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Do foot orthotics have an immediate effect on the gait mechanics of inidividuals [sic] with chronic incomplete spinal cord injury?

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Do foot orthotics have an immediate effect on the gait mechanics of inidividuals [sic] with chronic incomplete spinal cord injury?
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Do foot orthotics have an immediate effect on the gait mechanics of individuals with chronic incomplete spinal cord injury?
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Jagger, Kristen Lyn
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viii, 106 leaves : ill. ; 29 cm.

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Ankle ( jstor )
Feet ( jstor )
Gait ( jstor )
Kidnapping ( jstor )
Knee joint ( jstor )
Knees ( jstor )
Lower extremity ( jstor )
Physical trauma ( jstor )
Spinal cord ( jstor )
Walking ( jstor )
Dissertations, Academic -- Exercise and Sport Sciences -- UF ( lcsh )
Exercise and Sport Sciences thesis, Ph.D ( lcsh )
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bibliography ( marcgt )
theses ( marcgt )
non-fiction ( marcgt )

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Thesis (Ph.D.)--University of Florida, 2002.
Bibliography:
Includes bibliographical references.
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Printout.
General Note:
Vita.
Statement of Responsibility:
by Kristen Lyn Jagger.

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DO FOOT ORTHOTICS HAVE AN IMMEDIATE EFFECT ON THE GAIT MECHANICS
OF INIDIVIDUALS WITH CHRONIC INCOMPLETE SPINAL CORD INJURY?



















By

KRISTEN LYN JAGGER


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


2002














This work is dedicated to my father, Robert Thornton Jagger, and my grandfather,
Ernest Thornton Jagger, for passing on the desire to succeed.














ACKNOWLEDGEMENTS

I would like to offer special thanks to each of my academic committee members for their assistance during this dissertation process. Dr.'s Chow, Tillman, Cauraugh, Vander Griend, Falsetti, and Siders provided leadership and guidance that helped me to grow as a student, a teacher, and a researcher.

I also extend my heartfelt appreciation to Mike Jones at Shepherd Center and Ben Johnson at Georgia State University in Atlanta, Georgia. Without their support throughout the recruiting and data collection processes I would not have been able to pursue this line of research.

Most of all, I would like to acknowledge my family and friends for their role in my doctoral pursuits. I thank Joy for giving unconditional daily support and being a part of my "think tank"; Chris, for helping make sense of the senseless; Mum, for always showing me the silver lining; and Dad, for asking me the questions that no one else would.














TABLE OF CONTENTS
pagqe

ACKNOW LEDGEMENTS ......................................................................................... iii

ABSTRACT .................................................................................................................. vii

CHAPTER

1 INTRODUCTION ........................................................................................... 1


2 REVIEW OF LITERATURE .......................................................................... 5

Human Spinal Cord Anatomy ................................................................. 5
Spinal Cord Injury .................................................................................. 7
Classification of Spinal Cord Injury .................................................... 8
Com plications of Spinal Cord Injury .................................................... 9
Normal Human Gait ................................................................................ 11
Phases of Gait ..................................................................................... 12
W eight acceptance ....................................................................... 12
Single lim b support ........................................................................ 13
Limb advancement .......................................................................... 14
Normal Joint Mechanics During Gait ................................................. 14
The Foot .............................................................................................. 15
Role of the Foot During Gait ............................................................... 17
The Ankle ............................................................................................20
Role of Ankle During Gait ................................................................... 20
The Knee ............................................................................................ 21
Role of the Knee During Gait ............................................................. 21
The Hip .............................................................................................. 23
Role of the Hip During Gait ................................................................. 23
Electromyography During Normal Gait ............................................... 24
Ankle musculature .......................................................................... 24
Knee musculature ......................................................................... 26
Hip m usculature ............................................................................ 27
Foot Orthotics ......................................................................................... 28
Relevant Orthotic Research ................................................................... 30









3 MATERIALS AND M ETHO DS ..................................................................... 33

Subjects .................................................................................................. 33
M ethodology ............................................................................................... 34
Data Reduction ....................................................................................... 39
Clinical G ait M easures ........................................................................ 40
Frontal Plane Joint M echanics M easures ........................................... 42
EM G M easures ................................................................................... 45
Statistical Analysis .................................................................................. 45


4 RESULTS .................................................................................................... 50

Effects of O rthotic Intervention ............................................................... 50
Clinical G ait Variables ........................................................................ 50
Gait variable comparisons within the less-involved
lower extrem ity ..................................................................... 51
Gait variable comparisons within the more-involved
lower extrem ity ..................................................................... 51
Frontal Plane Joint M echanics ............................................................. 52
Joint mechanics comparisons within the
less-involved lower extrem ity .................................................. 53
Joint mechanics comparisons within the
m ore-involved lower extrem ity ............................................... 54
Electrom yography ................................................................................... 54
Electromyographic Changes In the Less-Involved
Lower Extrem ity ............................................................................ 55
Electromyographic Changes In the More-Involved
Lower Extrem ity ............................................................................ 56


5 DISCUSSIO N .................................................................................................. 66

O rthotic Intervention ................................................................................ 66
Clinical G ait Variables ........................................................................ 66
Clinical gait comparisons within the less-involved
lower extrem ity ..................................................................... 67
Clinical gait comparisons within the more-involved
lower extrem ity ..................................................................... 68
Frontal Plane Joint M echanics ............................................................. 68
Joint mechanics comparisons within the
less-involved lower extrem ity .................................................. 69
Joint mechanics comparisons within the
m ore-involved lower extrem ity ............................................... 75
Sym m etry ............................................................................................... 76
G ait Sym m etry ................................................................................... 76
Frontal Plane Joint M echanics Sym m etry ........................................... 80
Electrom yography ................................................................................... 82
Sum m ary and Conclusions ..................................................................... 89









APPENDIX

A CO NSENT FO RM .......................................................................................... 91

B SUBJECT INTAKE FO RM ............................................................................ 98


REFERENCES .......................................................................................... .......

BIOG RAPHICAL SKETCH ......................................................................................... 105











































vi















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

DO FOOT ORTHOTICS HAVE AN IMMEDIATE EFFECT ON THE GAIT MECHANICS
OF INIDIVIDUALS WITH CHRONIC INCOMPLETE SPINAL CORD INJURY?

By

Kristen Lyn Jagger

December, 2002

Chair: John W. Chow

Major Department: Exercise and Sport Sciences

Research regarding the effectiveness of foot orthotics during functional activities is primarily limited to the able-bodied population. No research determining the impact of orthotic intervention on specific gait mechanics of ambulatory individuals after spinal cord injury (SCI) has been identified, even though this group often exhibits poor foot and ankle alignment. Purpose: This research was designed to ascertain if orthotic intervention had an immediate effect on gait measures and frontal plane kinematics and kinetics in ambulatory individuals with SCI. Methods: Nine community ambulators with an American Spinal Injury Association "D" classification were recruited. Prior to testing subjects were measured and fitted with temporary orthotics. Testing included both shoeonly and orthotic conditions, administered in random order. A six-camera optical motion analysis system captured three-dimensional motion at 120Hz, while force plate and electromyographic (EMG) data, collected at 1200Hz, provided information regarding ground reaction forces and muscle activity, respectively. An inverse dynamic approach was used to calculate frontal plane joint angles, moments, and powers at the ankle,









knee, and hip. Mean and peak EMG data were recorded from five subjects for qualitative comparison. Statistics: A 2x2 doubly multivariate repeated measures multivariate analysis of variance was used to discern the effect of orthotics on each leg. Results: Orthotics significantly increased the stance duration, step duration, and ratio of stance on the less-involved limb. Orthotics reduced the gait velocity, swing duration, maximum hip abduction angle, maximum ankle power generation, and the maximum knee power absorption on the less-involved limb. Only the maximum ankle supination angle from the more-involved limb was decreased by orthotic intervention. Qualitative results of the mean and peak EMG amplitudes suggest increases in tibialis anterior and rectus femoris muscle activity with orthotic intervention. Conclusions: There is an immediate effect of orthotic intervention on selected gait and biomechanical variables within the lessinvolved limb. There is less immediate influence on the more-involved extremity. Results of this study justify further research in this area.














CHAPTER 1
INTRODUCTION

In the United States there are more than 11,000 new cases of spinal cord injury (SCI) reported each year, with most occurring in individuals between sixteen and thirty years of age.60 One of the major consequences of SCI is paralysis or paresis of muscles below the level of injury often resulting in the inability to walk. As a result many persons with SCI must, at least in the more acute phases of their recovery, use wheelchairs for household or community mobility. Currently, between 183,000 and 230,000 American citizens with SCI function at least in part from their wheelchairs.60 Still, despite its catastrophic nature, SCI no longer guarantees that an individual will necessarily function from a wheelchair for his/her entire lifetime. Within the past few decades SCI has been classified into two major categories: complete and incomplete injuries. Incomplete injuries are defined as a partial or complete preservation of sensory and motor function below the level of the spinal cord lesion.71 Persons with incomplete spinal cord injuries comprise the majority of injuries at 50.8%. In contrast, complete injuries (no motor or sensory sparing below the level of the lesion) account for 45.8%.60

Due to advances in emergency medical intervention and spinal cord trauma

management, more individuals with SCI are surviving their initial injuries and the early stages of their care. In addition these individuals are more likely to have incomplete injuries.13,60 In part because of the growing number of persons with incomplete SCI, therapists can no longer assume that persons with SCI need only to be taught techniques and prescribed equipment which allow them to adapt to their injuries. The rehabilitation philosophy is shifting from pure adaptation to a focus on motor recovery.









Ongoing research at model care centers around the nation is being conducted to ascertain the degree of central nervous system plasticity following SCl. This research trend focuses on the recovery of walking through body-weight-supported treadmill training (BWSTT).60 For this training the patient is placed in a torso harness that provides variable body-weight support. Trained individuals assist and control the patient's lower extremity and trunk movement at different treadmill speeds. The general purpose of this training is to stimulate central pattern generators (CPG) to create a rhythmic locomotor pattern that increases the individual's capacity for ambulation. Central pattern generators are neural networks located in the central nervous system that are responsible for creating a pattern of repetitive motor actions.22 It is thought that commands for initiation and termination of these rhythmic generators originate at supraspinal levels16 and that the generators themselves lie within the spinal cord of felines.22

Evidence for the presence of CPGs in humans is more tenuous, but according to

MacKay-Lyons4�, "The 'best guess' at this point is a cautious affirmation." Central pattern generators in humans are thought to be responsible for the presence of electromyographic activity during gait that is not measurable during voluntary muscle contractions.',8" Recent studies of the electromyographic output from individuals with spinal cord injury during treadmill locomotion have shown that loading and unloading of the limbs enhances the activity of the antigravity muscles.14'23 It has yet to be seen, however, if these same electromyographic characteristics are noted during over ground ambulation.

Although, with advanced treatment techniques, there is new hope for individuals with SCI to regain the ability to walk, rehabilitation following SCI remains extremely costly. Current estimates for medical care in the first year post-injury range from $168,627 to $572,178 and from $11,817 to $102,491 for each subsequent year depending on the level and severity of injury.6� Careful management of the cost of rehabilitative services offered to persons with SCI has become imperative for facilities and insurers hoping to maintain their









viability. As a result, early SCI research into recovery of ambulation has focused on individuals with incomplete injuries. Because their injuries are "incomplete," these individuals have a theoretical advantage over persons with complete injuries when attempting to regain function, including the ability to walk. However, their levels of impairment may vary dramatically.

Incomplete injuries often result in significant sensory or motor deficits of the upper

extremities, lower extremities, and/or trunk. Even if they become ambulators, persons with incomplete SCI must often contend with significant gait dysfunction resulting from their physical impairments (i.e., increased muscle tone, spasticity, sensory ataxia, balance deficits, reduced strength). Often these impairments are addressed with bracing and/or use of assistive devices. For research purposes gait in this population may be summarily described as unable to walk, slow walking with an antalgic gait, and walking over ground at normal speeds.1

Though there is a large amount of research directed at increasing the locomotor

capacity of the spinal cord injured population, there is little available information regarding the impact of ambulation on these individuals' skeletal systems. Specifically, there is a dearth of evidence indicating appropriate interventions used to address skeletal alignment in these persons during gait training. Considering individuals are living longer with chronic SCI, the impact of orthopaedic complications resulting from abnormal joint mechanics, especially during gait, must be elucidated.

Abnormal lower extremity mechanics, particularly foot alignment, have been widely researched in able-bodied and athletic populations. 10, 29,45, 53, 55-57, 59, 67, 75, 79 Studies have focused on the use of in-shoe orthotic devices designed to alter the mechanics of the lower extremity and reduce the occurrence of secondary orthopaedic complications (i.e., pain or injury). Despite positive clinical results for orthotic intervention, this information has not been consistently applied to neurologically involved populations. Therefore, the









purpose of this research is to investigate the effect of in-shoe orthotics on the gait mechanics of individuals with chronic incomplete spinal cord injury. In order to limit the number of variables under consideration, clinical gait measures and frontal plane kinematic and kinetic parameters were chosen for study. Clinical gait measures were chosen to describe gross functional changes. Kinematic and kinetic variables were chosen to illustrate changes in frontal plane mechanical alignment due to the physical position and theoretical action of the orthotics.

In order to determine if orthotics have a significant effect on this population of individuals with incomplete spinal cord injury, the following research hypotheses are stated:

1. There will be a significant difference in selected clinical gait variables between
orthotic and shoe-only conditions during walking at a self-selected speed in this
sample of individuals with chronic incomplete spinal cord injuries.

2. There will be a significant difference in selected frontal plane kinematic variables
during walking at a self-selected speed between orthotic and shoe-only conditions
in this sample of individuals with chronic incomplete spinal cord injuries.

3. There will be a significant difference in selected frontal plane kinetic variables
during walking at a self-selected speed between orthotic and shoe-only conditions
in this sample of individuals with chronic incomplete spinal cord injuries.















CHAPTER 2
REVIEW OF LITERATURE


Human Spinal Cord Anatomy

The human central nervous system (CNS) is divided into six main parts: (1) the cerebral hemispheres; (2) the diencephalon; (3) the midbrain; (4) the pons and cerebellum; (5) the medulla; and (6) the spinal cord. The spinal cord is an elongated cylinder fewer than two centimeters in diameter.21 It is a continuation of the medulla oblongata (inferior brainstem) and lies protected within a canal formed by the alignment of successive vertebral foraminae.48 In addition to having the protection afforded by the osseous vertebrae, the spinal cord is also enveloped by the spinal meninges (i.e., pia mater, arachnoid mater, and dura mater) and a network of blood vessels lying within the surrounding fatty and loose connective tissue. In the average adult the spinal cord ranges from 42-45 cm in length and extends from the foramen magnum in the occipital bone to the level of the second lumbar vertebra. At this level we find the tapered, caudal termination of the spinal cord known as the conus medullaris.48

The spinal cord may be divided vertically into five regions: cervical (Cl -C8), thoracic (T1-T12), lumbar (L1-L5), sacral (Sl-S5), and coccygeal (one level). In cross-section, the spinal cord is divided into right and left sides, multiple laminae (gray matter), and a network of myelinated nerve axons (white matter) and connective tissue. A transverse section of the spinal cord reveals a butterfly-shaped, gray area flanked by a white region. The gray matter is composed of the nerve cell bodies that extend cranial-caudal and medial-lateral. The "wings" of the butterfly are actually the dorsal and ventral horns of the









gray matter. In the cervical and lumbar regions these horns have a greater crosssectional area. These enlargements of the gray matter regions correspond with the increased nerve supplies necessary to innervate the upper and lower extremities. The cervical enlargement extends from C4-T1, while the lumbar enlargement extends from L2-S3.48 Because the spinal cord is only two-thirds the length of the vertebral canal, position of the enlargements is not necessarily representative of vertebral levels. In the cervical region, for example, cord levels and vertebral levels are closely associated, while in the lumbosacral region cord levels L2-S3 are found at vertebral levels T1 1-L1. Termination of the spinal cord at L2 necessitates longer nerve roots for the lower lumbar, sacral, and coccygeal levels. These nerve roots exit the cord as a bundle at L2 and descend caudally before passing laterally through their respective intervertebral foraminae. Named after its appearance, this bundle of nerves is called the cauda equina (horse's tail).48

The various layers of the spinal gray matter are known as laminae, and each of

these laminae is responsible for receiving and relaying information from other portions of the CNS or the periphery to higher CNS centers or to skeletal muscle, viscera, or other organs/tissues. Cells within the dorsal horns transmit information regarding sensory input from the spinal nerve afferents, while those in the ventral horns contain alpha, beta, and gamma motoneurons that send efferent information through the ventral roots to innervate skeletal muscle fibers. Afferent and efferent axons from the various laminae of the spinal cord emerge as root fibers and converge laterally to form mixed nerves prior exiting the vertebral canal. Because the cord is divided into right and left sides, a pair of mixed spinal nerves originates from each of the above spinal levels. These nerves escape the vertebral canal laterally through the intervertebral foraminae and travel to the periphery. These nerves are named according to their associated vertebral level. Spinal nerves C1-C7 depart through intervertebral foraminae superior to their









matched cervical vertebrae. Because there are eight cervical levels within the spinal cord and only seven cervical vertebrae, spinal nerve eight exits inferior to C7. All other spinal nerves caudal to this point pass below their name-matched vertebrae.21

The white matter of the spinal cord is composed of myelinated nerve axons that run longitudinally through the cord. These fibers are responsible for the transmission of information between the spinal cord and the brain.21 Information is shared between the gray matter and white matter of the spinal cord by interneurons. This network of neurons provides the brain with a system of control and feedback loops for the entire body. Subsequently, injury to any part of the spinal cord can have a profound effect on both the motor and sensory function of the human body.

Spinal Cord Injury

Injuries to the spinal cord can be grouped into two categories: traumatic and nontraumatic. Traumatic injuries are typically, but not always, the result of high velocity impact forces. These include gunshot wounds, stab and other penetrating wounds, indirect forces generated by movement of the head and/or trunk (e.g., shearing forces from a motor vehicle accident), and direct forces at the level of the vertebrae (e.g., blunt trauma, crush injuries, falls). In addition, trauma to the spinal cord may be the result of radiation exposure or surgical procedures. Non-traumatic injuries include circulatory compromise (i.e., spinal stroke); cord compression (e.g., spinal stenosis and spondylolisthesis); demyelenating diseases (e.g., multiple sclerosis and primary lateral sclerosis); diseases of the anterior horn cells (i.e., amyotrophic lateral sclerosis and polio); inflammatory processes (e.g., spinal meningitis and acquired immunodeficiency induced neurophathy); space-occupying lesions (e.g., tumors, cysts, and syringomyelia); or congenital malformations (e.g., myelomeningoceole).









Classification of Spinal Cord Injury

There are two generally accepted methods for classifying spinal cord injury, regardless of the mechanism of injury. The first classification scheme involves a description of the spinal level at which the injury occurred, followed by a description of the most caudal level of intact motor or sensory function. For instance, a spinal "fracturedislocation of C3 on C4" describes the location of the insult, while "C5 incomplete tetraplegia" details the last level of fully preserved motor and sensory function. In the second classification scheme, defined by the American Spinal Injury Association (ASIA), it is necessary to describe the level of injury as well as the level of functional impairment. For example, included in this examination are items concerning deep anal pressure, position sense, stereognosis, pinch/grasp, and jaw chuck.

