S) (N-kg-')
Male
Female
ShF(L/S)ant (N.kg-1)
Male
Female
ShF(L/S)post (N.kg-)
Male
Female
FlxM(L/S)
(N.m.kg-1.BH-1)
M
F
ExtM(L/S)
(N-m-kg-1.BH-1)
Male
Female
C/T junction
AxF(C/T) (N-kg-')
Male
Female
ShF(C/T)ant (N.kg-')
Male
Female
ShF(C/T)post (N-kg-')
Male
Female
FlxM(C/T)
(N-m-kg-1.BH-')
Male
Female
ExtM(C/T)
(N.m.kg-1.BH-1)
Male
Female


8.52 (2.70)
7.72 (2.43)
9.33 (2.81)
1.73 (1.01)
2.16 (1.24)
1.30 (0.45)
9.73 (3.54)
10.48 (3.21)
8.97 (3.81)

1.37 (0.44)
1.41 (0.37)
1.33 (0.50)

3.44 (1.15)
3.68 (1.06)
3.19 (1.23)


4.98 (4.07)
4.36 (4.0)
5.60 (4.20)
3.22 (2.28)
2.35 (1.84)
4.08 (2.42)
5.54 (1.28)
6.26 (1.42)
4.83 (0.53)

2.54 (0.91)
2.40 (0.77)
2.68 (1.04)

3.85 (1.33)
4.21 (1.4
3.50 (1.19)


5.77 (1.93)
5.55 (1.77)
5.99 (2.13)
0.83 (0.51)
0.82 (0.55)
0.84 (0.48)
4.35 (2.32)
4.66 (2.45)
4.05 (2.23)

0.91 (0.43)
0.95 (0.36)
0.87 (0.51)

1.94 (0.70)
2.02 (0.64)
1.85 (0.77)


1.65 (2.28)
1.96 (2.25)
1.34 (2.36)
2.56 (1.73)
2.56 (1.58)
2.56 (1.94)
4.26 (0.81)
4.35 (0.97)
4.17 (0.65)

1.75 (0.81)
1.80 (0.70)
1.70 (0.93)

2.24 (0.71)
2.25 (0.60)
2.24 (0.83)


0.182


0.085


0.297



0.583



0.31




0.749



0.197



0.026*



0.758



0.302


<0.001*


<0.001*


<0.001*



<0.001*



<0.001*




0.001*



0.117



<0.001*



<0.001*



<0.001*


0.268


0.022


0.476



0.961



0.418




0.303



0.042*



0.001*



0.342



0.145


Note: NL (self-selected normal landing), SL (soft landing), AxF (peak axial compressive force), ShFant(or
post) (peak ant. or post. shear force), FlxM (peak flexor moment), ExtM (peak extensor moment), (L/S) (for
lumbosacral junction), (C/T) (for cervicothoracic junction), (significant in univariate tests)









Table 4-4. Collapsed mean and SD values of different landing variables for different genders
and fatigue levels.
Landing SL1 SL2 Fatigue x
variables Mean (SD) Mean (SD) Gender: p Fatigue: p Gender: p
PVGRF (N.kg-1) 15.72 (3.03) 16.47 (3.43) 0.312 0.087 0.007
Male 15.8(3.04) 15.2(3.36)
Female 15.62 (3.19) 17.99 (2.98)
t(LP) (S) 0.335 (0.147) 0.342 (0.163) 0.304 0.628 0.666
Male 0.307 (0.088) 0.308 (0.093)
Female 0.368 (0.197) 0.382 (0.220)
OTD(KFA) () -10.2(6.7) -11.0(7.1) 0.008* 0.322 0.281
Male -13.2 (5.8) -14.6 (6.5)
Female -6.6 (6.2) -6.6 (5.2)
9p(KFA) (0) 70.5 (15.0) 71.7 (15.9) 0.271 0.44 0.112
Male 72.8 (16.2) 75.9 (18.0)
Female 67.6 (13.8) 66.5 (11.9)
OTD(HFA) (0) -65.2 (10.9) -64.1 (15.5) 0.094 0.608 0.648
Male -66.3 (11.8) -64.2 (18.4)
Female -64.0 (10.2) -63.9 (12.2)
9p(HFA) (0) 52.0 (10.8) 53.6 (12.9) 0.094 0.264 0.749
Male 48.1 (6.3) 50.0 (8.5)
Female 56.8 (13.3) 57.9 (16.3)
Note: SL1 (soft landing before the fatigue procedures), SL2 (soft landing after the fatigue procedures),
PVGRF (peak vertical GRF), t(LP) (time for landing phase), OTD(KFA) (knee flexion angle at
touchdown),Op(KFA) (ROM of knee flexion angle from touchdown to initial peak of knee flexion), OTD(HFA)
(hip flexion angle at touchdown), OP(KFA) (ROM of hip flexion angle from touchdown to initial peak of hip
flexion), (significant in univariate tests)