In both schemes the terms incomplete and complete are used to denote the degree of preservation of sensory or motor function below the level of injury. A complete injury indicates that both motor and sensory function below the level of the lesion have been fully compromised, including sacral reflexes in the peri-anal region. With a complete injury, however, there may be zones of preservation of sensory or motor function. These do not, however, include the sacral elements. Preservation of sacral reflexes and sensation, also known as sacral sparing, defines an incomplete injury. Incomplete spinal cord injuries, therefore, are those injuries in which there is sparing of at least partial sensory and/or motor function below the level of the spinal injury, including deep anal sensation.71 The terminology employed by the ASIA impairment scale can be found below in Table 1-1.

In order to obtain a more accurate portrayal of spinal cord injury and its influence on function, it is necessary to use both classification schemes described previously. Yet these measures alone are incapable of fully describing an individual's functional abilities and impairments after SC. Secondary complications may limit the function of one







9
individual more than another individual who had a similar mechanism and level of injury, despite being classified at the same ASIA level. For example, an individual with C5 tetraplegia with zones of sensory and motor preservation in C6.7 will be able to extend the wrist to produce a tenodesis grasp and to extend the elbows to assist with transfers and activities of daily living. A person with a C5 injury and preservation of only spotty sensation in some zones caudal to the lesion will not have the same function. Both individuals are classified at the C5 level, but the classification of A, B, C, D, or E helps to tease out the functional potential of the individual. Table 2-1. ASIA Impairment Classification Scheme ASIA A Complete loss of sensory and motor function below lesion ASIA B Partial sensory preservation; no motor sparing ASIA C Sensory and motor preservation; <1/2 muscles grade > 3/5 ASIA D Sensory and motor preservation; >1/2 muscles grade > 3/5 ASIA E No permanent sensory or motor deficits
From the National Spinal Cord Injury Statistical Center.6" Complications of Spinal Cord Injury

There are various complications that result from spinal cord injury and alter the

functional potential of its survivors. These complications can be grossly separated into orthopaedic, neurologic, and cardiorespiratory. Each of these complications leads to longer lengths of stay in the hospital or more frequent returns for follow-up care, delayed or more extensive rehabilitation, and subsequent delays in return to home and work. It is the responsibility of the rehabilitation team, therefore, to limit the occurrence and morbidity produced by these complications.

Orthopaedic problems include potential for contracture, joint ankylosis, osteoporosis or osteopenia, hypercalcemia, fractures, spinal deformity, heterotopic ossification, and degenerative joint abnormalities.71 Osteoporosis, a progressive loss of bone mass, occurs primarily in the bones below the level of the lesion and advances rapidly during the first year after injury before stabilizing. Most bone loss occurs in the trochlear bone in







10

the vertebral bodies, femoral heads/necks, the pelvis, and the bones of the leg, foot, and ankle. The cause of this rapid progression is not known, but the appearance of increased levels of calcium in the urine signals the initial onset. It is acknowledged that immobility is a partial cause of bone loss, and mobilization and weight bearing are recommended to slow its progression. Indeed, Plum and Dunning2 advocate ambulation as the only type of weight bearing that has any effect on the amount of calcium in the blood. Keeping this in mind, any intervention that increases the efficiency of ambulation may increase the amount of time that an individual spends walking each day, thereby reducing risk of additional bone loss. At the same time, responsible clinicians must consider the acuity of the person's injury, the person's age and sex, and the degree of bone loss prior to initiation of gait or weight bearing activity.

Degenerative joint abnormalities have been noted in both ambulatory and

nonambulatory spinal cord injured clients. Wylie and Chakera89 have noted that the higher the level of injury, the more likely the individual is to develop joint degeneration. It has been hypothesized that joint degeneration would be greater at the hip and sacroiliac joints of ambulatory individuals than in nonambulatory individuals.71

Sensory deficits resulting from SCI may include altered awareness of nociceptive input, altered light touch or pinprick/point discrimination, reduced baroreceptor input, reduced temperature discrimination, and altered proprioception.71 Loss of nociception and baroreception places an individual at risk for integumentary complications, such as pressure sores, secondary to a decreased awareness of physical pressure and pain. This has a direct impact on an individual's capacity to regularly perform weight shifts. Insufficient weight shifting results in areas of ischemia where regional vasculature is compressed. This eventually leads to tissue breakdown and risk of infection or need for surgical intervention. Decreased temperature sensation is dangerous for obvious reasons. A hot car seat or a prolonged wait at a cold bus stop can lead to serious problems for an









individual with SCI. Reduced light touch and point discrimination make it difficult for individuals with SCI to sense changes in their support surface. They subsequently rely more heavily on their vision for information regarding surface changes and environmental obstacles. Lack of proprioceptive input affects an individual's perception of where the body's joints are in space. Loss of deep position sense can exacerbate joint laxity if joints are positioned at their end range of motion for prolonged periods of time. Lack of position sense in the foot and ankle, in combination with reduced muscle activity, can lead to reliance on the ligamentous structures to support the joints during weight-bearing. This holds serious implications for ambulatory individuals, considering the tensile and deformation properties of connective tissue. Lack of structural support of the joints of the foot and ankle is guaranteed to facilitate joint deformation and mechanical instability.

Within the human body are several receptors designed to sense joint position (i.e., Ruffini corpuscles). These joint proprioceptors are known to profoundly impact physical function, because they are crucial for controlled voluntary movement of the body segments.71 Joint proprioceptors have also been implicated as playing a key role in central pattern generation.15 The higher CNS levels, both cortical and cerebellar, rely on feedback from joint proprioceptors, skin, muscle, and connective tissue to help refine mechanical efficiency and to accomplish purposeful movement. Therefore, it is the feedback from these various systems that helps an individual to "normalize" patterns of movement.

Normal Human Gait

Normal gait has been defined in detail by many researchers."'24- 616 - 87 The normative values generated by their research serve as a standard for the patterns of lower limb, trunk, and upper body function in humans. According to these authors, the segments of the skeletal system move in predictable arcs, the joints rotate through specified ranges of motion, and the neuromusculature has a defined phasic pattern of activation. Bianchi and







12

colleagues2 have even used the rotation of the body segments to arrive at a description of normal gait that demonstrates how kinematic parameters drive the nervous system. Phases of Gait

Although the phases of gait are often termed differently, there is consistency in the

description of events throughout the full cycle. For ease of further discussion, however, the terminology defined by Perry61 will be used.

The gait cycle can be broken up into three tasks and eight distinct phases. The tasks are termed weight acceptance, single limb support, and limb advancement. These tasks define the functional demands of the gait cycle. Within each of the tasks are the more descriptive phases of gait. These are initial contact, loading response, mid stance, terminal stance, pre-swing, initial swing, mid swing, and terminal swing. Each of these phases is named according to specific events of gait and is further described as percentages of the gait cycle (Figure 11).61 Weight acceptance

Within the task of weight acceptance there are two phases. Initial contact defines the moment when the foot touches the ground and occurs during the first 0-2% of the gait cycle. The objective of this phase is to position the limb for loading. Loading response (010% gait cycle) begins when one foot contacts the ground and continues until the opposite foot is lifted for swing. This period has also been termed the first double support or stance phase-the first time during which both feet are in contact with the ground.' 87 The objectives of this stage include shock absorption, weight-bearing stability, and preservation of progression.























* Swing


Initial Single Limb Terminal Swing Double Limb
Double Limb Stance Double Limb Stance
Stance Stance
(0%- 10%) (10%- 50%) (50%- 60%) (60%- 100%) (0%- 10%)

Figure 2-1. Perry's phases of gait.61

Single limb support

Upon termination of loading response, the body enters the period of single limb

support. This period consists of mid stance and terminal stance and makes up 40% of the gait cycle. Mid stance is the first half of single limb support and occurs between 10% and 30% of the gait cycle. It begins when the opposite foot is lifted off the floor and continues until the body's center of mass is aligned over the forefoot. The primary objectives of this phase are progression of body weight over the stance foot and lower limb and trunk stability. Advancement of single limb support results in the heel lifting off the ground and signals the beginning of terminal stance (30-50% of the gait cycle). This phase continues until the opposite foot strikes the ground. The only objective of this phase is progression of the body past the supporting foot.61









Limb advancement

Limb advancement is the final task of the limb during the gait cycle. It is composed of four phases: pre-swing, initial swing, mid swing, and terminal swing. Pre-swing typically occurs at 50-60% of the gait cycle and is also known as the second double support or stance interval.86'87 It begins when the contralateral foot hits the ground and continues until the ipsilateral foot leaves the ground. Weight transfer from one limb to the other occurs during this phase, and the main objective is to prepare and position the ipsilateral limb for the rapid demands of swing.61

Initial swing makes up the first one-third of the swing period (60-73%). It starts with the swing foot leaving the ground and ends when the foot is opposite the stance limb. The objectives of this phase are to clear the foot from the floor and advance the limb from its trailing position. The second one-third (73-87%) of the swing period is mid swing. Limb advancement and foot clearance are objectives of this phase, also. Finally, terminal swing completes the gait cycle. It begins when the tibia is vertical and ends when the foot contacts the floor. The objectives of this phase are to complete the advancement of the limb and prepare the limb for stance.61

Normal Joint Mechanics During Gait

Locomotion occurring under normal conditions incurs predictable forces on the

structures of the lower extremity. Gait mechanics that do not resemble those defined as normal represent a greater challenge to the skeletal, articular, and neuromuscular systems that bear the load and resist the applied forces. The following are descriptions of the anatomy and normal mechanics of the foot, ankle, knee, and hip joints during gait.









The Foot

In clinical and scientific literature, function of the foot is often divided theoretically into that pertaining to the forefoot, midfoot, and rearfoot (Figure 1-2). While this nomenclature is convenient for qualitative descriptions, it should be noted that the foot's composite function can not be divided so clearly. Motion at any one joint can not be measured independently during weight bearing due to the interdependence of all of the structures.







Forefoot





~'Midfoot



' Rearfoot



Figure 2-2. Functional segments of the foot. 8 The forefoot is a composition of six functional joints: the first through fifth rays and the first metatarsophalangeal joint. The first ray is composed of the first metatarsal and the first cuneiform bone. It functions in three planes, producing dorsiflexion/inversion and plantarflexion/eversion about its axis.17 Rays two through four are formed by the intersection of each metatarsal and its opposing cuneiform and appear to produce only sagittal plane motion. The fifth ray is composed of only the fifth metatarsal and is capable of triplanar motion into pronation and supination.5 The first metatarsophalangeal joint is comprised of the first metatarsal head and the base of the proximal phalanx of the hallux







16
and is capable of abduction/adduction in the coronal plane and dorsiflexion/plantarflexion in the sagittal plane.68

The midtarsal or transverse tarsal joint is located between the forefoot and rearfoot

structures and transmits forces and motion proximally and distally. This joint is made up of the talonavicular joint medially and the calcaneocuboid joint laterally. The axes at these joints allow a complex triplanar motion that consists of inversion/adduction and eversion/ abduction about the longitudinal axis (movement of the cuboid on the calcaneus) and dorsiflexion/abduction and plantarflexion/adduction of the forefoot about the oblique axis.43

Lastly, the rearfoot, comprised of the talus and calcaneus, houses the articulation called the subtalar joint (STJ). Formally called the talocalcaneal joint, this composite of three articulations (anterior, middle, and posterior) is triplanar and capable of pronation and supination.5 Pronation can be described as a combination of foot eversion, abduction, and slight dorsiflexion. Supination refers to inversion, adduction, and slight plantarflexion. Figure 2-3 shows a foot with normal forefoot and rearfoot alignment, while Figures 2-4 and 2-5 show a forefoot varus deformity and a rearfoot varus deformity, respectively. The normal range of supination and pronation during walking gait is subject to controversy. Brown and Yavorsky5 state that between 60 and 100 of combined supination and pronation are necessary for normal walking. Subotnick76 reported a necessary 4-60 of pronation and 8-120 of supination for normal gait, and Wright and colleagues note that an average of 100 of pronation is necessary in the first 8% of the stance phase. Other authors have presented calcaneal eversion requirements for normal gait.8'50 Close and colleagues8 report that a peak of 4-60 of eversion is reached by 14% gait cycle, while Moseley and associatess� report a maximum eversion of 7.30 at approximately 57% gait cycle. There remains much debate regarding the best method of describing subtalar joint motion.









Recent reports have demonstrated that measurement of calcaneal eversion/inversion is not representative of the complex three-dimensional motion of the subtalar joint.' 74
















Figure 2-3. Normal foot alignment.
Adapted from McPoil and
Brocato.'

Role of the Foot During Gait

The foot has four primary roles during the stance phase of gait:47 accommodation to uneven terrain,49 attenuation of the impact load of the ground reaction,49, 69 formation of a rigid lever for effective toe-off,49 and transferal of axial rotation of the leg to pronation/supination of the foot during stance.'49' Each of these functions is accomplished by a combination of actions at the midtarsal and subtalar joints. The specific concept of transferal of axial rotation is often called "coupling behavior" and is also used to describe the effect of the rearfoot on the proximal structures of the lower extremity32. Rodgers65 stated, "The body requires a flexible foot to accommodate the variations in the external environment, a semirigid foot that can act as a spring and lever arm for the push-off during gait, and a rigid foot to enable BW (body weight) to be carried with adequate stability."

At initial contact the heel touches the ground lateral to the ankle joint center. The ground reaction force creates a pronatory moment at the STJ that may stress the







18
structures of the medial arch.65 During pronation the talus rotates medially about the STJ axis and forces the calcaneus into a more valgus position. Talar rotation increases the freedom of motion at the transverse tarsal joint and creates a more flexible foot distal to the talus and navicular.65 A flexible foot, as mentioned above, is more able to comply with the ground beneath.

During mid stance the foot is fixed on the ground and the lower leg rotates laterally placing the talocrural joint in a tightly approximated position. Once the talus is closepacked in the mortise, the foot begins to supinate, increasing the stability of the transverse tarsal joint and the longitudinal arch of the foot.65 This stability is further increased by the firm fit of the talus on the navicular.26'41,42













00

Figure 2-4. Uncompensated Figure 2-5. Uncompensated forefoot varus deformity. rearfoot varus
Adapted from McPoil deformity. Adapted
and Brocato.46 McPoil and Brocato.46

As the limb enters terminal stance, the ankle reverts from dorsiflexion to plantarflexion at heel rise and forces the metatarsophalangeal joints to dorsiflex. Dorsiflexion of the phalanges lengthens the plantar aponeurosis due to its distal insertion on the metatarsal heads, thereby increasing tension and providing greater stability throughout the longitudinal arch. This is known as the "windlass effect" (see Figure 1-3). Thus, the







19
combination of greater weight bearing, supination, and the windlass effect contribute to the formation of the rigid lever needed for maximum stability at toe-off.65 Should any, or all, of these mechanisms fail, the foot would be unable to form the same effective lever for pushoff. Unpublished observations6 indicate that while individuals after SCI tend to present with a cavus foot in non-weight-bearing situations, they have a rigid rearfoot and an excessively mobile forefoot during weight bearing. This combination results in a foot that lacks rearfoot compliance at initial contact and the inability to form a rigid lever during terminal stance, or push-off. Figure 2-7 shows the result of an uncompensated forefoot and rearfoot varus.








Figure 2-6. The windlass effect. Adapted from Norkin and Levangie.58

Decreased rearfoot compliance results in a proximal transfer of ground reaction

forces.2873 77 Excessive motion in the forefoot decreases the ambulatory stride length, because there is little or no formation of a rigid lever.44 Placement of orthotics, therefore, is intended to reposition, or support, the bony structures of the foot in a more normal mechanical orientation. For example, an individual that presents with a rigid rearfoot inversion and a hypermobile forefoot varus would benefit from rearfoot "wedging", or posting, that brings the floor up to the medial rearfoot, and forefoot posting that prevents the forefoot from collapsing mediallly during weight bearing (see Figure 2-8). These additions would provide stability as well as increase the foot's capacity to create a rigid lever for push-off.











The Ankle

The ankle joint, or talocrural joint, is defined as the junction between the tibia, the fibula, and the talus. The junction of these three bones forms the tibiotalar, fibulotalar, and tibiofibular joints. It is also the point at which body weight is transferred from the leg to the foot. Motion at the talocrural joint has 1 0 of freedom, allowing flexion and extension in the sagittal plane.58











/ a,


Figure 2-7. Uncompensated varus Figure 2-8. Use of posting to
in weight-bearing, compensate for varus
Adapted from deformity. Adapted
McPoil and Brocato.4 from McPoil and Brocato.'

Role of the Ankle During Gait

When the foot comes into contact with the surface, a ground reaction force (GRF) results. The geometric centroid of instantaneous applied force distribution is termed the center of pressure (COP).6 During initial contact the GRF vector is directed posterior to the ankle joint center. The result is a moment, or torque, about the ankle joint called a plantar flexion moment, because it is trying to force the ankle into plantar flexion. The ankle, in response to this moment, plantar flexes until the foot is flat on the floor at an average angle of seven degrees.61 Once the foot is flat, the limb progresses forward









throughout the remainder of stance. In late mid stance the GRF moves anterior to the ankle, creating a dorsiflexion moment. To stabilize the ankle joint and continue progression of the limb, the plantar flexor muscles increase their activity, first allowing an eccentric increase in dorsiflexion range of motion to a maximum of 100 and then concentrically plantar flexing to a maximum of 300 by the end of the stance phase.8' Once the foot leaves the ground, the ankle rapidly dorsiflexes to neutral (00) in order to clear the foot as the limb swings through. There is often an increase in plantar flexion during terminal swing (3-50) that prepares the foot for initial contact.6' In reference to the example above, a hypermobile forefoot that collapses during mid and terminal stance will theoretically cause a greater dorsiflexion range of motion secondary to the delayed push-off that results from the lack of a rigid foot formation. Placement of orthotics, therefore, should decrease the dorsiflexion range of motion by promoting a rigid forefoot for earlier push-off.

The Knee

The knee joint is the junction of the femur proximally and the tibia distally. Due to the general lack of congruency of the tibiofemoral articulation, motion in all three planes is possible. The primary arcs of motion are sagittal plane flexion and extension, while the lesser rotations are transverse plane internal and external rotation and frontal plane abduction and adduction. The knee provides stability during the stance phase of gait and mobility during the swing phase"8.

Role of the Knee During Gait

Sagittal plane knee motion during the gait cycle ranges from 0-700 of flexion. At initial contact the knee is close to full extension (20 hyperextension to 50 flexion)27 and is experiencing an extensor moment secondary to the anterior position of the GRF vector. During loading response, there is a rapid rate of flexion as the body progresses anteriorly









and the limb is loaded. Approximately 180 of flexion is necessary to provide shock absorption and is achieved at rates as high as 3000/sec. This is the time at which maximal weight bearing occurs at the knee joint. After this loading, the knee extends once more to a maximum of 30 flexion by terminal stance, after which it rapidly flexes to a maximum of 600-700 during the swing phase.5'61 The peak flexion angle is achieved by mid swing, at which point there is a brief pause followed by a reverse in direction toward full extension by late terminal swing and initial contact.

Transverse plane rotation has been described using electrogoniometers and 3-D videography with bone pins and surface markers.9',33 36 At initial contact the femur is slightly externally rotated on a fully rotated tibia.6' During loading response the tibia and, at a slightly slower rate, the femur internally rotate approximately 70 to their maximum angle by the beginning of mid stance.38 Electrogoniometric data presented by Kettelkamp and colleagues33 indicate that tibial internal rotation persists through mid stance until the knee fully extends in terminal stance. By the time the knee is fully extended, the tibia and femur are, once again, fully externally rotated. This rotation remains until body weight is shifted to the contralateral limb and the ipsilateral limb prepares for swing. At this point the tibia and femur again internally rotate. In opposition to each of these reports, Lafortune et a136, using bone pins to quantify tibiofemoral rotation, report that internal rotation measuring less than 50 occurs at initial contact, while 00 rotation is the mean during the remainder of stance. A maximum external rotation of 9.40 was recorded at toe-off until 75 ms prior to initial contact.m

Coronal plane motion of abduction and adduction has been described by Kettelkamp et a133 and Lafortune and colleagues.3 There are, again, conflicting results for these values. Both groups of researchers agree that the knee abducts during the stance phase (mean peak value of 6.40 according to Lafortune et al.3 During the swing phase, however,







23
Kettelkamp and colleagues33 reported a maximum value of 80 adduction, while Lafortune et a13 reported abduction and not adduction in four out of five subjects. The Hip

The hip joint is composed of the head of the femur distally and the acetabulum of the pelvis proximally. These two structures form a ball-and-socket joint with 30 of freedom. Flexion and extension occur in the sagittal plane, abduction and adduction in the coronal plane, and internal and external rotation in the transverse plane. The primary role of the hip joint is to support the head, arms, and trunk in both static and dynamic situations.5 Role of the Hip During Gait

The hip joint serves three major functions during gait.61 First, it provides the junction between the lower extremities (the locomotor units) and the trunk (the passenger unit). Secondly, it provides stability during stance; and thirdly, it controls the limbs during advancement in swing phase. The hip joint is designed to provide more three-dimensional motion than the knee or ankle and its muscles are divided into stabilizers and prime movers.

Sagittal plane motion consists of flexion and extension of the thigh segment relative to the vertical frontal plane or the pelvic segment. The rotational angle can be defined as the thigh angle (between the thigh segment and the vertical) or as the hip angle (between the thigh and the pelvis segment). Hip angle measures have been used in clinical literature, because they account for the degree of pelvic tilt present during gait. The normal range of motion of the hip is approximately 400 during gait.3'31, 52 At initial contact the hip is flexed 300 and remains relatively steady during loading response. At 38% gait cycle a neutral hip angle (00) is achieved, and as mid stance is approached the hip continues its progression toward extension. A peak hip extension of 100 is reached when the contralateral foot contacts the ground at 50% gait cycle, signaling the reversal of hip rotation into flexion by







24
initial swing. At 60% gait cycle the hip passes through neutral once more and continues to its peak near 350 during terminal swing.61

Small rotations of coronal plane adduction and abduction occur at the hip during gait. At initial contact the hip is adducted approximately 100 and increases another five degrees during loading response.70 During mid stance there is a reversal in rotation and the hip reaches a neutral angle (00) by the end of terminal stance. This rotation toward abduction continues into initial swing, reaching 50 of abduction before altering direction toward adduction once more.

Measurements of transverse plane rotation at the hip have been highly variable when comparing results across various studies.3 The total arc of motion averages 80, with peak internal rotation occurring during loading response and peak external rotation at the end of pre-swing.3

Electromyography During Normal Gait

Just as there is a definable pattern of motion during human locomotion, there is also a structured and repetitive sequence of muscle activation. The timing of muscle activation is normally described in terms of the phase of gait during which it is active. Magnitude of muscle activation, on the other hand, is defined as a percentage of the maximal EMG produced by a maximal voluntary isometric contraction (MVIC) during a manual muscle test (MMT) muscle activity is presented according to the joint upon which it has the greatest effect. Only those muscles tested for this study will be described in detail here. Ankle musculature

The muscles crossing the ankle joint are grossly divided into those that cross the

anterior aspect of the ankle joint, termed dorsiflexors, and those that cross the posterior aspect of the joint, termed plantarflexors. The dorsiflexor muscles predominate during the swing period of gait secondary to their role in lifting the foot for ground clearance.









The plantarflexor muscles are more active during stance as they are responsible for controlling progression of the body.61

The primary movers during dorsiflexion are the tibialis anterior (TA), the extensor

digitorum longus (EDL), and the extensor hallucis longus (EHL). The TA has the largest cross-sectional area and mass, has a similar lever arm length when compared to the EDL and EHL, and therefore produces the greatest amount of torque. The onset of dorsiflexor activity begins during pre-swing with the EHL firing first and is quickly followed by the TA and EDL. Tibialis anterior activity peaks initially during swing at approximately 35% MVIC, drops to 10% during mid swing, and rises sharply to 45% by early loading response.1 The double peaks of the TA and EHL designate the pretibial muscles as biphasic.

The plantarflexor muscle group is composed of seven primary movers that are subdivided into the triceps surae and the perimalleolar muscles. The triceps surae is composed of the soleus and gastrocnemius. The perimalleolar muscles are made up of the tibialis posterior (TP), flexor digitorum longus (FDL), flexor hallucis longus (FHL), peroneus brevis (PB), and peroneus longus (PL). The primary role of the triceps surae is production of plantar flexor torque, accounting for 93% of the theoretical total. The perimalleolar muscles are concerned with control of the subtalar joint and other foot articulations during gait and account for only 7% of the theoretical torque.61

The soleus and gastrocnemius muscle activity begins during the loading response phase of gait. The medial head of the gastrocnemius and the soleus activate in near unison, while the lateral head of the gastrocnemius is delayed until mid stance.78 The increase in gastrocnemius activity follows a more linear path until late mid stance when it rises rapidly, culminating at a peak of 60% MVIC values in terminal stance. It returns to zero by the initiation of pre-swing.









Knee musculature

The muscles that cross the knee are grossly divided into those that extend the knee joint and those that flex the knee joint. Knee extensors cross the joint anteriorly while knee flexors generally cross the joint posteriorly. Muscle activity is present to assist with the three primary functions of the knee: shock absorption, extensor stability, and limb advancement.61

The knee extensors are comprised mainly of the quadriceps muscle. The vastus medialis, vastus lateralis, vastus intermedius, and rectus femoris form the quadriceps. The vasti muscles cross only the knee joint, while the rectus femoris crosses the knee and hip joints. The vasti muscles act in synchrony, extending the knee during terminal swing (90% gait cycle) to position the limb for initial contact and rapidly increasing activity that peaks at 25% MVIC during early loading response (5% gait cycle). At the onset of mid stance these muscles drastically reduce their level of activity and are quiet by 15% of the gait cycle. The rectus femoris has a very different activation pattern due to its dual role of extending the knee and flexing the hip. It is active between late pre-swing (56% gait cycle) and early initial swing (64% gait cycle) at approximately 20% MVIC. This muscle is seldom active with the other vasti muscles during loading response.61

Knee flexor muscles can be divided into those that cross one or two joints, as was

seen in the knee extensor musculature. The popliteus and the biceps femoris short head (BF-SH) comprise the knee flexors that cross only the knee joint, while there are three knee flexors that cross both the hip and knee joints.6 These muscles, in addition to BFSH, make up the hamstrings. They are the biceps femoris long head (BF-LH), semimembranosus (SM), and semitendinosus(ST). Although each of these is a primary hip extensor, they have dual roles as knee flexors. They share an onset time during late mid swing (75% gait cycle) and continue through terminal swing. The SM and ST often continue their activity throughout mid stance. Biceps femoris long head and ST reach









peak activities near 20% MVIC during late mid swing and terminal swing, while SM peaks near 30% MVIC during early terminal swing.61

There are other muscles that add to the knee flexion effort even though their primary responsibilities lie at the hip and ankle. The gastrocnemius, mentioned previously, is a primary ankle plantarflexor that assists knee flexion from 15% to 50% of the gait cycle and peaks at 75% MVIC during mid terminal stance.6' Hip musculature

The muscles of the hip joint can be divided into extensors, abductors, flexors, and

adductors. The roles of these muscles include stabilization of the trunk on the lower limb and limb advancement. The hip extensors and abductors are active during stance to stabilize the limb, the flexors are active during swing to advance the limb, and the adductors are active during the intervals between stance and swing to assist transition from one role to another. Due to limitations in electromyographic techniques there is no reliable information present regarding the activity of the deep external rotators.6'

The hip extensor muscles include the two-joint hamstring muscles, the adductor magnus, and the lower fibers of the gluteus maximus. The semimembranosus, semitendinosus, and biceps femoris long head act to eccentrically slow the forward progression of the limb during late mid swing (80% gait cycle) and throughout terminal swing. They continue their activity until mid loading response (8% gait cycle). The adductor magnus and lower fibers of the gluteus maximus have a slightly later onset, initiating in mid terminal swing and continuing through most of loading response.8'

The biceps femoris long head and semimembranosus peak during early terminal swing at 20% and 30% MVIC, respectively, and taper off to 10% by initial contact. The adductor magnus and lower fibers of the gluteus maximus do not peak until initial contact, at which time they have been recorded at 40% and 25%, respectively.6'







28
The hip flexor muscles are composed of the adductor longus, rectus femoris, gracilis, sartorius, and iliacus. According to Perry61, activity of the flexor muscles is most significant during the first stride of gait at a normal speed. After this initial acceleration the muscle activity subsides greatly (<5% MVIC). During faster walking speeds, however, greater muscle activity is seen.

Flexor muscle activity begins during pre-swing and continues through initial swing and into early mid swing. There appears to be a definable sequence of muscle activity during this time. The adductor longus is the first muscle to become active (during terminal stance) and also remains active the longest (into early mid swing). The adductor brevis may have a similar pattern of activity, but unreliable EMG techniques preclude its accurate collection. Rectus femoris, as stated earlier, becomes active during pre-swing and remains active into early initial swing. At self-selected speeds for walking, the rectus femoris showed little to no activation in half of the subjects tested.61 Hip flexor activity is not typically witnessed during mid swing. This is largely due to the release of potential energy that is stored in the muscles when the limb moves into hip extension.1 Foot Orthotics

Foot orthotics are orthopaedic devices that are used to provide support under the metatarsal shafts and calcaneus in order to position the foot in a neutral position and prevent compensatory motion.4 These inserts can be described in lay terms as "shims" that transpose the floor to the plantar surface of the foot while maintaining subtalar neutral (see Figure 2-4).46 Functional orthotics should not be confused with arch supports that create a weight-bearing surface within the longitudinal arch. Orthotics are intended to support the load-bearing structures of the foot while leaving the longitudinal arch to function in its normal mechanical role of suspension. A suspension bridge is designed with vertical supports at each end of an expanse and cables to support the central section. The foot, on the other hand, has a talus, a calcaneus, and metatarsals that are present









proximally and distally for support, with longitudinal ligaments and fascia that act as the cables. Central support is not necessary and can be detrimental to the structural integrity of suspension-based structures.

Orthotics are Type I medical devices that can be classified according to the density of materials and processes used during fabrication. Soft orthotics are constructed of lowdensity materials that deform easily and are intended for accommodative support46 and shock absorption.39 These can be bought off-the-shelf or can be custom-made but are not used in conjunction with posting (wedging) to correct biomechanical alignment.46 Semirigid orthotics incorporate softness for shock absorption and rigidity for biomechanical control.46 Unlike soft orthotics, they use extrinsic posting to control forefoot and/or rearfoot motion. Extrinsic posting refers to the addition of materials to a negative cast of the foot. This method of posting requires casting of the foot while it is maintained in neutral position, followed by formation of a "negative cast", or model, of the foot. Once the neutral negative cast is made, materials are added to the plantar surface until the desired amount of posting/control is achieved. Rigid orthotics are the least forgiving of all inserts. They are typically made of a single layer of heat-moldable plastic in order to minimize material and provide the highest degree of biomechanical control. Intrinsic posting is used in these orthotics. This method requires careful addition of plaster to the neutral foot cast in order to post the foot without further addition of materials. Inaccuracies in this technique will result in an orthotic with an unacceptable fit for the client.

Temporary orthotic posting falls into the semi-rigid category. The materials are of moderate density and are intended to biomechanically support the foot, without the preparation of a neutral cast. Soft inserts are cut to the size of the individual's insoles and posting is applied to the appropriate areas. Temporary orthotics are often manufactured to determine the effectiveness of this intervention and to modify the degree of posting necessary before a custom pair is fabricated for the client.









Relevant Orthotic Research

As stated above, the fundamental goal of orthotic intervention is to support the foot

during the stance phase of gait such that the need for the foot to compensate for structural malalignment is reduced.39 Many researchers have attempted to quantify rearfoot motion. A wide variety of analyses have led a large number of authors to report changes in kinematic variables as a direct result of orthotic intervention.10, 29, 45, 53, 55-57, 59, 67, 75, 79 Some researchers specifically noted reductions in the maximum pronation of the foot, maximum pronation velocity, time to maximum pronation, total rearfoot movement,45.59, 67 rearfoot pronation,45 57, 67, 79 and calcaneal eversion.57,59,67,79 Others noted alterations in the amount of intemal tibial rotation as a result of orthotic use, 10,53, 56 75 while still others focused on the electromyographic changes that resulted from the use of orthotics.'' 81 However, because other authors found no significant difference in many of the same variables,4 . 7 results remain inconclusive63 and further analyses of the methodologies must be performed.

One of the many discrepancies noted among experiments is the type of kinematic analysis used. It is well accepted that proper human biomechanical analyses must be performed using three-dimensional techniques. Human segmental rotations and translations are not purely angular or linear, but a combination of the two. Therefore, the use of two-dimensional analyses to describe three-dimensional motions are suspect. Some authors have adhered to three-dimensional analyses, and others have not.10 29,67

In addition to questionable videographic methodology, there are concerns regarding the placement of rearfoot markers when attempting to measure calcaneal motion. Superficial markers are generally used in preference to intracortical bone markers, because they are noninvasive, cheaper, easy to apply, and do not typically hinder performance of the activity of interest. They are, however, more susceptible to







31
measurement error. Calcaneal motion would best be measured noninvasively by marker placement directly over the heel, but this is not possible with activities that require shoes. Attempts have been made to represent calcaneal motion more accurately within the shoe. Nigg and associates cut windows in the heel counters of running shoes and placed markers directly over the calcaneus, while Genova and colleagues" affixed a piece of molded plastic to the heel and bent it in a "U" fashion over the heel counter of the shoe. Both methods are good attempts at measuring calcaneal motion more accurately, but window cutouts in shoes raise the question of lost stability for the rearfoot within the shoe, and molded plastic inserts may move on the heel and misrepresent calcaneal motion. Sandals have also been used in place of running shoes when quantifying rearfoot motion.53 While providing adequate access to the heel for rearfoot markers, they lack the stability of a full lace-up running shoe. Altering shoe stability in any way may cause an individual to change the mechanics of their stride.

The most accurate method of measuring true skeletal calcaneal motion to date is through the use of intracortical bone pins. Stacoff and colleagues75 used bone pins to demonstrate the differences between skeletal motion and shoe movement during running. Their findings demonstrated an average 5.80-7.30 difference between calcaneal and shoe eversion. A critique of the methodology must also be performed here. A small cutout had to be made in the heel counter of the shoe to allow attachment of the marker set to the bone. In addition, all testing was performed with a local anesthetic. Both of these procedures, though necessary, may have altered normal shoe motion or rearfoot biomechanics.

Perhaps realizing the difficulties associated with accurate quantification of rearfoot

motion, and in attempting to learn more about the impact of the ankle/foot complex on the structures higher in the kinematic chain, some authors have attempted to determine alterations in tibial rotations.7' 10,53 Here again, the results remain inconclusive. Cornwall







32
and McPoil1� demonstrated a high correlation between rearfoot motion and tibial rotation (r = .953), and noted a decrease in internal tibial rotation velocity and acceleration with orthotic use compared with a barefoot condition. Nawoczenski and colleagues53 noted a decrease in tibial rotation during the first 50% of stance when semi-rigid orthotics were used, but stated that no there was no significant difference in mean tibial rotation throughout stance between orthotic and non-orthotic conditions.

Given the drawbacks and difficulties listed above, it does not seem prudent to attempt measurement of rearfoot motion in a population with abnormal gait characteristics when that of able-bodied "normals" is still under scrutiny. More appropriate global measures consist of general gait characteristics, joint angular measures, joint moments, joint powers, and average EMG amplitudes. The purpose of this research is to determine if orthotic intervention has an impact on mechanical components of gait affecting overall function. Because the functional variability between subjects would limit the usefulness of the data, a within-subject comparison of orthotic and shoe-only conditions will be employed.















CHAPTER 3
MATERIALS AND METHODS

Because of larger subject pools in major cities, the data collection for this project was conducted in the Georgia State University Biomechanics and Ergonomics Laboratory in Atlanta, Georgia.

Subjects

Eight males and one female were recruited for this research project (Table 3-1). Prospective subjects were located via a Shepherd Center (Atlanta, GA) database search, and a mailing describing this research was sent to those who resided within a 250-mile radius of Atlanta, Georgia. Individuals were included in the mailing if they were between the ages of 18 and 65 years and classified by the American Spinal Injury Association (ASIA) as Impairment Scale category D80 as of discharge from Shepherd Center. Interested individuals were asked to return a reply card that included their name, address, and phone number. A phone interview was conducted with each respondent prior to scheduling. Inclusion criteria required that each subject was: (a) medically diagnosed with a first-time spinal cord injury from trauma, or vascular or orthopaedic pathology at cervical, thoracic, or lumbar levels; (b) at least six months post-SCI prior to participation in this research and had completed all spinal cord rehabilitation; (c) was categorized by ASIA as Impairment Scale category D;60 (d) a community ambulator (50 m) with or without an assistive device; (e) a household ambulator (15 m) without an assistive device; and (f) able to give informed consent. Subjects were excluded if any of the above criteria were not met, or they were symptomatic for pain or other significant medical complications that would prohibit or interfere with testing of walking function.









Table 3-1. Subject data with level of injury and degrees of orthotic posting. Subject Sex Age Height Weight SCI More-Involved RF Posting FF Posting
(years) (cm) (kg) (level) (lower extremity) (Left, Right) (Left, Right)
BS Male 62 185.4 97.5 C6 Left 60,60 60,60 CO Female 54 165.1 63.1 T9-10 Right 60, 40 60, 00 CP Male 18 172.7 65.8 C1-2 Right 6, 20 40, 00 DS Male 50 190.5 104.5 T10-11 Right 60,60 60,60 EB Male 50 175.3 68.2 C4-5 Right 60, 60 40 60 JD Male 45 188.0 78.0 C5-6 Left 50 30 60, 60 MS Male 22 172.7 59.0 C3-5 Right 60, 60 60, 20 MT Male 63 185.4 90.9 C6-7 Right 30 50 60, 60 RM Male 34 177.8 78.9 T12 Left 60, 40 40, 30 Note: SCI=Spinal Cord Injury; RF = rearfoot; FF = forefoot; All posting material was placed medially.