Table 4-5. Collapsed mean and SD values of touchdown angle and extension ROM of each
spinal region for different genders and fatigue levels.
Kinematic SL1 SL2 Fatigue x
variables Mean (SD) Mean (SD) Gender: p Fatigue: p Gender: p
Touchdown angle of each spinal region
YTD(L/S) () 15.1 (9.0) 14.0 (9.3) 0.473 0.343 0.558
Male 13.5 (10.0) 13.1 (9.8)
Female 17.0 (7.7) 15.1 (9.1)

YTD(TL/L) () 11.5 (9.0) 11.4 (9.9) 0.005* 0.988 0.545
Male 6.8 (6.4) 6.5 (7.0)
Female 17.0 (8.7) 17.3 (9.8)


YTD(TH/TL) ()
Male
Female

YTD(LC/TH) (0o)
Male
Female


-22.9(8.1)
-25.7 (8.2)
-19.6 (7.1)
0.3 (13.6)
0.7 (15.5)
-0.1 (11.6)


-22.9 (7.8)
-25.7 (7.0)
-19.6 (7.6)
3.2 (13.4)
1.6 (15.2)
5.0 (11.5)


Extension ROM of each spinal region
YP(LS) () 0.0(0.1) 0.4(1.2)
Male 0.0 (0.0) 0.0 (0.0)
Female 0.0(0.1) 0.9(1.7)


YP(TL/L) ()
Male
Female

YP(TH/TL) ()
Male
Female


2.1 (2.4)
1.3 (1.8)
3.1 (2.7)
7.8 (5.5)
5.2 (4.0)
11.0 (5.5)


1.5 (1.5)
1.1 (1.2)
1.9 (1.6)
8.1 (5.5)
5.7 (4.4)
11.0 (5.5)


0.068



0.819




0.074



0.085



0.013*


0.924



0.056




0.072



0.083



0.624


0.993



0.184




0.087



0.212



0.624


YP(LC/TH) () 16.6 (10.6) 19.2 (12.3) 0.776 0.355 0.042
Male 15.0 (11.2) 22.0 (14.0)
Female 18.6 (10.1) 15.9 (9.5)
Note: SL1 (soft landing before the fatigue procedures), SL2 (soft landing after the fatigue procedures),
YTD(L/S) (lumbar regional angle at touchdown), YTD(TL/L) (thoracolumbar regional angle at touchdown),
YTD(TH/TL) (thoracic regional angle at touchdown), YTD(LC/TH) (lower cervical regional angle at touchdown),
YP(L/S) (extension ROM of lumbar region from touchdown to initial peak during landing phase), YP(TL/L)
(extension ROM of thoracolumbar region from touchdown to initial peak during landing phase), YP(TH/TL)
(extension ROM of thoracic region from touchdown to initial peak during landing phase), YP(LC/TH)
(extension ROM of lower cervical region from touchdown to initial peak during landing phase), *
(significant in univariate tests)









Table 4-6. Collapsed mean and SD values of kinetic variables at L/S and C/T junctions for
different genders and fatigue levels.
SL1 SL2 Fatigue x
Kinetic variables Mean (SD) Mean (SD) Gender: p Fatigue: p Gender: p
L/S junction
AxF(L/S) (N-kg-') 5.79 (2.01) 5.76 (1.86) 0.063 0.9 0.019*
Male 5.46 (1.82) 4.80 (1.38)
Female 6.19(2.26) 6.91 (1.73)
ShF(L/S)ant (N-kg-1) 0.84 (0.53) 0.85 (0.41) 0.763 0.877 0.185
Male 0.88 (0.53) 0.76 (0.40)
Female 0.80 (0.54) 0.95 (0.43)
ShF(L/S)post (N-kg-') 4.40 (2.48) 5.33 (2.45) 0.703 0.002 0.001*
Male 4.71 (2.55) 4.66 (2.39)
Female 4.02 (2.46) 6.13 (2.39)
FlxM(L/S)
(N.m.kg-'.BH-1) 0.93 (0.45) 0.98 (0.42) 0.709 0.576 0.961
Male 0.96 (0.37) 1.01 (0.40)
Female 0.90 (0.55) 0.94 (0.47)
ExtM(L/S)
(N.m.kg-1.BH-) 1.93 (0.73) 2.21 (0.68) 0.7 0.01 0.011*
Male 2.02 (0.67) 2.02 (0.63)
Female 1.82 (0.83) 2.44 (0.70)
C/T junction
AxF(C/T) (N-kg-') 1.74 (2.40) 2.21 (2.04) 0.541 0.096 0.002*
Male 2.04 (2.33) 1.42 (1.28)
Female 1.38(2.56) 3.17(2.42)
ShF(C/T)ant (N-kg-') 2.71 (1.71) 2.66 (1.72) 0.398 0.907 0.715
Male 2.51 (1.64) 2.35 (1.48)
Female 2.95 (1.84) 3.04 (1.98)
ShF(C/T)post (N-kg-') 4.19 (0.86) 4.53 (0.90) 0.438 0.036 0.699
Male 4.28 (0.98) 4.69 (0.93)
Female 4.07 (0.71) 4.35 (0.88)
FlxM(C/T)
(N-m-kg-1.BH-1) 1.84 (0.80) 1.94 (0.76) 0.782 0.474 0.787
Male 1.78 (0.73) 1.91 (0.63)
Female 1.90 (0.92) 1.97 (0.92)
ExtM(C/T)
(N.m.kg-'.BH-1) 2.21 (0.72) 2.53 (0.70) 0.454 0.022 0.054
Male 2.24 (0.63) 2.30 (0.59)
Female 2.17 (0.86) 2.79 (0.77)
Note: SL1 (soft landing before the fatigue procedures), SL2 (soft landing after the fatigue procedures),
AxF (peak axial compressive force), ShFant(or post) (peak ant. or post. shear force), FlxM (peak flexor
moment), ExtM (peak extensor moment), (L/S) (for lumbosacral junction), (C/T) (for cervicothoracic
junction), (significant in univariate tests)














* /TD(THfrL) in Male
r' .. .l in Female


*ShF(CIT)oP, in Male
ShF(C/T)oP, in Female


* ShF(C/T).n in Male
E ShF(C/T),nt in Female


Figure 4-1. Significant interactions of the touchdown angle and C/T kinetic variables between
gender and landing technique: (A) YTD(TH/TL) was decreased from NL to SL condition
in males, but not decreased in females (thoracic regional angle becomes more flexed
from NL to SL condition in males, but not in females), (B) ShF(C/T)post of males was
greater than that of females during NL, but not different from each other during SL
condition; ShF(C/T)ant of females was greater than that of males during NL, but not
different from each other during SL condition; ShF(C/T)ant was decreased from NL to
SL condition in females, but not decreased in males. Note: NL (self-selected normal
landing), SL (soft landing).

















D-20-

-40 -

-60-

-80-
40
2n

0--

-20-
S-40--

-60-

-80-

-100-




Figure 4-2


. Representative kinematics of each spinal region (YL/S: lumbar regional angle relative
to the sacral region, YTL/L: thoracolumbar regional angle relative to the lumbar region,
YTH/TL: thoracic regional angle relative to the thoracolumbar region, YLC/TH: lower
cervical regional angle relative to the thoracic region), knee flexion angle (OKFA) and
hip flexion angles (OHFA) in one male and one female subjects during NL and SL
conditions. For the direction of motion, negative slope represents the flexion motion.
Flexion of a spinal region means the forward rotation of a region relative to the
adjacent lower region. Note: LP (landing phase), NL (self selected normal landing),
SL (soft landing).