Methodology

Prior to the arrival of each subject, all experimental equipment was inspected and calibrated by the investigator. The AMTI (model BP400600) force plate (Advanced Medical Technology, Inc., Watertown, MA) was allowed at least 15 minutes to warm up per manufacturer guidelines. Six 120-Hz optical digital cameras were placed as shown in Figure 3-1 and a Peak Performance@ calibration frame (Peak Performance Technologies, Inc.) was used to define a 2m x 2m x 2m (length x width x height) three-dimensional space in which all coordinate data would be gathered. Successful calibration was determined when the true length of a 0.914 m calibration rod was accurately estimated by the Peak Motus� 2000 video system and had a standard deviation from its true length of less than

0.003 m.

When a subject arrived at the Georgia State University Biomechanics and Ergonomics Laboratory, he/she was provided with a consent form that was approved by the Georgia State University (GSU) Institutional Review Board (Appendix A). After reading and signing the consent form, the subject was assessed during a single four-hour period. The subject first changed into appropriate clothing (dark colored shorts and shirts) and was then weighed and measured. Measurements of anthropometric data (Figure 3-2) from the










Camera 1 Camera 2 Camera 3












> Force olate







Camera 6 Camera 5 Camera 4 Figure 3-1. Experimental setup.

lower extremities and pelvis were recorded for use in determining segment inertial properties. A manual muscle test and sensory evaluation were performed as defined by ASIA to confirm category of impairment. More- and less-involved lower extremities were defined by summing the muscle grades for each leg. The leg with the lower score was labeled as the more-involved lower extremity. Further bilateral measures were taken to define the available rearfoot and forefoot motion. All measurements were recorded on a subject evaluation form (Appendix B).

The method described by McPoil and Brocato46 was used to measure available

forefoot and rearfoot motion. Briefly, this method required the subject to lie prone while two lines were drawn on the back of each lower extremity, one bisecting the calf and the other bisecting the calcaneus. These lines were used to determine the position of subtalar neutral and to measure the available calcaneal inversion and eversion and forefoot varus







36
or valgus using a hand-held goniometer. The only variation from the procedure described by McPoil and Brocato was the use of overpressure when measuring forefoot varus. Overpressure was used to obtain a more realistic measure of forefoot motion in a weightbearing position.

ASIS Breadth



Thigh Length
Midthigh Circumference


Knee Diameter
Calf Length (
/ Calf
Circumference
Malleolus Width

Foot Breadth HeightL j

Foot Length
Figure 3-2. Anthropometric measurements. Adapted from Vaughn.2

All measures were used to determine the amount of forefoot and/or rearfoot posting necessary for optimal positioning for each foot. Temporary orthotics were then fabricated from semi-rigid posting material (EVA 70 shore A; Orthofeet, Inc., North Vale, NJ) and affixed to a semi-rigid felt backing that was cut to the size and shape of the subject's own shoe insoles. A hand-held electric rotary grinder (Black and Decker RTXTM, The Black and Decker Corporation, Hampstead, MD) was used to smooth the edges of the posting material to reduce plantar surface irritation. Rearfoot posting material was placed under the calcaneus and forefoot posting was placed immediately posterior to the metatarsal heads in the forefoot. All posting materials were measured and placed according to the methods described by McPoil and Brocato,4 with a maximum of 60 allowed at the forefoot







37
or rearfoot. The 60 limit was deemed necessary to protect individuals from injury to the fifth metatarsal caused by shifting the center of pressure laterally. Medial forefoot posting was used to support an excessive forefoot varus and medial rearfoot posting was used either to stabilize a greater than normal calcaneal inversion (rearfoot varus) or to limit an extreme calcaneal eversion (rearfoot valgus) during stance. The goal of the orthotics was to place the subtalar joint as close to neutral as possible during functional weight bearing tasks.

Once the physical measures were recorded and orthotics were manufactured,

electromyography (EMG) electrodes were placed at predetermined sites (Figures 3-3, 3-4, and 3-5) to record the average and maximum amplitudes of selected muscles during walking. Electrode placement was determined using descriptions and diagrams provided by Cram.12 Each electrode site was shaved and vigorously cleansed with alcohol prior to placement. Blue Sensor (Medicotest Inc., Rolling Meadows, Illinois) 2-cm diameter surface electrodes and a telemetered Noraxon Telemyo 900 (Noraxon USA, Inc., Scottsdale, AZ) EMG system were used to detect muscle activity, while the Peak Motus 2000@ (Peak Performance Technologies, Inc.) displayed and recorded the raw signals.

The order of testing was randomly assigned. Initial testing condition (orthotic or shoeonly) and leg (right or left) were logged on the subject's evaluation form. Reflective markers (1cm and 2cm diameter) were placed according to the Modified Helen Hayes marker system82 (Figure 3-6) over pre-defined bony landmarks in preparation for over ground walking trials. A minimum of 15 minutes was allowed for accommodation prior to testing.

Once five successful practice passes were made, the recorded trials began. On

average, 10-15 practice passes had to be made before the five successful passes were accomplished. Five successful self-selected speed test trials were then recorded. Next, the subject was repositioned for practice passes for the contralateral foot. Five successful



























Figure 3-3. Electrode placement for
tibialis anterior. Adapted
from Cram.12


Figure 3-5. Electrode placement for
medial hamstrings and medial gastrocnemius.
Adapted from Cram.12


Figure 3-4. Electrode placement for
rectus femoris. Adapted
from Cram.12









practice passes were required for this leg also prior to collecting the test trials. After successful practice passes, five successful test trials were recorded. An unsuccessful pass or trial occurred when one foot was not fully and solely on the plate during a single pass, or the subject obviously altered his/her stride length (targeted) to step on the plate. Unsuccessful trials that occurred during the recording sessions were discarded. A 15minute rest and accommodation period was provided after the first condition. Additional rest periods were provided at any time the subject requested. Upon completion of all testing, electrodes and reflective markers were removed and subjects were paid for mileage to and from the laboratory.

Data Reduction

Raw analog data from the force plate and telemetric EMG were saved directly onto the hard drive of the Peak Motus� 2000 workstation computer. Digital optical data from each camera were recorded and used to describe the three-dimensional path of motion of each reflective marker during each walking trial. Each path was defined manually as that pertaining to a specific reflective marker and then automatically tracked by the Peak Motus� 2000 system throughout the remainder of the trial. Occasional aberrant light would interrupt the automatic processing and trigger a query regarding the identity of one or more markers. The marker(s) of interest was (were) redefined manually at that time and automatic tracking was resumed. This process was followed for each subject's 20 walking trials.

After pathways were identified in each trial, lower extremity coordinate, raw analog, and event data were extracted from the Peak Motus� 2000 system and loaded into a LabVIEW 6.0� (National Instruments Corporation, Austin, TX) kinematic and kinetic data calculation program. This program used the inverse dynamic approach described by









Robertson and Winter" to calculate three-dimensional kinematic and kinetic data, in addition to providing clinical gait measures.

JR. ASIS 14 L.ASIS



R. Femoral wand R. Femoral /
epicondy.o / R. Ttbial~a wfd\1 S 4Y
R. iee Malleolus /
", I- L. Metatarsal head II


Global reference frame


Y
Figure 3-6. Modified Helen Hayes marker system." Clinical Gait Measures

Clinical gait variables were analyzed because they are measures commonly used in rehabilitation to quantify the effect of specific interventions. Gait values are typically easy and cost-effective to measure and provide valuable information regarding functional capacity. The clinical gait variables of interest are defined below. For each subject, the









average value over five trials for each measure and each condition was used in

subsequent analysis.

* Stride length (m) - the distance either foot travels during one stride (foot
contact to ipsilateral foot contact) (Figure 3-7)
* Gait velocity (m/s) - walking speed (measured as the stride length divided by
the time taken to complete that stride)
" Step length (m) - the distance one foot moves ahead of the other foot during
the gait cycle (Figure 3-7)
" Step width (m) - the distance between the heels of the feet during gait
(Figure 3-7)
* Toe clearance (m) - the distance from the toe of the foot to the floor
" Toe-out angle (0) - the angle of the foot segment from the line of progression;
also known as foot angle (Figure 3-7)
" Step duration (s) - the time for the test foot to take one step length
" Stance duration (s) - the time that the test foot is in contact with the ground
" Swing duration (s) - the time that the test foot is not in contact with the
ground
* Ratio of stance (%) - the percentage of the gait cycle during which the test
foot is in contact with the ground
* Ratio of swing (%) - the percentage of the gait cycle during which the test
foot is not in contact with the ground
" Double limb support (%) - the percentage of the gait cycle during which both
feet are in contact with the floor (Figure 3-8)
* Single limb support (%) - the percentage of the gait cycle during which the
test foot only is in contact with the floor (Figure 3-8)

[ Left step length - Right step length

Awq% Left foo~t angle




Stride length I Figure 3-7. Selected gait variables.























* swingn


Initial Single Limb Terminal Swing Double Limb
Double Limb Stance Double Limb Stance
Stance Stance
(0%- 10%) (10%- 50%) (50% 60%) (60%- 100%) (0%- 10%) Figure 3-8. Perry's phases of gait.61

Frontal Plane Joint Mechanics Measures

Frontal plane joint mechanics were chosen for analysis because the orthotic posting material acts primarily by altering frontal plane position of the bones of the foot. Frontal plane kinematic and kinetic measures were computed using reference frames embedded at each segment's center of mass (Figure 3-9). Segment reference frames are used to define the position and orientation of each segment with respect to a global reference frame.82 Changes in position and/or orientation over time, in conjunction with the application of external forces, allow computation of joint angles, joint moments, and joint powers. The following briefly describes the placement of segment reference axes and the computation of frontal plane angles, moments, and powers.

Each segment reference frame is composed of an x-, y-, and z-axis. The origin of

each reference frame is at the center of mass of each segment, and the global reference frame lies outside the body within the laboratory setting (Figure 3-9). The x-axis of each







43
segment reference frame is generally directed from the distal joint center to the proximal joint center. In each thigh, for example, the x-axis is directed toward the hip on a line between the knee and hip joint centers. Each leg's x-axis is directed toward the knee on a line between the knee and ankle joint centers, and each foot's x-axis is directed from the foot's center of mass toward the heel marker on a line between the toe and heel markers. The pelvis must also have an embedded reference frame for hip computations to be performed. The pelvis's x-axis is simply directed from the pelvic center of mass vertically and is parallel to the global Z-axis.

Once the x axis is defined, external markers (Figure 3-6) are used to define a plane that bisects the segment. In the hip, the xz-plane is defined by the x-axis and the thigh wand marker; and in the leg, the xz-plane is formed by the x-axis and the calf wand marker. The xy-plane is described in the foot by the ankle joint center, the heel marker, and the toe marker. Lastly, the xz-plane in the pelvis is formed by the x-axis and an axis parallel to the line between bilateral anterior superior iliac spines. Using the right hand rule, the remaining axis, which is perpendicular to the previously defined plane, is determined.

Joint angles in the frontal plane are defined as the rotation of the distal segment

relative to the proximal segment about the anterior-posterior (AP) axis. Positive angles represent hip and knee adduction and foot pronation. Negative values represent hip and knee abduction and foot supination. For each subject, the average value over five trials for each measure and each condition was used in subsequent analysis.

* Frontal plane hip angle (0) - the rotation of the thigh reference frame relative
to the pelvic reference frame about the AP axis (y axis)
* Frontal plane knee angle (0) - the rotation of the leg reference frame relative
to the thigh reference frame about the AP axis (y axis)
" Frontal plane ankle angle (0) - the rotation of the foot reference frame relative
to the leg reference frame about the AP axis (x axis)










XPeMs


z Y






X ~ 4 --_,
x
0(





Figure 3-9. Segmental reference frames.82

Moments are the cross product of forces applied at known distances from a point of rotation and are a measure of how much force is required to produce a rotation about the axis. Moments, like linear forces, are three dimensional in nature, and can be broken down into components acting about each of the cardinal axes. Lower extremity frontal plane moments are rotary forces that act about a floating axis that is perpendicular to the mediolateral and longitudinal axes.82 Free body diagrams are often employed to display forces acting on a body at specific instants in time. A free body diagram of the foot (Figure 3-10) depicts the resultant force measured by the force plate and the subsequent component reference frame forces acting on the foot segment's center of mass during pre-swing. These component forces must be balanced by equal and opposite forces and resistant moments at the ankle joint center. The force acting in the reference z plane creates a frontal plane moment of









resistance about the ankle joint center. This moment, along with frontal plane moments at the knee and hip, were used in the analysis.

Joint power is the product of the resultant joint moment and the joint angular

velocity. The sign of the product determines whether power is being generated (+) or absorbed (-). Maximum values of frontal plane power generation and absorption at the hip, knee, and ankle were analyzed here. For each subject, the average value over five trials for each measure and each condition was used in subsequent analysis.

" Frontal plane hip power (W) - the product of the frontal plane hip joint moment
and the frontal plane angular velocity of the hip joint
" Frontal plane knee power (W) - the product of the frontal plane knee joint
moment and the frontal plane angular velocity of the knee joint
* Frontal plane ankle power (W) - the product of the frontal plane ankle joint
moment and the frontal plane angular velocity of the ankle joint EMG Measures

Raw EMG data were extracted from the Peak Motus� 2000 system and processed using Microsoft� Excel (The Microsoft Corporation, Redmond, WA). Data were smoothed using a sliding averaging with a window of 13 data points. Once all channels were smoothed, the mean and maximum values during one stride were calculated for each trial and then averaged by side and condition. Because of technical problems encountered during data collection, complete EMG data were only available in five subjects. Statistical analyses, other than descriptive measures, were not performed on EMG values. Trends noted in mean and maximum EMG amplitudes are discussed.

Statistical Analysis

A single 2x2 (Side x Intervention) doubly multivariate repeated measures multivariate analysis of variance (DM MANOVA) was used to determine significant differences between side conditions and intervention conditions (Figure 3-6). In the presence of a main effect, univariate tests were used to identify the level of significance of each









\MR.
\


\


FR _. v" WF


IRz


,OT


FR~


Figure 3-10. Free body diagram of the foot during pre-swing. The forces acting at the
foot's center of mass create moments of resistance at the ankle joint center about their respective axes of rotation. FR=resultant force; XA=force in the x plane; YA=force in the y plane; ZA=force in the z plane; WF=force due to the
weight of the foot; FAx=force at the ankle in the x plane; FAy=force at the ankle
in the y plane; FA,=force at the ankle in the y plane; Mp.=moment of
resistance about the x-axis; MRy=moment of resistance about the y-axis;
Mz=moment of resistance about the z-axis.

dependent variable within each condition (Figures 3-7 and 3-8). Pairwise comparisons using Bonfferoni adjustments were used when appropriate as post hoc analyses to determine the effect of one condition level on another (Figure 3-9). Differences between orthotic and shoe-only conditions were used to determine the impact of the orthotics on clinical gait parameters and frontal plane gait mechanics in this sample of individuals.

Electromyography data were not included in the statistical analysis. Data from only five subjects were available and were not deemed appropriate for a quantitative







47
analysis. Mean and maximum values were used to detect trends in EMG outcomes as a result of orthotic intervention.

All variables that achieved statistical significance at a=. 10 are represented in the results and discussion. This a-level was chosen due to the experimental nature of this research. The functional level of individuals with chronic incomplete SCI required to participate in this research also resulted in recruitment of a small number of subjects. In order to determine valid conclusions about the effectiveness of orthotic intervention among a representative population of individuals with chronic incomplete SCI, a larger number of subjects would be needed. However, experimental research was necessary to justify the time and expense of a large sample and/or long-term study of this nature. The present research was intended to do just that.
Orthotic Shoe






More- I IN=9 1
Involved






Less- N9I =
Involved



Figure 3-6. 2x2 RM MANOVA layout.










I Orthoti


I Shoe


Figure 3-7. Univariate comparisons within the orthotic and shoe-only conditions.


Shoe I


II







Figure 3-8. Univariate comparisons within the more- and less-involved lower extremities.


N=9


-"---- - --
L~z~ L~N=9


MoreInvolved







LessInvolved


MoreInvolved







LessInvolved


I Orthotics _









Orthotic


LIII


LE~I1


- - - ----


L~I1


L~I~I1


i u i
Figure 3-9 Pairwise comparisons between sides and conditions.


MoreInvolved


LessInvolved


I Shoe














CHAPTER 4
RESULTS

The results of the MANOVA showed a significant main effect for the intervention

condition (Wilk's Lambda=0.010; F=5665.89; error df=1) and no significant main effect for the side condition (Wilk's Lambda=0.642; F=1.052; error df=1). There was also no significant interaction between the side and intervention conditions (Wilk's Lambda=0.143; F=29.035; error df=1). The results of univariate analyses for the intervention condition are reviewed below and are followed by pairwise comparisons of intervention results to determine specific effects on bilateral limbs. Descriptive changes in EMG are reviewed last.

Effects of Orthotic Intervention

The univariate analyses for intervention used analyzed to determine which variables were significantly affected by the orthotic intervention, regardless of limb. These differences depicted the general effect of orthotics on bilateral lower extremities. Sphericity-assumed p-values were used to determine significance because there was no violation of sphericity according to Mauchly's test of sphericity. Clinical Gait Variables

Univariate analyses revealed significant differences between the orthotic and shoeonly conditions on gait velocity, step duration, stance duration, ratio of stance, and ratio of swing (Table 4-1). Gait velocity was decreased by 8% during the orthotic condition (F=4.435; df=1,8). The step duration was lengthened by 1% when the orthotic was in use (F=6.530; df=1,8). The stance duration (F=5.614; df=1,8) and the ratio of stance (F=3.738; df=1,8) were also longer during the orthotic condition (2% and 1%,









respectively), and the swing duration was 1% shorter with the orthotics (F=3.738; df=1,8).

Table 4-1. Univariate analysis results for the clinical gait variables.
Orthotic Shoe-only
Variable Mean S.E. Mean S.E. p-value


Gait velocity 1.01 Stride length 1.34 Step length 1 (m) 0.670 Step length 2 (m) 0.667 Step width (m) 0.093 Toe clearance (m) 0.028 Toe-out angle (0) 8.7 Step duration (s) 1.318 Stance duration (s) 0.891 Swing duration (s) 0.427 Ratio of stance (%) 67.1 Ratio of swing (%) 32.9 Double limb support (%) 15.1 Single limb support (%) 37.0 S.E. = standard error; * significant at x =


0.082 0.055 0.028 0.027 0.009 0.003
1.2
0.092 0.079 0.020
1.2 1.2 1.3 1.3 10 level;


1.1 0.100 0.068*
1.34 0.058 0.923
0.671 0.029 0.905 0.667 0.030 0.956 0.094 0.010 0.801 0.028 0.003 0.780
8.2 1.5 0.347
1.299 0.085 0.034** 0.872 0.072 0.045**
0.427 0.020 0.867
66.7 1.2 0.089* 33.3 1.2 0.089* 14.7 1.2 0.213 37.2 1.3 0.484
** significant at a = .05 level.


Gait variable comparisons within the less-involved lower extremity

There was a significant difference between the orthotic and shoe-only conditions for step duration, stance duration, ratio of stance, and ratio of swing in the less-involved lower extremity according to pairwise comparison results (Table 4-2). The step duration was 2% longer with the orthotics in place. The stance duration and ratio of stance were both lengthened on the less-involved leg as a result of the orthotic intervention (3% and 1%, respectively). The ratio of swing was reduced by 1% during orthotic trials. Gait velocity and stride length were not included in pairwise comparisons as they are measures of a composite more- and less-involved lower extremity function. Gait variable comparisons within the more-involved lower extremity

Orthotics did not have a measurable effect on the more-involved lower extremity gait variables in this sample of individuals. No clinical gait variables were significantly changed in the presence of orthotics (Table 4-3). Though not significantly different, there was a









trend toward an increase in stance duration on the more-involved side dung the orthotic condition.