0KFA OHFA 7L/S 7TL/L -'TH.TL 7LCYTIH









40 60-
LP Male in SL1 Female in SL
2 0 ,, 4 0 L P F l 1
20
0-\

--62, i ; -603'" /
20 Time. -10- 0


-100_ -120 -
-120- -140-
hip feio Male inSl~~L one .e Female in SL2
40 40 ,-q--


2o0 -" 1 /;8 -2
-40


-80_ -100 /
-100 -120 '
-120- -140
0KFA 0HFA 'US '/TUL "/xTwL 7LCH

Figure 4-3. Representative kinematics of each spinal region (YL/s: lumbar regional angle relative
to the sacral region, YTL L: thoracolumbar regional angle relative to the lumbar region,
YTI/TL: thoracic regional angle relative to the thoracolumbar region, YLC/TH: lower
cervical regional angle relative to the thoracic region), knee flexion angle (OKFA) and
hip flexion angles (OHFA) in one male and one female subjects during SL1 and SL2
conditions. For the direction of motion, negative slope represents the flexion motion.
Flexion of a spinal region means the forward rotation of a region relative to the
adjacent lower region. Note: LP (landing phase), SL1 (soft landing before the fatigue
procedures), SL2 (soft landing after the fatigue procedures).









0 AxF(L/S) in Male
(A) AxF(L/S) in Female

Qr^i^ n


* AxF(C/T) in Male
AxF(C/T) in Female


X ExtM(L/S) in Male

X ExtM(L/S) in Male
* ExtM(L/S) in Female


Figure 4-4. Significant interactions of the kinetic variables at L/S and C/T junctions between
gender and fatigue level: (A) AxF(L/S) was increased from SL1 to SL2 condition in
females, but not increased in males; ShF(L/S)post was increased from SL1 to SL2
condition in females, but not increased in males; AxF(C/T) was increased from SL1
to SL2 in females, but not increased in males, (B) ExM(L/S) was increased from SL1
to SL2 condition in females, but not increased in males. Note: SL1 (soft landing
before the fatigue procedures), SL2 (soft landing after the fatigue procedures).









CHAPTER 5
DISCUSSION

Effects of Landing Technique

Previous studies on landing mechanics focused mostly on biomechanical characteristics of

lower extremities. Majority of these studies focused on defining injury mechanism (Fagenbaum

& Darling, 2003; Wikstrom et al., 2006), gender differences (Kernozek et al., 2005) and

conditional variability (Schot et al., 2002) associated with knee and ankle joints. To our best

knowledge, no attempt has been made to evaluate the mechanical characteristics of upper body

segments above the hip joints during drop landings.

The landing variables used in the current study identified the overall landing characteristics

including knee and hip joints. The landing technique used in soft landing significantly decreased

the PVGRF, extended the landing phase, exhibited flexed landing postures of knee and hip

joints, and exhibited more flexion motion of knee and hip joints. Additionally, female subjects

demonstrated a greater degree of hip flexion than male subjects, and this may indicate that

females demonstrate absorbing type landing without any different instructions.

Horita et al. (2002) proposed two types of landing performances with regard to the pre-

landing motion of the knee joint. The proper pre-landing movement was characterized by knee

flexion just before touchdown, which is associated with a high initial joint stiffness coupled with

high joint power. This was called bouncing type in their study, which is similar to plyometrics.

On the other hand, an inadequate pre-landing movement, associated with incomplete knee

flexion induced subsequent deep-knee flexion after touchdown, was called the absorbing type of

landing and was regarded as a poor strategy. The absorbing movement comes too late and

demands longer contact time and lower takeoff velocity. A recent study by Park et al. (2006a)

revealed that the soft landing strategy used by males was close to the bouncing type which is fit









for the effective subsequent movements, and that of females was close to absorbing type of

landing which is just pursuing the dissipation of increased loads from ground impact.

The lumbar region was in a flexion posture immediately before touchdown and reached its

maximal flexion around the end of landing phase, and this was followed by a rapid extension to

neutral after the landing phase. For the purpose of discussion, the neutral angle of a spinal region

is defined as the spinal regional angle recorded during a standing posture at the end of a landing

trial. Flexion motion of a spinal region refers to forward rotation of the spinal region relative to

the lower adjacent spinal region. The thoracic region started an extension motion after

touchdown followed by a flexion to neutral toward the end of landing phase. However, the

thoracic spine moved more gradually compared to the movement of lumbar region. The

thoracolumbar region had a short duration of slight extension around touchdown and reached its

initial peak at the beginning of the landing phase followed by a flexion motion similar to the

lumbar region. The lower cervical region exhibited an extension motion around touchdown, and

reached its full extension during the landing phase followed by a rapid flexion to neutral. Most

spinal regions demonstrated biphasic motions (i.e., extension followed by flexion or flexion

followed by extension), but the thoracolumbar region demonstrated multiphasic motion during

the landing phase of drop landing.

The biphasic pattern of spinal motion has been observed during walking. Crosbie et al.

(1997) studied the patterns of spinal motion during walking using a model including upper and

lower trunks, lumbar and pelvis segments. They used a surface marker in each spinal segment on

the back surface of the subjects. The pattern of flexion/extension of each segment was generally

biphasic throughout the gait cycle. The pelvis rotated into negative pelvic tilt at heel strike, and

this was followed by a counter-motion to a maximum positive pelvic tilt in the single support









phase. The lumbar spine showed maximum flexion at heel strike, and this was followed by a

rapid extension to neutral until the single support phase. The lower thoracic segment extended

maximally at heel strike, and returned to a neutral at mid-stance, then extended through the late

stance phase.

Kinematics of the spinal column after touchdown of drop landing was addressed with

touchdown angle and extension ROM variables. Relative to research question Q (any gender

differences in spinal kinematics?), the null hypothesis la was rejected by the finding of gender

differences in thoracolumbar touchdown angle and extension ROM of the thoracic region.

Females demonstrated a more extended posture in the thoracolumbar region at landing and

more extension motion in the thoracic region during the landing phase compared to males. The

greater thoracic region extension may be related to the greater hip flexion in females (balance

control). An extended lower extremity landing posture and more flexion motion of lower

extremity joints are considered as specific landing characteristics in females (Decker et al.,

2003). According to the work-energy relationship, the average vertical GRF experienced by the

subject during a drop landing depends on the vertical displacement of whole body CG (center of

gravity) during the landing phase. Because an increased flexion motion of the hip joint will lower

the body CG, the action of the thoracic region may compensate the lowered CG during the

landing phase. As a result, the energy absorbing procedure in the landing phase may become less

effective in females because of the movements of spinal column against the energy absorption

such as the extension motion of the thoracic region.

Relative to research question Q2 (any changes in spinal kinematics by the different landing

technique?), the null hypothesis 2a was rejected by the finding of significant changes in YP(TH/TL),

YP(LC/TH), and YTD(TH/TL) across different landing techniques. When going from NL to SL









condition, the thoracic and lower cervical regions exhibited more extension motion. Males

exhibited more flexed landing posture in the thoracic region from NL to SL, while females did

not or showed extended posture.

Active movement of each spinal region in both landing conditions may indicate that the

spinal column is actively involved in energy absorption during the landing phase. Without

analyzing the flexion ROM in each spinal region (only extension ROMs were analyzed), the

flexion ROM of each spinal region could not be apparently differentiated with each other.

However, the very small initial peaks of extension observed in the lumbar region across different

landing techniques indicate that the lumbar region exhibited flexion motion in most cases.

Additionally, during a soft landing, the thoracic and lower cervical regions demonstrated active

extension motions against the lumbar motion. This may suggest that the entire spinal column is

involved in energy absorbing procedure during the NL condition, while only part of spinal

column is involved during the SL condition. The partial involvement is more apparent in females

during the soft landing. Futhermore, the gender differences in thoracic landing posture during

soft landings may suggest that the spinal column is less involved in energy absorbing procedure

of females during soft landings (see Figure 4-4).

The results also suggest that the thoracolumbar region could be highly stressed by the

simultaneous motions of thoracic and lower cervical extension and lumbar flexion during the

landing phase of soft landing and is more pronounced in females. Schache et al. (2003)

performed a kinematic study of the lumbo-pelvic-hip complex during running to define the

gender differences. They found that females displayed a shorter stance time, swing time, stride

time and stride length, and a higher stride rate than males. Mean waveforms were different in the

peak-to-peak oscillations and the offset of pelvis anterior/posterior tilt. Females displayed greater









amplitudes of lumbar spine lateral bend and axial rotation, pelvis anterior/posterior tilt, obliquity

and axial rotation, and hip adduction/abduction than their male counterparts. The mean positions

of anterior pelvic tilt across the running cycle were 20.20 for females and 16.90 for males. The

prevalence of pelvic-femoral stress fractures in female runners might be explained by these

findings (Bennell et al., 1996; Pavlov et al., 1982). The greater amplitude of lumbar spinal

movements in females during drop landings is similar to the findings observed in running.