Table 4-2. Pairwise comparisons between the orthotic and shoe-only conditions for


the less-involved leg gait varibles.
Orhtotic
Mean
I1 (M) 0.663 0 2 (m) 0.659 C
(M) 0.091 C nce (m) 0.025 C gle (0) 8.0 on (s) 1.322 C ation (s) 0.914 C tion (s) 0.408 C since (%) 68.6 ing (%) 31.4 b support(%) 15.3
* support (%) 38.0


S.E.
.032 .027 ).011
).003 1.8 ).093
).081 ).019 1.3
1.3 1.3 1.5


Sh
Mean 0.674 0.656 0.101
0.024
7.7 1.294 0.887
0.407 68.0 32.0
14.9 38.2


oe-only
S.E.
0.032
0.032 0.010 0.003
2.0
0.085
0.074 0.019
1.3 1.3
1.2 1.2


p -value
0.099* 0.806 0.286
0.389 0.340
0.034** 0.030**
0.756 0.058*
0.058* 0.287 0.718


S.E. = standard error; * significant at c = .10 level; ** significant at ax = .05 level. Table 4-3. Pairwise comparisons between the orthotic and shoe-only conditions for


Variable Step length Step length Step width Toe clearan Toe-out ant Step durati( Stance dur, Swing durat Ratio of sta Ratio of swi Double lim Single limb S.E. = stan


the more-involved leg gait variables.
Orthotic
Mean S.
1 (M) 0.678 0.0 2 (m) 0.674 0.0
(M) 0.094 0.0 ice (m) 0.030 0.0 gle (0) 9.4 2 on (s) 1.314 0.0 nation (s) 0.869 0.0 tion (s) 0.445 0.0 nce (%) 65.6 1 ng (%) 34.4 1 b support (%) 14.8 1 support (%) 36.0 1


dard error


Shoe-only
Mean S.E.


.E.
30
31
09 04 2.1
91
77 22 .3
.3
.2
.3


0.668 0.678 0.087 0.032
8.7 1.304 0.858 0.446 65.4 34.6 14.6 36.3


0.031
0.030 0.012 0.004
2.3 0.087
0.071 0.022
1.2 1.2 1.2 1.5


p -value
0.329 0.581 0.292
0.214 0.458 0.220 0.189 0.922 0.329 0.329
0.206 0.359


Frontal Plane Joint Mechanics

When data for both legs were pooled, significant differences between the orthotic and shoe-only conditions were found in the maximum hip abduction angle, maximum frontal plane ankle power generation, and maximum frontal plane knee power absorption (Table 4-4). When compared to the shoe-only condition, the maximum hip abduction angle decreased by 16% during the orthotic trials (F=4.244; df=1,8), and the maximum


Variable
Step length Step length Step width Toe clearar Toe-out an Step durati Stance dur Swing dura Ratio of sta Ratio of sw Double lim Single limb









frontal plane ankle power generation was reduced by 22% (F=4.365; df=1,8). The

maximum knee power absorption was also decreased with the orthotic intervention

(F=3.570; df=1,8).

Table 4-4. Univariate results for the intervention kinematic and kinetic variables.
Orthotic Shoe-only
Variable Mean S.E. Mean S.E. p-value
Max ankle pronation angle (0) 14.9 2.2 15.2 2.2 0.746 Max knee adduction angle (0) -4.1 1.2 -3.6 0.9 0.297 Max hip adduction angle (0) 9.2 0.8 9.0 1.0 0.593 Max ankle pronation moment (Nm) 35.45 2.80 42.91 7.24 0.276 Max knee adduction moment (Nm) 18.79 3.57 24.59 6.98 0.170 Max hip adduction moment (Nm) 50.98 6.81 47.53 8.67 0.317 Max ankle power generation (W) 52.84 14.34 61.93 18.30 0.070* Max knee power generation (W) 22.48 5.20 26.22 7.54 0.214 Max hip power generation (W) 40.00 3.04 40.56 6.03 0.920 Max ankle supination angle (0) -3.9 1.6 -4.4 1.8 0.394 Max knee abduction angle (0) -16.9 1.6 -17.1 1.4 0.820 Max hip abduction angle (0) -3.0 0.6 -3.5 0.7 0.073* Max ankle supination moment (Nm) -14.61 2.09 -20.23 5.40 0.182 Max knee abduction moment (Nm) -26.91 2.28 -34.08 6.30 0.299 Max hip abduction moment (Nm) -42.86 9.00 -44.14 10.14 0.728 Max ankle power absorption (W) -53.13 15.90 -77.33 36.33 0.298 Max knee power absorption (W) -19.44 7.50 -22.23 7.30 0.095* Max hip power absorption (W) -35.18 6.86 -32.80 4.17 0.642 S.E. = standard error; * significant at a = .10 level.

Joint mechanics comparisons within the less-involved lower extremity

There were significant differences in the pairwise comparisons between the orthotic

and shoe-only conditions in the less-involved limb for the following variables: maximum

hip abduction angle, maximum frontal plane ankle power generation, and maximum

frontal plane knee power absorption (Table 4-5). The mean maximum hip abduction

angle was decreased by 24% on the less-involved side as a result of orthotic

intervention. The mean maximum frontal plane ankle power generation and knee power

absorption were also reduced during orthotic trials (23% and 34%, respectively). While

not achieving statistical significance, the maximum knee abduction moment was reduced


with the orthotic intervention.









Joint mechanics comparisons within the more-involved lower extremity

Orthotic intervention pairwise comparisons only revealed one frontal plane variable

that was significantly different in the more-involved limb. The maximum ankle supination

angle was significantly different between the orthotic and shoe-only conditions for the

more-involved extremity (Table 4-6). This supination angle was reduced by 70% with the

use of orthotics.

Table 4-5. Pairwise comparisons between the orthotic and shoe-only conditions for the
less-involved leg kinematic and kinetic variables.
Orhtotic Shoe-only
Variable Mean S.E. Mean S.E. p -value
Max ankle pronation angle (degrees) 12.9 3.8 12.7 3.6 0.851 Max knee adduction angle (degrees) -4.3 1.4 -3.5 1.1 0.253 Max hip adduction angle (degrees) 9.2 1.6 8.6 1.8 0.258 Max ankle pronation moment (Nm) 36.98 3.54 38.79 3.37 0.309 Max knee adduction moment (Nm) 17.06 3.19 18.50 3.96 0.387 Max hip adduction moment (Nm) 42.51 8.68 43.78 8.68 0.800 Max ankle power generation (W) 41.95 10.07 54.72 12.12 0.004** Max knee power generation (W) 14.44 2.17 17.54 1.97 0.197 Max hip power generation (W) 36.03 4.58 34.41 4.92 0.670 Max ankle supination angle (degrees) 7.1 3.7 6.7 3.4 0.705 Max knee abduction angle (degrees) 14.2 1.6 14.6 1.4 0.757 Max hip abduction angle (degrees) 2.6 1.4 3.4 1.5 0.032** Max ankle supination moment (Nm) 9.27 1.27 9.60 1.57 0.707 Max knee abduction moment (Nm) 27.95 3.20 31.33 3.17 0.117 Max hip abduction moment (Nm) 46.43 10.89 44.20 10.29 0.557 Max ankle power absorption (W) 32.42 7.27 34.96 5.21 0.614 Max knee power absorption (W) 9.32 1.45 14.11 2.93 0.043** Max hip power absorption (W) 29.94 5.90 30.65 2.85 0.900 S.E. = standard error; * significant at a = .10 level; ** significant at a = .05 level.

Electromyography

A qualitative analysis of mean and maximum EMG amplitudes during one stride was

performed in five subjects. An examination of mean and maximum EMG values for the

tibialis anterior (TA), medial gastrocnemius (MG), medial hamstrings (MH), and rectus

femoris (RF) during shoe-only and orthotic conditions revealed changes in muscle activity

in both the more- and less-involved lower extremities due to the orthotic intervention.









Table 4-6. Pairwise comparisons between the orthotic and shoe-only conditions in the
more-involved leg kinematic and kinetic variables.
Orthotic Shoe-only
Variable Mean S.E. Mean S.E. p -value Max ankle pronation angle (0) 16.9 3.4 17.7 3.6 0.562 Max knee adduction angle (0) -4.0 1.3 -3.7 1.0 0.514 Max hip adduction angle (0) 9.2 1.4 9.5 1.4 0.398 Max ankle pronation moment (Nm) 33.93 2.95 47.03 13.34 0.333 Max knee adduction moment (Nm) 20.51 4.37 30.67 11.19 0.210 Max hip adduction moment (Nm) 59.46 11.26 51.27 12.26 0.150 Max ankle power generation (W) 63.73 25.68 69.13 32.88 0.537 Max knee power generation (W) 30.51 10.23 34.90 15.59 0.461 Max hip power generation (W) 43.97 7.02 46.72 11.15 0.822 Max ankle supination angle (0) 0.6 3.5 2.1 3.8 0.049** Max knee abduction angle (0) 19.7 2.2 19.6 2.1 0.796 Max hip abduction angle (0) 3.4 1.1 3.7 1.3 0.545 Max ankle supination moment (Nm) 19.95 4.51 30.87 11.39 0.200 Max knee abduction moment (Nm) 25.88 1.69 36.83 12.66 0.408 Max hip abduction moment (Nm) 39.29 10.94 44.08 12.08 0.340 Max ankle power absorption (W) 73.83 30.02 119.69 70.41 0.317 Max knee power absorption (W) 29.56 14.34 30.34 13.80 0.715 Max hip power absorption (W) 40.41 10.63 34.94 7.45 0.634 S.E. = standard error; ** significant at a = .05 level.

Electromyographic Changes In the Less-involved Lower Extremity

Mean TA amplitudes were generally higher during the orthotic condition, while mean

MG activity showed little change in two subjects, a small increase in two other subjects,

and a small decrease in one subject (Figures 4-1 and 4-2). Mean MH amplitudes were

generally unaffected by orthotic intervention, while RF activity showed an increase in four

subjects and no change in the fifth (Figures 4-3 and 4-4). Of these muscles, only the

increases in mean TA and RF activity showed consistent changes in most subjects during

the orthotic condition.

Changes in TA maximum EMG activity showed a similar trend with orthotic

intervention as that seen among mean amplitudes (Figures 4-1 and 4-5). Four of the five

subjects demonstrated higher maximum values during the orthotic condition on the lessinvolved side. The remaining subject had a reduction in maximal TA activity (Figure 4-5).

Maximum MG activity varied among the five individuals. Three subjects exhibited little or









no change, while the other two subjects showed small decreases in activity on the lessinvolved limb with orthotic intervention (Figure 4-6). Little change was noted in maximal MH activity on the less-involved side in three subjects when orthotics were in use, one subject experienced a mild increase in maximum amplitude, and one subject demonstrated a substantial reduction in maximal activity (Figure 4-7). The use of orthotics had a mixed affect on maximum RF activity. One subject showed little change on the lessinvolved side during orthotic use, two subjects had small reductions, and two subjects showed increases in maximum RF activity (Figure 4-8). Maximum EMG values were more variable than mean amplitudes, with only the TA showing consistent changes (increases) in most subjects on the less-involved extremity due to the orthotic intervention. Electromyographic Changes In the More-Involved Lower Extremity

When going from the shoe-only to the orthotic condition, increases in mean TA activity were observed in four out of five subjects (Figure 4-9). Mean MG activity increased in three subjects and showed little change in two subjects during the orthotic condition (Figure 4-10). Increases in mean MH activity were noted in two subjects during the orthotic condition, while two were largely unaffected, and one showed a substantial reduction (Figure 4-11). Mean RF activity in the more-involved limb increased minimally in two subjects, moderately in two subjects, and decreased moderately in one subject when the orthotics were in use (Figure 4-12). These findings are very similar to those on the lessinvolved side, but changes were generally of smaller magnitude. The most marked exception was the variability in mean MH activity between the orthotic and shoe-only conditions. Unlike the minimal change accompanying the orthotic intervention on the lessinvolved side, mean MH amplitudes varied considerably with orthotics on the moreinvolved side.

Maximum EMG amplitudes on the more-involved lower extremity also changed during the orthotic intervention (Figure 4-13). The maximum TA amplitude increased during









orthotic use in the same four subjects as was seen when observing the mean activity levels. The other subject demonstrated a decrease in maximum TA activity on the moreinvolved side. Maximum MG activity was increased in three subjects during the orthotic condition, while one subject showed a decrease and another remained unchanged (Figure 4-14). The more-involved extremity MG maximum EMG amplitude during the orthotic condition was minimally affected in three subjects, decreased in one subject, and increased in another (Figure 4-15). Lastly, the RF maximum EMG amplitude was reduced in four subjects and increased in one (Figure 4-16).

U Orthotic U Shoe-only I

350 300
250 L 200
'~150 0100
50
0
BS RM CP JD CO Subject

Figure 4-1. Mean tibialis anterior EMG activity in the less-involved lower
extremity during orthotic and shoe-only conditions.










mOrthotic m Shoe-only


TI


... T


T.T


CP JD Subject


Figure 4-2. Mean medial gastrocnemius EMG activity in the less-involved
lower extremity during orthotic and shoe-only conditions.

UOrthotic m Shoe-only


TI


Subject

Figure 4-3. Mean medial EMG activity in the less-involved lower extremity
during orthotic and shoe-only conditions (n=5).


80
70 > 60
50 2 40 C 30 (D
0 10 0










I Orthotic N Shoe-only


BS RM


CP JD


Subject

Figure 4-4. Mean rectus femoris EMG activity in the less-involved lower
extremity during orthotic and shoe-only conditions.


UOrthotic E Shoe-only


1800 1600
1400 1200 1000 800
600 400 200
0


BS RM


CP JD
Subject


Figure 4-5. Maximum tibialis anterior EMG activity in the less-involved lower
extremity during orthotic and shoe-only conditions.


200 180 160 140 120 100
80 60 40 20
0


TT


T T


I T










U Orthotic U Shoe-only


1200


100
0 800 E 600 E 400 5 200


T1


LT1


LI


BS RM CP JD CO Subject

Figure 4-6. Maximum medial gastrocnemius EMG activity in the lessinvolved lower extremity during orthotic and shoe-only
conditions.


= Orthotic U Shoe-only


1600
1400 1200 1000 800 600
400 200
0


CO


Subject


Figure 4-7. Maximum medial hamstring EMG activity in the less-involved
lower extremity during orthotic and shoe-only conditions.


I-L










U Orthotic m Shoe-only


L T


T.


0 - .i_BS RM CP JD CO Subject

Figure 4-8. Maximum rectus femoris EMG activity in the less-involved lower
extremity during orthotic and shoe-only conditions.


HOrthotic m Shoe-only


250

200 CD 150 c 100 (U


TI I.


TT


Subject

Figure 4-9. Mean tibialis anterior EMG activity in the more-involved lower
extremity during orthotic and shoe-only conditions.


1200 S1000 O 800 u 600 E 400
x
S200









lOrthotic U Shoe-only


150 1-


Subject


Figure 4-10.


Mean medial gastrocnemius EMG activity in the more-involved lower extremity during orthotic and shoe-only conditions.


U Orthotic n Shoe-only


T T


T T


Subject


Figure 4-11.


Mean medial hamstring EMG activity in the more-involved lower extremity during orthotic and shoe-only conditions.


200


IL i










m Orthotic m Shoe-only


T T


T T


RM CP
Subject


Figure 4-12.


Mean rectus femoris EMG activity in the more-involved lower extremity during orthotic and shoe-only conditions.


mOrthotic m Shoe-only


1200 5 1000 O 800 LE 600 E 400 M 200
0


T T


TT


CP JD Subject


Figure 4-13. Maximum tibialis anterior EMG activity in the more-involved
lower extremity during orthotic and shoe-only conditions.










U Orthotic U Shoe-only


v.i


CP JD Subject


Figure 4-14.


Maximum medial gastrocnemius EMG activity in the moreinvolved lower extremity during orthotic and shoe-only conditions.


KOrthotic EShoe-only


_LT


BS RM


LT


JD CO


Subject

Figure 4-15. Maximum medial hamstring EMG activity in the more-involved
lower extremity during orthotic and shoe-only conditions.


1200 1000 800 600
400 200

0


TT


TT


500 450 400 L 350 2 300 Lu 250
E
= 200 .E 150 (= 100
50 0










n Orthotic U Shoe-only


RM CP
Subject


Figure 4-16. Maximum rectus femons EMG activity in the more-involved
lower extremity during orthotic and shoe-only conditions.


1200 1000 800 600

400 200

0


TP


L.T















CHAPTER 5
DISCUSSION

All of the nine subjects who participated in this research had an excessive forefoot varus bilaterally and six had a hypomobile rearfoot valgus bilaterally. Had these individuals presented without neurologic injury to an orthopaedic clinic, it is likely that they would have been measured for and fitted with orthotics to improve their lower extremity mechanical alignment. Because the medical paradigm for prescribing orthotics has not yet shifted to include those with neurologic injury, these individuals are typically overlooked. Results of this study indicate that orthotics have an immediate effect on specific clinical gait characteristics and frontal plane joint mechanics of this small subject sample and suggest that ambulatory individuals with chronic incomplete SCI may benefit from foot evaluation and appropriate orthotic placement. This study represents the first known attempt to determine the effects of orthotic intervention among ambulatory individuals with chronic incomplete SCI.

Orthotic Intervention

Statistical analyses revealed that orthotics had a measurable effect on specific clinical gait parameters and frontal plane joint mechanics in this group of individuals with chronic incomplete SCI. The following sections discuss the greater impact of orthotics on the lessinvolved leg and the lesser impact on the more-involved lower extremity. Clinical Gait Variables

Five clinical gait measures were significantly affected by the use of orthotics. Step duration, stance duration, and ratio of stance experienced increases, while gait velocity and ratio of swing were reduced in the orthotic condition. The magnitude of change for









each of these variables was minimal and not clinically significantly, but such immediate changes in these gait parameters may be an indicator of future adaptations. The presence of these immediately measurable changes justifies the need for future research that includes follow-up testing after several weeks of orthotic use. Follow-up testing and a larger sample size would help to discern the long-term benefits of orthotic use in this population.