The increased extension movement of the thoracic region relative to the thoracolumbar

region from NL to SL condition suggests that the soft landing procedure intended to decrease

lower extremity loads may be a risk factor for developing thoracolumbar or upper lumbar

degeneration and spondylolysis in physically active individuals. Degenerative disc disease and

spondylolysis are the most common structural abnormalities associated with low-back pain in

athletes. Disc degeneration appears to be influenced by the type and intensity of the sport.

Videman et al. (1995) demonstrated that weight lifters have a higher rate of and more severe

degenerative changes in the upper lumbar spine, whereas back problems in soccer players are

almost exclusively in the L4 to S levels. Cappozzo et al. (1985) found that, when a person

performed half-squat exercises with weights approximately 1.6 times body weight, compressive

loads across the L3/L4 motion segment were about 10 times body weight. The prevalence of

spondylolysis in athletes is variable, but some sports appear to be associated with a higher

prevalence rate. Rossi and Dragoni (1990) reported a rate of 43% in divers, 30% in wrestlers,

and 23% in weight lifters. Although the exact mechanism for the development of spondylolysis

is not known, there is some suggestion that it may be a fatigue fracture following repeated

hyperextension.









Joint resultants at L/S and C/T junctions after touchdown of drop landing were addressed

with the L/S and C/T kinetic variables. Relative to research question Q (any gender differences

in L/S or C/T kinetics?), the null hypothesis lb was rejected by the finding of significant gender

differences. Posteriorly directed shear force of the trunk segment at C/T junction was

significantly greater in male subjects than in female subjects during NL, but no gender difference

for the SL condition was found. Females demonstrated decreased ShF(C/T)ant from NL to SL

condition, while males did not demonstrate any changes in ShF(C/T)ant across landing techniques

(see Figure 4-1). ShF(C/T)ant of females was greater than that of males during NL, but no

difference between genders was found for the SL condition. Increased posteriorly directed shear

force in males and increased anteriorly directed shear force in females for the NL condition can

place much stress on supporting anatomical structures.

The cervical spine can be injured due to increased shear forces at the lower cervical region

by a whiplash. During a whiplash, hyperextension of head and neck is the basic mechanism for

cervical spine injury and it commonly occurs in the rear-end impact in motor vehicle accidents

(Luan et al., 2000). Studies of the natural history of whiplash-associated disorders have

suggested that chronic pain with continued symptoms develops in 6-33% of acutely injured

patients (Hildingsson & Toolanen, 1990). Previous biomechanical studies have focused on injury

mechanisms of the cervical facet joints and the intervertebral discs as potential structures to

develop whiplash-associated disorders. Following a whiplash, the lower cervical spine

experiences complex loading consisting of an extension moment, posterior shearing and

compressive forces. This loading pattern has been hypothesized to injure the cervical

intervertebral disc and facet joints (Panjabi et al., 2004).









Facet joints of the lower cervical spine can be compressed by the increased posterior shear

force, and a capsular ligament strain can be distracted by the increased anterior shear force

during the NL condition. Pearson et al. (2004) evaluated peak facet joint compression and

capsular ligament strain using a whole cervical spine specimen with muscle force replication and

a bench-top trauma sled to simulate whiplash of increasing severity. Peak facet joint compression

was greatest at C4/C5 and reached over physiologic limits during the 3.5g simulation (low-level

acceleration). Capsular ligament strains exceeded the physiologic strains at 6.5g and were largest

at C6/C7 during the 8g simulation (high-level acceleration). They concluded that peak facet joint

compression occurred at maximum intervertebral extension, whereas peak capsular ligament

strain was reached as the facet joint was returning to its neutral position after the maximum

intervertebral extension. The greater degree of posteriorly directed shear force in males than in

females may indicate that male subjects could be in risk of increased facet joint compression,

while females could be in risk of increased capsular ligament strain due to the increased

anteriorly directed shear force during NL condition.