Univariate statistics showed that gait velocity was decreased during the orthotic

condition. While no orthotic research was found that has studied the effect of orthotics on gait velocity, an increase was anticipated due to the function of the forefoot posting. The medial forefoot posting was placed to provide support for a hypermobile forefoot varus. that had limited ability to form an effective and/or timely lever for push-off during terminal stance. One of the functions of orthotics is to bring the ground up to the foot instead of allowing the foot to become overly compliant with the ground. Improving the effectiveness of push-off should, theoretically, have led to increases in stride length and gait velocity. In these subjects, however, this was not the case. It is likely that the short period between orthotic placement and testing did not allow each individual enough time for sensory accommodation. The altered sensory stimulus may have caused these subjects to be hesitant during gait. Further testing would be beneficial to determine the effect of orthotics on gait velocity after a longer period of accommodation. Clinical gait comparisons within the less-involved lower extremity

Differences in less-involved lower extremity clinical gait characteristics between the

orthotic and shoe-only conditions were limited to measures of phase duration. Increases in step duration, stance duration, and ratio of stance all referred to the lengthened period of time that the less-involved foot was in contact with the ground. Ratio of swing, having a direct relationship with ratio of stance (ratio of stance + ratio of swing = 100% gait cycle), showed an expected decrease. The only research report found that discusses changes in









the stance duration of gait as a result of orthotic placement was a gait study among children with Down Syndrome.72 The authors did not report significant changes in stance duration as a result of orthotic use but did find a decrease in the inter-trial variability of stance duration with orthotics. A review of stance duration standard errors in the current study revealed a very small increase in stance duration variability resulting from orthotic placement (Table 4-2). This finding is in direct opposition to that of Selby-Silverstein and colleagues.72

Clinical gait comparisons within the more-involved lower extremity

Orthotics did not have a measurable impact on the more-involved lower extremity gait variables in this sample of individuals. This result was not entirely unexpected given the neurologic capacity of the more-involved limb versus that of the less-involved limb. The more-involved lower extremity in all subjects demonstrated greater strength and sensory impairments than the less-involved leg. Altered sensation may not have allowed for adequate awareness of the placement of the orthotics and, subsequently, may not have triggered enough of a response to alter the gait characteristics of the limb. Impaired strength, on the other hand, may have reduced the individual's physical ability to make adjustments to their altered mechanical alignment. Frontal Plane Joint Mechanics

Selected hip, knee, and ankle kinematic and kinetic characteristics were significantly affected by the orthotic intervention. Maximum hip abduction angles were decreased, and maximum knee power generation and maximum ankle power absorption were reduced when orthotics were used. Pairwise comparison results showed that the less-involved lower extremity was influenced to a greater degree than the more-involved lower extremity.









Joint mechanics comparisons within the less-involved lower extremity

A 24% decrease in the maximum hip abduction angle was noted on the less-involved limb during the orthotic condition. Graphic representation of the average frontal plane hip range of motion of one subject over five trials shows this reduction occurring during the stance phase of gait (Figure 5-1). During single limb support, while the foot is fixed in a closed kinetic chain relationship to the trunk segment, frontal plane hip range of motion varies as a function of lateral pelvic tilt and thigh position. A lateral lean toward the stance limb was typically employed by most of the subjects during the swing phase in order to clear the contralateral (swing) foot from the floor. When the trunk laterally flexes and the pelvis tilts laterally to a greater degree over the stance limb, a larger abduction range of motion results.

Frontal plane hip angles during gait among able-bodied subjects have a very different pattern than the one found here. During initial stance the hip is adducted an average of 100 in normal gait.70 During the shoe-only condition this group of subjects did not adduct the less-involved limb at all during initial stance. They averaged 10 of abduction until orthotics were added, at which point, they achieved 1-20 of adduction in early stance. By mid stance, normal range of motion values reach 150 of adduction. In these subjects, however, the hip angle progressed into abduction and peaked in early mid stance at 70 during the shoe-only trials and 50 during orthotic trials.

While the impact of the orthotics is not large in its magnitude of change, one must remember that the orthotics were in place for only 30-45 minutes prior to testing. Further research is justified to perform a longer-term study that tests this population immediately and after four to eight weeks of orthotic use. In this group, specifically, immediate decreases in hip abduction during orthotic intervention trials had a very positive effect. Due to the lateral lean employed by the majority of the subjects to clear the contralateral limb








70
from the floor, lower maximum hip abduction angles indicate that there was a decrease in lateral flexion over the less-involved limb during the stance phase. A decrease in lateral flexion could partially be the result of more appropriate contralateral limb toe clearance during swing. Though it did not achieve statistical significance, five of the nine subjects actually demonstrated a decrease in toe clearance height on the more-involved side during the swing phase while using orthotics. This finding is consistent with the decrease in lateral trunk flexion over the less-involved limb during stance. Reducing lateral trunk flexion over the less-involved leg brings the more-involved foot closer to the ground, and toe clearance height is thus reduced. If the reduction in the amount of lateral flexion exceeds the corresponding change in toe clearance height, however, another mechanism must be adopted.

-Orthotic -Shoe

10
_e 8
6
4 Adduction
� 2
*0
0\
10 -2 20 30 40 50 60 70 80 90
S -4
o -6 Abduction
u -8
-10
% Gait Cycle

Figure 5-1. Less-involved limb frontal plane hip angles from one subject.

A review of sagittal plane hip flexion angles, though not included in this analysis,

revealed an increase in maximum flexion angles in six subjects during the swing phase of gait in the more-involved leg. This suggests that the subjects in this study altered their lower extremity compensatory motions in favor of a more normal pattern of sagittal plane hip flexion to clear the foot from the floor. Though the changes in joint angles were small








71
in magnitude, they represent an immediate and distinct alteration in gait strategies as a result of orthotic intervention.

A significant decrease (23%) in frontal plane ankle power generation in the lessinvolved limb was noted during early stance (Figure 5-2). Inspection of the power curves for the orthotic and shoe-only conditions reveals that there was a smaller magnitude of change from peak power generation (-10% gait cycle) to peak power absorption (-16% gait cycle) in the orthotic condition. Joint power is the product of joint angular velocity and the resultant moment of the joint. Angular data indicate that the time rate of change of ankle angle (angular velocity) from supination to pronation is relatively similar for the two conditions. This finding fails to support reports in the literature that pronation velocity decreases with orthotic intervention.45 s Orthotic -Shoe

100
0 80
0
0. 60 Generation
40
3: 20

-20 10 40 50 60 70 80 90
0 -40
. Absorption
LL -60 Asrto

% Gait Cycle

Figure 5-2. Less-involved lower extremity frontal plane ankle power from one subject. Data represent the average of five trials per
intervention.

A graph of the frontal plane ankle joint range of motion from one subject shows that the slopes of both orthotic and shoe-only curves are almost identical at approximately 10% of the gait cycle (Figure 5-3). Because the curves represent the change in joint angle over time, the slope of the line is descriptive of frontal plane ankle joint angular








72
velocity. Lack of variation between the two conditions indicates that any change in power generation is due to differences in the joint moment.

One subject's average frontal plane ankle joint moments were graphed for each

intervention to illustrate the change that caused the reduction in power generation (Figure 5-4). The pronation moment was found to be lower for the orthotic condition, suggesting that the smaller moment was primarily responsible for the lower power generation with the orthotic intervention. Changes in frontal plane ankle moment and power as a result of orthotic placement have not been reported previously.

Orthotic -Shoe

6

- SPronation



0
2
0
00
-20 30 40 50 60 7 80 0


0 Supination

-10
% Gait Cycle
Figure 5-3. Less-involved lower extremity frontal plane ankle angle from one subject. Data represent the average of five trials per intervention.

There was a significant decrease (34%) in maximum knee power absorption as a result of orthotic intervention (Figure 5-5). As with ankle power generation, the two components that determine knee joint power are joint angular velocity and joint moment. During terminal stance (-60% gait cycle), when the maximum knee absorption power was most notably different between conditions, the knee angle was progressing into abduction (Figure 5-6) with an adduction moment of force (Figure 5-7). An absorption power indicates that the knee was being eccentrically controlled late in the stance phase (i.e., slowing the progression of the knee into abduction).











-Orthotic - Shoe

E 40
S35
Z
o30
o 25 Pronation
=20
15 o 10
5
0
10 20 30 40 50 70 80 90
U -10
% Gait Cycle

Figure 5-4. Less-involved lower extremity frontal plane ankle moment from one subject. Data represent the average of five trials per
intervention.


-Orthotic -Shoe

20
15


S0
1 01 Adduction
0



0 -1 0

ine-o Abduction
C -20
0
U. 25
% Gait Cycle

Figure 5-5. Less-involved lower extremity frontal plane knee power from
one subject. Data represent the average of five trials per
intervention.

During late stance the COP is acting through the forefoot. A GRF that acts lateral to

the ankle joint center causes a pronation moment that forces the foot into pronation. Foot

pronation has been shown to increase tibial internal rotation angles' 53 and femoral

internal rotation angles.80 A combination of internal tibial and femoral rotation will

necessarily increase the knee abduction angle due to the congruency of the tibial and








74
femoral condyles. From the graph of knee abduction angle (Figure 5-6), it appears that the abduction angular velocity (the slope of the lines) did not change between the orthotic and shoe-only conditions. Thus, the change in knee power absorption must have been the result of different frontal plane moments at the knee joint.

The frontal plane knee moment curves (Figure 5-7) reveal that there was a reduction in the knee adduction moment late in the stance phase. This could have been caused by a decrease in the magnitude of the GRF and/or the moment arm length between the knee joint center and the point of application of the GRF. Regardless of the cause, the result was a decrease in the adduction moment that was required to control the knee's progression into abduction during late stance. A reduction in the frontal plane moment should, theoretically, result in less compressive force placed on the lateral condyles of the tibia and femur. A smaller compressive force, in turn, would reduce the stress (force per unit area) on the articular cartilage of the knee. Less stress on the joint's passive soft-tissue structures over time could result in decreased wear and tear of the joint and even reduce pain.


Orthotic -Shoe

4
2 Adduction
�< 0
0
0 30 40 50 60 70 80 90
"�-4

a' -6
= -8 Abduction
-8
U- -10
-12
% Gait Cycle

Figure 5-6. Less-involved lower extremity frontal plane knee angle from one subject. Data represent the average of five trials per
intervention.











Orthotic -Shoe

E 15
Z10
C Adduction
E 0^
0 A
; -5 10 20 30 40 50 60 70 80 90
c:-10
-15
.2 -20
-20 Abduction
~-25 2 -30
- 35
% Gait Cycle

Figure 5-7. Less-involved lower extremity frontal plane knee moment from
one subject. Data represent the average of five trials per
intervention.

Joint mechanics comparisons within the more-involved lower extremity

Orthotic intervention did not have as great an effect on the more-involved limb as it did on the less-involved limb. Only the maximum ankle supination angle was significantly different between the orthotic and shoe-only conditions for the more-involved lower extremity. This supination angle was reduced 70% with the use of the orthotic. As previously mentioned, the relatively limited impact of the orthotic on the more-involved limb was not altogether surprising considering the type and degree of neuromuscular deficits noted within the extremity. Because the neuromuscular system responsible for the limb function was not as sensitive to or able to adequately respond to minor alterations in mechanical positioning, segments higher in the chain may have been unable to respond to the same degree. Peripheral sensory deficits result in less afferent information that is available to the spinal cord and consequently influences reflexive or voluntary output to the musculature. Weakness in the more-involved musculature also affects outcomes secondary to the inability to increase or decrease muscle activity to









optimal levels. The fact that the orthotic was able to produce an immediate change in frontal plane ankle alignment, however, is quite promising.

Symmetry

When discussing the effect that orthotics have on specific gait variables, it is also necessary to discuss the bilateral symmetry of these variables with and without orthotic intervention. Gait is a product of symmetric segmental motions and any decrease in asymmetry that results from orthotic intervention should be seen as a positive outcome. There were several variables that did not achieve significance in the statistical analysis that did, nonetheless, demonstrate mean changes between the two intervention conditions.

Gait Symmetry

Variations in lower extremity gait symmetry between the two interventions can be used to deduce benefits and drawbacks of orthotics in this population. In the following graphs, the position of each line's intersection with the vertical axis can be used to determine the degree of symmetry between orthotic and shoe-only conditions. The closer the two points are together on an axis, the greater the symmetry. Step width and toe clearance height are two clinical gait parameters that demonstrated greater lower extremity symmetry during the orthotic condition.

There was a large reduction in step width on the less-involved side and an increase of similar magnitude on the more-involved side when orthotics were in use (Figure 5-8). Step width is the distance between the two heel markers during double limb stance (Figure 3-7). This distance is calculated during the double limb stance period that follows the test foot contact with the force plate. Though the magnitudes of change between intervention conditions for each lower extremity were small (more-involved: 0.69cm difference, lessinvolved: 0.94cm difference), the step width difference between bilateral extremities was noticeably smaller during the orthotic condition (orthotic: 0.28cm difference, shoe-only:








77
1.34cm difference). Alterations in the GRF's, line of progression, path of the COM, and other variables could have contributed to greater step width symmetry.

.102 .100 r

.098

E .096

.094

CD .092

090 Side

.088 3 More-involved .086 13 Less-involved
Orthotic Shoe Intervention
Figure 5-8. Step width maximum values for both sides and interventions.

Toe clearance height also showed greater symmetry during the orthotic condition

(Figure 5-9). As was seen with step width, there was a reduction in toe clearance height on one side as a result of orthotic intervention and an increase on the other with relatively small magnitudes of change (less-involved: 0.18cm difference, more-involved:

0.25cm difference). Here again, however, there was a greater disparity between extremities within the shoe-only condition (orthotic: 0.45 difference, shoe-only: 0.89 difference). No literature was located that supports or refutes greater toe clearance symmetry resulting from orthotic intervention.

While step width and toe clearance were both more symmetrical during the orthotic condition, step duration was actually more asymmetrical after the orthotic placement (Figure 5-10). One possible explanation for a loss of symmetry after orthotic placement








78
lies within the less-involved limb's ability to better accommodate to changes when compared to the more-involved extremity. Superior muscular and sensory capabilities on the less-involved side may have enabled these subjects to adapt and respond to the placement of the orthotic.

034


.032


E .030
()
0
.028
a)

o .026
I-

Side
.024
1 More-Involved .022 1 0 Less-Involved
Orthotic Shoe Intervention Figure 5-9. Maximum toe clearance height for both sides and interventions.

In their research of various gait parameters among individuals after CVA, Yavuzer and Erging� actually discovered immediate increases in the hemiparetic (more-involved) limb stance duration with an arm-sling intervention. The arm sling could not have altered the sensory or muscular ability of either leg, yet it caused an increase in stance time on the hemiparetic side. These investigators also discovered that the arm sling reduced the medial-lateral excursion of the center of mass (COM) and may have been the key to the longer stance duration on the hemiparetic side.










.92 .91

90


0
. 89


.88
C
(13
.87Side

.86 - More-Involved .85 0_ Less-Involved
Orthotic Shoe Intervention
Figure 5-10. Stance duration for both side and intervention conditions.

Use of a sling to place the arm across the torso decreases the arm's effective

perpendicular distance from midline, thus shifting the COM slightly toward the contralateral side. This shift necessitates a smaller medial-lateral translation from the line of progression toward the less-involved side to maintain upright balance. It also requires a greater mediallateral translation toward the more-involved side for the COM to pass over the foot. Because this distance is greater than that without the arm-sling intervention, more time would be necessary to travel the distance to and from the line of progression (assuming no change in velocity) and would be measured as an increase in stance duration. Center of mass variables were not included in the present study and, therefore, cannot be addressed directly. It is possible, given Yavuzer and Ergin's9� findings, that analysis of medial-lateral COM translation toward the more-involved leg would not reveal any significant changes between orthotic and shoe-only conditions because orthotics were unable to immediately increase the stance time on that limb.









Frontal Plane Joint Mechanics Symmetry

Improved proprioception and pain perception help the musculoskeletal system to control joint motion through stabilization and acceleration/deceleration. For example, during mid stance when the knee reaches terminal extension, proprioceptors signal the spinal cord to trigger a response from the surrounding musculature to facilitate cocontraction and prevent hyperextension.21 If an individual has a sensory impairment and lacks appropriate proprioception, the protective response to reduce knee extension is not elicited in time, and hyperextension occurs. Additionally, if an individual has a flexible flat foot, he/she would undergo excessive tibial internal rotation and knee abduction during weight bearing.4753, 8" Greater knee internal rotation and abduction angles place both the posterior capsule of the knee and medial collateral ligaments at risk for developing laxity. Knee abduction and internal rotation also place greater compressive and shear forces on the lateral femoral and tibial condyles.

Alternatively, if a person has appropriate proprioceptive awareness but not the necessary strength to respond to joint feedback, then a similar dilemma develops. Individuals with quadriceps weakness often lean forward and thrust their knees into hyper-extension, thereby eliminating the need for muscular control of the joint, relying instead on the ligaments of the knee to maintain the extended position. Here again, an additional excessive forefoot varus creates greater tibiofemoral internal rotation and knee abduction, placing even greater stress on the ligamentous structures of the knee. Over time, both poor sensation and poor muscular control increase the risk of joint laxity, articular cartilage damage, and eventual deformity.

There were two frontal plane moments that demonstrated greater symmetry between the intervention conditions. The bilateral difference between maximum ankle supination moments in the shoe-only condition was twice that of the orthotic condition (Figure 5-11), and the difference between bilateral maximum ankle pronation moments was three times








81
greater in the shoe-only condition (Figure 5-12). Interestingly, these figures clearly show

that the more-involved leg had greater magnitudes of change as a result of orthotic

intervention, yet statistically there were no significant differences found here. A review of

descriptive statistics (Tables 4-5 and 4-6) reveals that there was greater variability on the

more-involved side than the less-involved side and, therefore, no significant differences

were found.

In both these examples, there was a reduction in the joint moments as a result of

orthotic intervention. Here again, regardless of whether the reduction was due to a

decrease in the GRF magnitude and/or a decrease in the moment arm distance, a smaller

joint moment is a positive outcome. Decreasing the moment of resistance of the joint will

help decrease the stresses on the joint structures.


48

E 46
z C
G 44
E 0
C 42
C 40
2
ai)3
: 38
C
< E
E 36 S
"3 Side
E
'M 34 0]
3More-Involved

32 0_ Less-Involved
Orthotics Shoe

Intervention Figure 5-11. Maximum ankle supination moment for both sides and interventions.











E
Z -26 E 28
0 0
-30

< -32
W
-o I


E -34
E
' -36 More-Involved

-38 Less-Involved
Orthotic Shoe Intervention
Figure 5-12 Maximum ankle pronation moment for both sides and interventions.

Electromyography

Due to a small sample size and lack of statistical analysis, no solid conclusions are drawn based on EMG data. Instead, consistent trends between the orthotic and shoeonly conditions are noted. Each of these trends is discussed in light of available research, and the possible impact of each is explored.

Qualitative analysis of EMG changes in the more- and less-involved lower extremity during both orthotic and shoe-only conditions revealed consistent differences only in the mean and maximum TA muscle activity. Mean RF activity was the only other muscle to demonstrate a consistent change with orthotic intervention, but it was limited to the lessinvolved lower extremity. Mean MH amplitudes in the less-involved lower extremity, alternatively, were not substantially altered in any subject during orthotic use. The remaining muscles of interest demonstrated variable responses to orthotic intervention.









Mean and maximum TA amplitudes were increased during the orthotic condition in both the more and less-involved lower extremities. An increase in mean TA activity is consistent with findings reported by Nawoczenski and colleagues in a small sample of runners with structural malalignment of the foot. These researchers found that orthotic intervention caused an increase in mean TA activity during the first 50% of running stance. It was suggested that the increase in mean activity may have been due, in part, to the TA's role in resupination of the foot just before terminal stance, and that an earlier conversion from eccentric to concentric control may account for the noted increase in mean amplitude.54 Closer inspection of one subject's mean TA activity on the lessinvolved lower extremity indicates that this explanation may not hold true in this sample of ambulatory individuals with chronic incomplete SCI.