Panjabi and co-workers (2004) have studied a whole cervical spine model on a bench-top

sled with muscle force replication and a surrogate head to simulate rear-end impact. They

underwent standard flexibility testing to determine the sagittal disc deformation with the various

acceleration of the bench. They observed the greatest strain in the posterior 1500 fibers running

posterosuperiorly at C5/C6. They also observed increased disc shear strain at the posterior region

and increased axial deformation at the anterior region of the disc at C5/C6. They concluded that

the cervical intervertebral disc is at risk for injury during a whiplash because of excessive 1500

fiber strain, disc shear strain, and anterior axial deformation. The injury of the cervical spine by









facet joint distraction may occur at the level lower than the cervical injury by the facet

compression.

Relative to research question Q2 (any changes in L/S or C/T kinetics by the landing

technique?), the null hypothesis 2b was rejected by the finding of significant changes in all

kinetic variables except for ShF(C/T)ant across different landing techniques. During the SL, all

kinetic variables except for ShF(C/T)ant decreased significantly comparing to NL condition. This

may indicate that the soft landing technique used in the current study is an effective way to

decrease the overall loads applied to L/S and C/T junctions. However, ShF(C/T)ant was not

decreased by the soft landing.

The greatest loads at L/S junction were AxF(L/S) and ShF(L/S)post, while the ShF(C/T)post

was the greatest load at C/T junction during drop landing. Callaghan et al. (1999) conducted a

biomechanical study using two models to estimate loads applied to the L4/L5 level during

walking task: linked segment model with EMG technique and rigid segment model with inverse

dynamic technique. The joint loading at L4/L5 calculated by the EMG model resulted in large

increases in the maximum compressive forces (3.5 times of body weight), compared with the

joint reaction forces calculated using inverse dynamics (1.0 times of body weight). Including the

muscular component resulted in a more than three-fold increase in joint load. However, the joint

shear forces (anterior/posterior, lateral) obtained using the two techniques were quite similar.

The peak compressive axial forces at L/S junction were 8.5 and 5.8 times of body mass for NL

and SL conditions, respectively. The posteriorly directed shear force at L/S junction from the

current study (9.7 x body mass in NL, 4.5 x body mass in SL) was much higher than the values

for their walking trials (EMG model: 0.18 x BW, Inverse dynamics model: 0.19 x BW). In the

study of Callahan et al. (1999), there was a peak flexor moment at heel contact followed by a









peak extensor moment around toe-off. During a faster speed gait, the flexion/extension moment

at L4/L5 shifted to the extension side and demonstrated a high extensor moment around toe-off.

Callahan et al. concluded that the loads and motions for the lumbar spine during gait depended

on the walking speed. Increasing walking speed increases the lumbar spine ROM, activation of

spinal and trunk muscles, and anterior/posterior shear forces. Likewise, the loads and motions for

the lumbar spine during drop landing depend on the landing technique which controls the

involvement of body segments including lower extremity joints and spinal regions.

One of the limitations of this study is that sagittal spinal kinematics relative to the adjacent

spinal region based on the spinal skin markers cannot precisely determine the vertebral motions

in each spinal region. Additionally, there might be some errors in kinetic measures, because the

locations of joint center and spinal junction (used for joint resultants computations) were

estimated using surface markers instead of determined using radiographic imaging techniques.

Mechanical characteristics of the lower extremities were assumed to be symmetrical and only the

GRF data collected from the left leg were used to determine joint resultants at the L/S and C/T

junctions. Despite these limitations, the results from this study provide insight into the spinal

movement during two different landing techniques.

Effects of Knee Joint Muscles Fatigue

The purpose of the second part of this study was to determine whether spinal mechanics

were affected by lower extremity fatigue during drop landings. The overall landing mechanics

did not change significantly by the presence of knee joint muscles fatigue in the current study.

Instead, a significant gender difference in landing posture was found in the knee joint. Females

landed with a more extended knee joint posture than males which is consistent with the findings

reported in a previous study (Decker et al., 2003). These authors found that females

demonstrated a more erect landing posture and utilized greater hip and ankle joint range of