Anterior tibialis activity from one stride during each of five walking trials was

smoothed, rectified, averaged, and graphed for both the orthotic (Figure 5-13) and shoeonly (Figure 5-14) conditions. Results indicate that while the eccentric control during loading response was of similar mean magnitude and duration for both interventions, the onset of concentric activity during terminal stance actually occurred later during the orthotic condition. In this individual, the larger mean TA amplitude during one stride resulted from a larger maximum TA amplitude during initial swing, higher levels of activity throughout the remainder of swing, and a longer stride time recorded during the orthotic condition (1.04 s versus 1.00 s).

Possible explanations for increases in maximum TA activity during the orthotic

condition include the rigidity of the posting material, eccentric control of the foot, the role of the muscle in early stance, and orientation of the foot in terminal stance. Komi and colleagues found that increasing the rigidity of a shoe's heel counter caused an increase in EMG amplitude just prior to and at heel strike. The subject presented here actually had lower TA amplitudes at initial contact, which were followed by a rapid








84
increase during loading response. Though this subject did not exhibit higher maximum TA amplitudes in early stance, the timing and magnitude of activity in other subjects may have differed somewhat.


0 10 20 30 40 50 60 % Gait Cycle


Figure 5-13.




700
600
> 500
400 - 300 2 200
100
0-L


70 80 90 100


Mean tibialis anterior EMG activity of one subject on the less-involved side during the orthotic condition. Data represent the average of five strides.


0 10 20 30 40 50 60 70 80 90 100 % Gait Cycle

Figure 5-14. Mean tibialis anterior EMG activity of one subject on the
less-involved side during the shoe-only condition. Data
represent the average of five strides.

Nawoczenski et al and Komi et al34 have suggested that the presence of an

increase in TA activity during early stance may have been necessary to eccentrically









control the forefoot's rapid descent to the floor when orthotics were in use. Other researchers,45, 9, 67 however, have shown that orthotic intervention reduces foot pronation velocity during early stance (a finding not supported among these subjects). A decrease in pronation velocity should, theoretically, reduce the eccentric TA activity needed to control the foot during early stance. For the subject presented in Figures 5-13 and 5-14, there was an initial decrease in TA activity followed by an increase during loading response to the level measured in the shoe-only condition. The lower mean TA activity was recorded during the first 6% of the gait cycle, while the foot was being lowered to the floor, and the increase in activity occurred when the foot was flat and was loaded with body weight. The spike in TA activity in early stance during the orthotic condition may also have been the transition from an eccentric to a concentric role. A delay in the orthotic condition may represent a more effective push-off from the contralateral leg that delayed the need for concentric TA activity. This suggests that the changes in early mean TA activity in this subject may have been caused by a combination of concentric activity to pull the tibia anteriorly and by the altered sensory stimulus resulting from the rigid orthotic.

An increase in mean and maximum TA activity was seen during the swing phase in the same subject. Nawoczenski and colleagues proposed that this type of result may be due to a change in foot position during terminal stance. Several studies of orthotic intervention and its effect on gait kinematics have demonstrated a reduction in maximum pronation during stance.4 '67,79 Nawoczenski and associates54 have suggested that a reduction in the maximum pronation angle may place the TA in a position of greater mechanical advantage that leads to an increase in activity. Additionally, extrapolation of Komi and associates'34 work suggests that increases may also be a function of the rigid forefoot posting material. It is possible that intrinsic placement of high-density forefoot posting may









cause an increase in TA activity during terminal stance that carries over into the swing phase. No research has been found that addresses this possibility.

Verification of each of these potential explanations would require experimental

research focused on each specific issue and remains beyond the scope of this study. Regardless of the cause of increased swing phase TA activity, however, this population of individuals would benefit from greater dorsiflexion range of motion during the swing phase of gait to reduce the pathological compensatory mechanisms employed to clear the foot from the floor. Other joint motions that would affect swing phase mechanics in this group of subjects are hip and knee flexion.

Greater hip and knee flexion angles may decrease the need to circumduct the leg or hike the hip (both of which were used by this group of subjects) to provide an adequate toe clearance. Higher mean RF activity was noted on the less-involved side in four of the five subjects while no gross changes were seen in the mean MH activity. The RF is a two-joint muscle that acts as a knee extensor as well as a hip flexor. Graphical representation of one subject's RF EMG (Figures 5-15 and 5-16) demonstrates that the maximum RF activity was higher during early stance in the shoe-only condition, but that the mean activity was greater in the orthotic condition due to its more regular activity throughout the stance phase. The same trend was noted during the swing phase of gait. In addition to the more uniform activity seen during stance and swing, the timing of the RF activity varied notably between the two conditions. During the orthotic condition, the RF became most inactive at approximately 40% of the gait cycle and then increased near initiation of the swing phase at 60% of the gait cycle (Figure 5-15). Activity during the shoe-only condition was high during early stance, tapered off to a quiet state at approximately 40% of the gait cycle, and did not increase again until nearly 78% of the gait cycle (Figure 5-16).











800 700 600
500 2 400 to 300
200
100
0
0 10 20 30 40 50 60 70 80 90 100
% Gait Cycle

Figure 5-15. Mean rectus femoris EMG activity from one subject on the
less-involved side during the orthotic condition. Data represent
the average of five strides.

Neither the orthotic nor the shoe-only RF EMG graphs have normal phases or

amplitudes of activity. Of the two conditions, however, the orthotic condition more closely approximated the normal patterns described by Winter87 (Figure 5-17). During normal gait, the peak RF activity is reached during loading response as the muscle works to control knee flexion. Activity then tapers throughout the remainder of stance while the muscle assists in extending the knee. Once the knee is extended and the body's center of mass is passing over the foot, the RF activity decreases to minimal levels. This is followed by an increase in pre-swing activity when hip flexion is initiated and then tapers off slowly as the swinging foot and leg are decelerated. The final increase in RF activity occurs during terminal swing as the knee is extended in preparation for initial contact.87

During the orthotic condition peak RF activity did occur during the stance phase, but more importantly, the timing of activity at swing initiation was more appropriate than during the shoe-only condition. This earlier activation should improve hip flexion motion and decrease the compensatory mechanisms necessary to clear the foot from the floor







88
during the swing phase. There are no known experimental data that address the effect of

orthotics on RF activity to corroborate this finding.


800 700 600
*- 500 2 400
W
� 300 200
100

0
0 10 20 30 40 50 60 70 80 90 100
% Gait Cycle

Figure 5-16. Mean rectus femoris EMG activity from one subject on the
less-involved side during the shoe-only condition. Data
represent the average of five strides.


.''
U
* S a *
* a


so "" """'''

Figure 5-17. Rectus femoris EMG during
(n-16).Adapted from Winter.

In an experiment performed by Duysens and Pearson, 1 lower limb flexor activity


during treadmill walking was decreased in premammillary decerebrated felines as the result of a constant stretch placed on the triceps surae. Maintaining a constant stretch on the triceps surae mimicked continuous loading of the limb. When the nervous system did not interpret unloading of the test hindlimb, all flexor activity in that limb first decreased









and then ceased altogether. When all flexor activity was abolished, the hindlimb remained in a position of full extension while the feline ambulated on three legs. Duysens and Pearson15 speculated that proprioceptors in the triceps surae inhibited the rhythmic activity created by the CPG.

In the current sample of ambulatory individuals with SCI, mean RF hip flexor activity was increased during the orthotic condition primarily because it had an earlier onset during the swing phase of gait than that seen during the shoe-only condition. An increase in TA activity as a result of orthotic intervention, as discussed previously, was also seen when the orthotics were in use. It is possible that the increase in flexor activity was a direct result of altered triceps surae afferentation acting on the flexor burst generators of the CPG. Obviously, the present research was not designed to explore these findings. A study focused specifically on the impact of orthotics on EMG and central pattern generation would need to be performed to adequately address this proposition.

Summary and Conclusions

Orthotic inserts have played a significant role in the treatment and rehabilitation of orthopaedic dysfunction among able-bodied individuals for years. They have been used both as a preventive measure for correcting poor mechanical alignment and in response to pain caused by dysfunction. Much of the research that has focused on determining the impact of orthotic intervention has primarily been concerned with the quantitative measurement of rearfoot motion. Few studies have looked at the impact of this intervention on frontal plane lower extremity joint mechanics and on the functional parameters of gait. This study represents the first attempt to measure immediate changes in clinical gait parameters and frontal plane mechanics in a sample of ambulatory individuals with chronic incomplete SCI. Despite the small sample size, the







90
current study demonstrates that several of these parameters can be altered immediately by orthotic intervention in ambulatory individuals with chronic incomplete SCI.

The findings presented justify further research in this area. Orthotics had an

immediately measurable effect on clinical gait parameters and kinematic and kinetic parameters, as was hypothesized, in a small sample size of this population. Increasing the sample size would increase the internal and external validity of the results. Longitudinal data, on the other hand, would provide necessary information regarding long-term effects on the gait mechanics in this population. Immediate changes that were noted in the less-involved lower extremity in this subject sample were hypothesized to be the result of greater sensory and muscular function in this limb. Post-testing this group after a longer period of time would provide more conclusive results regarding the effect of orthotics on the more-involved lower extremity and the time necessary to see these results.

Other future research opportunities include the study of orthotic intervention in other neurologically involved populations. Individuals with multiple sclerosis, Gillian-Barre, cerebrovascular accident, and head injuries are examples. Any intervention that has the potential to improve quality of life should be investigated in detail. Thus far, orthotic intervention has been found to be an inexpensive and efficient method of improving mechanical alignment of the lower extremities.

As with any research, there are many limitations to this study. The primary limitations are the sample size and the variability in this subject population. Secondary limitations include those typical to biomechanics, such as marker movement, surface electrode noise caused by skin movement, and accuracy of the calibrated space and the digitizing process. These sources of error were controlled to the greatest extent possible in a human experimental setting.















APPENDIX A CONSENT FORM














IRB#


Orthotic Gait Analysis: Adults, 18 years and older

Informed Consent to Participate in Research


You are being asked to take part in a research study. This form provides you with information about the study. The Principal Investigator (the person in charge of this research) or a representative of the Principal Investigator will also describe this study to you and answer all of your questions. Before you decide whether or not to take part, read the information below and ask questions about anything you do not understand. Your participation is entirely voluntary.


1. Name of Participant ("Study Subject")




2. Title of Research Study

Do Orthotics Have an Immediate Impact on the Gait Mechanics of Individuals with
Chronic Incomplete Spinal Cord Injury?


3. Principal Investigator and Telephone Number(s)

Kristen Jagger MS, PT
(352) 392-0580 ext. 1321 (University of Florida)
(678) 642-3642 (Georgia State University)




Full Text

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DO FOOT ORTHOTICS HAVE AN IMMEDIATE EFFECT ON THE GAIT MECHANICS OF INIDIVIDUALS WITH CHRONIC INCOMPLETE SPINAL CORD INJURY? By KRISTEN LYN JAGGER 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 2002

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This work is dedicated to my father, Robert Thornton Jagger, and my grandfather, Ernest Thornton Jagger, for passing on the desire to succeed.

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ACKNOWLEDGEMENTS I would like to offer special thanks to each of my academic committee members for their assistance during this dissertation process. Dr.'s Chow, Tillman, Cauraugh, Vander Griend, Falsetti, and Siders provided leadership and guidance that helped me to grow as a student, a teacher, and a researcher. I also extend my heartfelt appreciation to Mike Jones at Shepherd Center and Ben Johnson at Georgia State University in Atlanta, Georgia. Without their support throughout the recruiting and data collection processes I would not have been able to pursue this line of research. Most of all, I would like to acknowledge my family and friends for their role in my doctoral pursuits. I thank Joy for giving unconditional daily support and being a part of my "think tank"; Chris, for helping make sense of the senseless; Mum, for always showing me the silver lining; and Dad, for asking me the questions that no one else would. iii

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TABLE OF CONTENTS page ACKNOWLEDGEMENTS i" ABSTRACT vii CHAPTER 1 INTRODUCTION 1 2 REVIEW OF LITERATURE 5 Human Spinal Cord Anatomy 5 Spinal Cord Injury 7 Classification of Spinal Cord Injury 8 Complications of Spinal Cord Injury 9 Normal Human Gait 11 Phases of Gait 12 Weight acceptance 12 Single limb support 13 Limb advancement 14 Normal Joint Mechanics During Gait 14 The Foot 15 Role of the Foot During Gait 17 The Ankle 20 Role of Ankle During Gait 20 The Knee 21 Role of the Knee During Gait 21 The Hip 23 Role of the Hip During Gait 23 Electromyography During Normal Gait 24 Ankle musculature 24 Knee musculature 26 Hip musculature 27 Foot Orthotics 28 Relevant Orthotic Research 30 iv

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3 MATERIALS AND METHODS 33 Subjects 33 Methodology 34 Data Reduction 39 Clinical Gait Measures 40 Frontal Plane Joint Mechanics Measures 42 EMG Measures 45 Statistical Analysis 45 4 RESULTS 50 Effects of Orthotic Intervention 50 Clinical Gait Variables 50 Gait variable comparisons within the less-involved lower extremity 51 Gait variable comparisons within the more-involved lower extremity 51 Frontal Plane Joint Mechanics 52 Joint mechanics comparisons within the less-involved lower extremity 53 Joint mechanics comparisons within the more-involved lower extremity 54 Electromyography 54 Electromyographic Changes In the Less-Involved Lower Extremity 55 Electromyographic Changes In the More-Involved Lower Extremity 56 5 DISCUSSION 66 Orthotic Intervention 66 Clinical Gait Variables 66 Clinical gait comparisons within the less-involved lower extremity 67 Clinical gait comparisons within the more-involved lower extremity 68 Frontal Plane Joint Mechanics 68 Joint mechanics comparisons within the less-involved lower extremity 69 Joint mechanics comparisons within the more-involved lower extremity 75 Symmetry 76 Gait Symmetry 76 Frontal Plane Joint Mechanics Symmetry 80 Electromyography 82 Summary and Conclusions 89 V

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APPENDIX A CONSENT FORM 91 B SUBJECT INTAKE FORM 98 REFERENCES 99 BIOGRAPHICAL SKETCH 105 vi

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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 DO FOOT ORTHOTICS HAVE AN IMMEDIATE EFFECT ON THE GAIT MECHANICS OF INIDIVIDUALS WITH CHRONIC INCOMPLETE SPINAL CORD INJURY? By Kristen Lyn Jagger December, 2002 Chair: John W. Chow Major Department; Exercise and Sport Sciences Research regarding the effectiveness of foot orthotics during functional activities is primarily limited to the able-bodied population. No research determining the impact of orthotic intervention on specific gait mechanics of ambulatory individuals after spinal cord injury (SCI) has been identified, even though this group often exhibits poor foot and ankle alignment. Purpose: This research was designed to ascertain if orthotic intervention had an immediate effect on gait measures and frontal plane kinematics and kinetics in ambulatory individuals with SCI. Methods: Nine community ambulators with an American Spinal Injury Association "D" classification were recruited. Prior to testing subjects were measured and fitted with temporary orthotics. Testing included both shoeonly and orthotic conditions, administered in random order. A six-camera optical motion analysis system captured three-dimensional motion at 120Hz, while force plate and electromyographic (EMG) data, collected at 1200Hz, provided information regarding ground reaction forces and muscle activity, respectively. An inverse dynamic approach was used to calculate frontal plane joint angles, moments, and powers at the ankle. vii

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knee, and hip. Mean and peak EMG data were recorded from five subjects for qualitative comparison. Statistics: A 2x2 doubly multivariate repeated measures multivariate analysis of variance was used to discern the effect of orthotics on each leg. Results: Orthotics significantly increased the stance duration, step duration, and ratio of stance on the less-involved limb. Orthotics reduced the gait velocity, swing duration, maximum hip abduction angle, maximum ankle power generation, and the maximum knee power absorption on the less-involved limb. Only the maximum ankle supination angle from the more-involved limb was decreased by orthotic intervention. Qualitative results of the mean and peak EMG amplitudes suggest increases In tibialis antehor and rectus femoris muscle activity with orthotic intervention. Conclusions: There is an immediate effect of orthotic intervention on selected gait and biomechanical variables within the lessinvolved limb. There is less immediate influence on the more-involved extremity. Results of this study justify further research in this area. viii

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CHAPTER 1 INTRODUCTION In the United States there are more than 1 1 ,000 new cases of spinal cord injury (SCI) reported each year, with most occurring in individuals between sixteen and thirty years of age.®° One of the major consequences of SCI is paralysis or paresis of muscles below the level of injury often resulting in the inability to walk. As a result many persons with SCI must, at least in the more acute phases of their recovery, use wheelchairs for household or community mobility. Cun-ently, between 1 83,000 and 230,000 American citizens with SCI function at least in part from their wheelchairs.®" Still, despite its catastrophic nature, SCI no longer guarantees that an individual will necessarily function from a wheelchair for his/her entire lifetime. Within the past few decades SCI has been classified into two major categories: complete and incomplete injuries. Incomplete injuries are defined as a partial or complete preservation of sensory and motor function below the level of the spinal cord lesion.^^ Persons with incomplete spinal cord injuries comprise the majority of injuries at 50.8%. In contrast, complete injuries (no motor or sensory sparing below the level of the lesion) account for 45.8%.^ Due to advances in emergency medical intervention and spinal cord trauma management, more individuals with SCI are surviving their initial injuries and the eariy stages of their care. In addition these individuals are more likely to have incomplete injuries. ^^ ®° In part because of the growing number of persons with incomplete SCI, therapists can no longer assume that persons with SCI need only to be taught techniques and prescribed equipment which allow them to adapt to their injuries. The rehabilitation philosophy is shifting from pure adaptation to a focus on motor recovery. 1

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2 Ongoing research at model care centers around the nation is being conducted to ascertain the degree of central nervous system plasticity following SCI. This research trend focuses on the recovery of walking through body-weight-supported treadmill training (BWSTT).®° For this training the patient is placed in a torso harness that provides variable body-weight support. Trained individuals assist and control the patient's lower extremity and tmnk movement at different treadmill speeds. The general purpose of this training is to stimulate central pattern generators (CPG) to create a rhythmic locomotor pattern that increases the individual's capacity for ambulation. Central pattern generators are neural networks located in the central nervous system that are responsible for creating a pattern of repetitive motor actions.^ It is thought that commands for initiation and termination of these rhythmic generators originate at supraspinal levels^® and that the generators themselves lie within the spinal cord of felines.^^ Evidence for the presence of CPGs in humans is more tenuous, but according to MacKay-Lyons'*°, "The 'best guess' at this point is a cautious affirmation." Central pattern generators in humans are thought to be responsible for the presence of electromyographic activity during gait that is not measurable during voluntary muscle contractions.^ ®^ Recent studies of the electromyographic output from individuals with spinal cord injury during treadmill locomotion have shown that loading and unloading of the limbs enhances the activity of the antigravity muscles.^" It has yet to be seen, however, if these same electromyographic characteristics are noted during over ground ambulation. Although, with advanced treatment techniques, there is new hope for individuals with SCI to regain the ability to walk, rehabilitation following SCI remains extremely costly. Cun'ent estimates for medical care in the first year post-injury range from $168,627 to $572,178 and from $11,817 to $102,491 for each subsequent year depending on the level and severity of injury.®" Careful management of the cost of rehabilitative services offered to persons with SCI has become imperative for facilities and insurers hoping to maintain their

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3 viability. As a result, early SCI research into recovery of ambulation has focused on individuals with incomplete injunes. Because their injuries are "incomplete," these individuals have a theoretical advantage over persons with complete injuries when attempting to regain function, including the ability to walk. However, their levels of impairment may vary dramatically. Incomplete injuries often result in significant sensory or motor deficits of the upper extremities, lower extremities, and/or trunk. Even if they become ambulators, persons with incomplete SCI must often contend with significant gait dysfunction resulting from their physical impairments (i.e., increased muscle tone, spasticity, sensory ataxia, balance deficits, reduced strength). Often these impairments are addressed with bracing and/or use of assistive devices. For research purposes gait in this population may be summarily described as unable to walk, slow walking with an antalgic gait, and walking over ground at normal speeds.^ Though there is a large amount of research directed at increasing the locomotor capacity of the spinal cord injured population, there is little available infomriation regarding the impact of ambulation on these individuals' skeletal systems. Specifically, there is a dearth of evidence indicating appropriate interventions used to address skeletal alignment in these persons during gait training. Considering individuals are living longer with chronic SCI, the impact of orthopaedic complications resulting from abnormal joint mechanics, especially during gait, must be elucidated. Abnormal lower extremity mechanics, particulariy foot alignment, have been widely researched in able-bodied and athletic populations.'"' ^^•^^^^^"'^^^^•^^^^ Studies have focused on the use of in-shoe orthotic devices designed to alter the mechanics of the lower extremity and reduce the occurrence of secondary orthopaedic complications (i.e., pain or injury). Despite positive clinical results for orthotic intervention, this information has not been consistently applied to neurologically involved populations. Therefore, the

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4 purpose of this research is to investigate the effect of in-shoe orthotics on the gait mechanics of individuals with chronic incomplete spinal cord injury. In order to limit the number of variables under consideration, clinical gait measures and frontal plane kinematic and kinetic parameters were chosen for study. Clinical gait measures were chosen to describe gross functional changes. Kinematic and kinetic variables were chosen to illustrate changes in frontal plane mechanical alignment due to the physical position and theoretical action of the orthotics. In order to detemriine if orthotics have a significant effect on this population of individuals with incomplete spinal cord injury, the following research hypotheses are stated: 1 . There will be a significant difference in selected clinical gait variables between orthotic and shoe-only conditions during walking at a self-selected speed in this sample of individuals with chronic incomplete spinal cord injuries. 2. There will be a significant difference in selected frontal plane kinematic variables during walking at a self-selected speed between orthotic and shoe-only conditions in this sample of individuals with chronic incomplete spinal cord injuries. 3. There will be a significant difference in selected frontal plane kinetic variables during walking at a self-selected speed between orthotic and shoe-only conditions in this sample of individuals with chronic incomplete spinal cord injuries.

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CHAPTER 2 REVIEW OF LITERATURE Human Spinal Cord Anatomy The human central nervous system (CNS) is divided into six main parts: (1) the cerebral hemispheres; (2) the diencephalon; (3) the midbrain; (4) the pons and cerebellum; (5) the medulla; and (6) the spinal cord. The spinal cord is an elongated cylinder fewer than two centimeters in diameter.^^ It is a continuation of the medulla oblongata (inferior brainstem) and lies protected within a canal formed by the alignment of successive vertebral foraminae."® In addition to having the protection afforded by the osseous vertebrae, the spinal cord is also enveloped by the spinal meninges (i.e., pia mater, arachnoid mater, and dura mater) and a network of blood vessels lying within the surrounding fatty and loose connective tissue. In the average adult the spinal cord ranges from 42-45 cm in length and extends from the foramen magnum in the occipital bone to the level of the second lumbar vertebra. At this level we find the tapered, caudal termination of the spinal cord known as the conus medullaris.''® The spinal cord may be divided vertically into five regions: cervical (C1-C8), thoracic (T1-T12), lumbar (L1-L5), sacral (S1-S5), and coccygeal (one level). In cross-section, the spinal cord is divided into right and left sides, multiple laminae (gray matter), and a network of myelinated nerve axons (white matter) and connective tissue. A transverse section of the spinal cord reveals a butterfly-shaped, gray area flanked by a white region. The gray matter is composed of the nerve cell bodies that extend cranial-caudal and medial-lateral. The "wings" of the butterfly are actually the dorsal and ventral horns of the 5

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gray matter. In the cervical and lumbar regions these horns have a greater crosssectional area. These enlargements of the gray matter regions correspond with the increased nerve supplies necessary to innervate the upper and lower extremities. The cervical enlargement extends from C4-T1 , while the lumbar enlargement extends from 12-83."® Because the spinal cord is only two-thirds the length of the vertebral canal, position of the enlargements is not necessarily representative of vertebral levels. In the cervical region, for example, cord levels and vertebral levels are closely associated, while in the lumbosacral region cord levels L2-S3 are found at vertebral levels T1 1-L1 . Termination of the spinal cord at L2 necessitates longer nerve roots for the lower lumbar, sacral, and coccygeal levels. These nerve roots exit the cord as a bundle at L2 and descend caudally before passing laterally through their respective intervertebral foraminae. Named after its appearance, this bundle of nerves is called the cauda equina (horse's tail)."® The various layers of the spinal gray matter are known as laminae, and each of these laminae is responsible for receiving and relaying information from other portions of the CNS or the periphery to higher CNS centers or to skeletal muscle, viscera, or other organs/tissues. Cells within the dorsal horns transmit information regarding sensory input from the spinal nerve afferents, while those in the ventral horns contain alpha, beta, and gamma motoneurons that send efferent information through the ventral roots to innervate skeletal muscle fibers. Afferent and efferent axons from the various laminae of the spinal cord emerge as root fibers and converge laterally to form mixed nerves prior exiting the vertebral canal. Because the cord is divided into right and left sides, a pair of mixed spinal nerves originates from each of the above spinal levels. These nerves escape the vertebral canal laterally through the intervertebral foraminae and travel to the periphery. These nerves are named according to their associated vertebral level. Spinal nerves C1-C7 depart through intervertebral foraminae superior to their

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7 matched cervical vertebrae. Because there are eight cervical levels within the spinal cord and only seven cervical vertebrae, spinal nerve eight exits inferior to C7. All other spinal nerves caudal to this point pass below their name-matched vertebrae.^^ The white matter of the spinal cord is composed of myelinated nerve axons that run longitudinally through the cord. These fibers are responsible for the transmission of Information between the spinal cord and the brain.^^ Information is shared between the gray matter and white matter of the spinal cord by Interneurons. This network of neurons provides the brain with a system of control and feedback loops for the entire body. Subsequently, injury to any part of the spinal cord can have a profound effect on both the motor and sensory function of the human body. Spinal Cord Injury Injuries to the spinal cord can be grouped into two categories: traumatic and nontraumatic. Traumatic injuries are typically, but not always, the result of high velocity impact forces. These include gunshot wounds, stab and other penetrating wounds, indirect forces generated by movement of the head and/or trunk (e.g., shearing forces from a motor vehicle accident), and direct forces at the level of the vertebrae (e.g., blunt trauma, crush injuries, falls). In addition, trauma to the spinal cord may be the result of radiation exposure or surgical procedures. Non-traumatic Injuries include circulatory compromise (i.e., spinal stroke); cord compression (e.g., spinal stenosis and spondylolisthesis); demyelenating diseases (e.g., multiple sclerosis and primary lateral sclerosis); diseases of the anterior horn cells (i.e., amyotrophic lateral sclerosis and polio); inflammatory processes (e.g., spinal meningitis and acquired immunodeficiency induced neurophathy); space-occupying lesions (e.g., tumors, cysts, and syringomyelia); or congenital malformations (e.g., myelomeningoceole).

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8 Classification of Spinal Cord Injury There are two generally accepted methods for classifying spinal cord injury, regardless of the mechanism of injury. The first classification scheme involves a description of the spinal level at which the injury occurred, followed by a description of the most caudal level of intact motor or sensory function. For instance, a spinal "fracturedislocation of 03 on C4" describes the location of the insult, while "C5 incomplete tetraplegia" details the last level of fully preserved motor and sensory function. In the second classification scheme, defined by the American Spinal Injury Association (ASIA), it is necessary to describe the level of injury as well as the level of functional impairment. For example, included in this examination are items concerning deep anal pressure, position sense, stereognosis, pinch/grasp, and jaw chuck. In both schemes the terms incomplete and complete are used to denote the degree of preservation of sensory or motor function below the level of injury. A complete injury indicates that both motor and sensory function below the level of the lesion have been fully compromised, including sacral reflexes in the peri-anal region. With a complete injury, however, there may be zones of preservation of sensory or motor function. These do not, however, include the sacral elements. Preservation of sacral reflexes and sensation, also known as sacral sparing, defines an incomplete injury. Incomplete spinal cord injuries, therefore, are those injuries in which there is sparing of at least partial sensory and/or motor function below the level of the spinal injury, including deep anal sensation.''^ The terminology employed by the ASIA impairment scale can be found below in Table 1-1. In order to obtain a more accurate portrayal of spinal cord injury and its influence on function, it is necessary to use both classification schemes described previously. Yet these measures alone are incapable of fully describing an individual's functional abilities and impairments after SCI. Secondary complications may limit the function of one

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g individual more than another individual who had a similar mechanism and level of injury, despite being classified at the same ASIA level. For example, an individual with C5 tetraplegia with zones of sensory and motor preservation in Ce-j will be able to extend the wrist to produce a tenodesis grasp and to extend the elbows to assist with transfers and activities of daily living. A person with a C5 injury and preservation of only spotty sensation in some zones caudal to the lesion will not have the same function. Both individuals are classified at the C5 level, but the classification of A, B, C, D, or E helps to tease out the functional potential of the individual. Table 2-1 . ASIA Impairment Classification Scheme ASIA A Complete loss of sensory and motor function below lesion ASIAB Partial sensory preservation; no motor sparing ASIAC Sensory and motor preservation; <1/2 muscles grade > 3/5 ASIAD Sensory and motor preservation; >1/2 muscles grade > 3/5 ASIAE No permanent sensory or motor deficits From the National Spinal Cord Injury Statistical Center. Complications of Spinal Cord Injury There are various complications that result from spinal cord injury and alter the functional potential of its survivors. These complications can be grossly separated into orthopaedic, neurologic, and cardiorespiratory. Each of these complications leads to longer lengths of stay in the hospital or more frequent returns for follow-up care, delayed or more extensive rehabilitation, and subsequent delays in return to home and work. It is the responsibility of the rehabilitation team, therefore, to limit the occurrence and morbidity produced by these complications. Orthopaedic problems include potential for contracture, joint ankylosis, osteoporosis or osteopenia, hypercalcemia, fractures, spinal deformity, heterotopic ossification, and degenerative joint abnormalities.'^^ Osteoporosis, a progressive loss of bone mass, occurs primarily in the bones below the level of the lesion and advances rapidly during the first year after injury before stabilizing. Most bone loss occurs in the trochlear bone in

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10 the vertebral bodies, femoral heads/necks, the pelvis, and the bones of the leg, foot, and ankle. The cause of this rapid progression is not known, but the appearance of increased levels of calcium in the urine signals the initial onset. It is acknowledged that immobility is a partial cause of bone loss, and mobilization and weight bearing are recommended to slow its progression. Indeed, Plum and Dunning" advocate ambulation as the only type of weight bearing that has any effect on the amount of calcium in the blood. Keeping this in mind, any intervention that increases the efficiency of ambulation may increase the amount of time that an individual spends walking each day, thereby reducing risk of additional bone loss. At the same time, responsible clinicians must consider the acuity of the person's injury, the person's age and sex, and the degree of bone loss prior to initiation of gait or weight bearing activity. Degenerative joint abnormalities have been noted in both ambulatory and nonambulatory spinal cord injured clients. Wylie and Chakera^® have noted that the higher the level of injury, the more likely the individual is to develop joint degeneration. It has been hypothesized that joint degeneration would be greater at the hip and sacroiliac joints of ambulatory individuals than in nonambulatory individuals.^^ Sensory deficits resulting from SCI may include altered awareness of nociceptive input, altered light touch or pinprick/point discrimination, reduced baroreceptor input, reduced temperature discrimination, and altered proprioception.^^ Loss of nociception and baroreception places an individual at risk for integumentary complications, such as pressure sores, secondary to a decreased awareness of physical pressure and pain. This has a direct impact on an individual's capacity to regulariy perform weight shifts. Insufficient weight shifting results in areas of ischemia where regional vasculature is compressed. This eventually leads to tissue breakdown and risk of infection or need for surgical intervention. Decreased temperature sensation is dangerous for obvious reasons. A hot car seat or a prolonged wait at a cold bus stop can lead to serious problems for an

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11 individual with SCI. Reduced light touch and point discrimination make it difficult for individuals with SCI to sense changes in their support surface. They subsequently rely more heavily on their vision for information regarding surface changes and environmental obstacles. Lack of proprioceptive input affects an individual's perception of where the body's joints are in space. Loss of deep position sense can exacerbate joint laxity if joints are positioned at their end range of motion for prolonged periods of time. Lack of position sense in the foot and ankle, in combination with reduced muscle activity, can lead to reliance on the ligamentous structures to support the joints during weight-bearing. This holds serious implications for ambulatory individuals, considering the tensile and defomnation properties of connective tissue. Lack of structural support of the joints of the foot and ankle is guaranteed to facilitate joint deformation and mechanical instability. Within the human body are several receptors designed to sense joint position (i.e., Ruffini corpuscles). These joint proprioceptors are known to profoundly impact physical function, because they are crucial for controlled voluntary movement of the body segments.^^ Joint proprioceptors have also been implicated as playing a key role in central pattern generation. The higher CNS levels, both cortical and cerebellar, rely on feedback from joint proprioceptors, skin, muscle, and connective tissue to help refine mechanical efficiency and to accomplish purposeful movement. Therefore, it is the feedback from these various systems that helps an individual to "normalize" pattems of movement. Normal Human Gait Normal gait has been defined in detail by many researchers."' ^^'^^'^'^^ The normative values generated by their research serve as a standard for the pattems of lower limb, trunk, and upper body function in humans. According to these authors, the segments of the skeletal system move in predictable arcs, the joints rotate through specified ranges of motion, and the neuromusculature has a defined phasic pattern of activation. Bianchi and

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12 colleagues^ have even used the rotation of the body segments to amve at a description of normal gait that demonstrates how kinematic parameters drive the nervous system. Phases of Gait Although the phases of gait are often termed differently, there is consistency in the description of events throughout the full cycle. For ease of further discussion, however, the terminology defined by Pen7®^ will be used. The gait cycle can be broken up into three tasks and eight distinct phases. The tasks are termed weight acceptance, single limb support, and limb advancement. These tasks define the functional demands of the gait cycle. Within each of the tasks are the more descriptive phases of gait. These are initial contact, loading response, mid stance, terminal stance, pre-swing, initial swing, mid swing, and terminal swing. Each of these phases is named according to specific events of gait and is further described as percentages of the gait cycle (Figure 1-1).®^ Weight acceptance Within the task of weight acceptance there are two phases. Initial contact defines the moment when the foot touches the ground and occurs during the first 0-2% of the gait cycle. The objective of this phase is to position the limb for loading. Loading response (010% gait cycle) begins when one foot contacts the ground and continues until the opposite foot is lifted for swing. This period has also been termed the first double support or stance phase-the first time during which both feet are in contact with the ground.®^' ®^ The objectives of this stage include shock absorption, weight-bearing stability, and preservation of progression.

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13 Initial Single Limb Terminal Swing Double Limb Double Limb Stance Double Limb Stance Stance Stance (0%-10%) (10% -50%) (50% -60%) (60% -100%) (0%-10%) Figure 2-1 . Perry's phases of gait. Single limb support Upon termination of loading response, the body enters the period of single limb support. This period consists of mid stance and terminal stance and makes up 40% of the gait cycle. Mid stance is the first half of single limb support and occurs between 10% and 30% of the gait cycle. It begins when the opposite foot is lifted off the floor and continues until the body's center of mass is aligned over the forefoot. The primary objectives of this phase are progression of body weight over the stance foot and lower limb and trunk stability. Advancement of single limb support results in the heel lifting off the ground and signals the beginning of terminal stance (30-50% of the gait cycle). This phase continues until the opposite foot strikes the ground. The only objective of this phase is progression of the body past the supporting foot.^^

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14 Limb advancement Limb advancement is the final task of the limb during the gait cycle. It is composed of four phases: pre-swing, initial swing, mid swing, and terminal swing. Pre-swing typically occurs at 50-60% of the gait cycle and is also known as the second double support or stance interval.®®'®'' It begins when the contralateral foot hits the ground and continues until the ipsilateral foot leaves the ground. Weight transfer from one limb to the other occurs during this phase, and the main objective is to prepare and position the ipsilateral limb for the rapid demands of swing. ®^ Initial swing makes up the first one-third of the swing period (60-73%). It starts with the swing foot leaving the ground and ends when the foot is opposite the stance limb. The objectives of this phase are to clear the foot from the floor and advance the limb from its trailing position. The second one-third (73-87%) of the swing period is mid swing. Limb advancement and foot clearance are objectives of this phase, also. Finally, terminal swing completes the gait cycle. It begins when the tibia is vertical and ends when the foot contacts the floor. The objectives of this phase are to complete the advancement of the limb and prepare the limb for stance.®^ Normal Joint IVIeclianics During Gait Locomotion occurring under normal conditions incurs predictable forces on the structures of the lower extremity. Gait mechanics that do not resemble those defined as normal represent a greater challenge to the skeletal, articular, and neuromuscular systems that bear the load and resist the applied forces. The following are descriptions of the anatomy and normal mechanics of the foot, ankle, knee, and hip joints during gait.

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15 The Foot In clinical and scientific literature, function of the foot is often divided theoretically into that pertaining to the forefoot, midfoot, and rearfoot (Figure 1-2). While this nomenclature is convenient for qualitative descriptions, it should be noted that the foot's composite function can not be divided so clearly. Motion at any one joint can not be measured independently during weight bearing due to the interdependence of all of the structures. Figure 2-2. Functional segments of the foot. The forefoot is a composition of six functional joints: the first through fifth rays and the first metatarsophalangeal joint. The first ray is composed of the first metatarsal and the first cuneiform bone.^ It functions in three planes, producing dorsiflexion/inversion and plantarflexion/eversion about its axis." Rays two through four are fonned by the intersection of each metatarsal and its opposing cuneiform and appear to produce only sagittal plane motion. The fifth ray is composed of only the fifth metatarsal and is capable of triplanar motion into pronation and supination.^ The first metatarsophalangeal joint is comprised of the first metatarsal head and the base of the proximal phalanx of the hallux

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16 and is capable of abduction/adduction in the coronal plane and dorsiflexion/plantarflexion in the sagittal plane.^ The midtarsal or transverse tarsal joint is located between the forefoot and rearfoot structures and transmits forces and motion proximally and distally. This joint is made up of the talonavicular joint medially and the calcaneocuboid joint laterally. The axes at these joints allow a complex triplanar motion that consists of inversion/adduction and eversion/ abduction about the longitudinal axis (movement of the cuboid on the calcaneus) and dorsiflexion/abduction and plantarflexion/adduction of the forefoot about the oblique axis."*^ Lastly, the rearfoot, comprised of the talus and calcaneus, houses the articulation called the subtalar joint (STJ). Formally called the talocalcaneal joint, this composite of three articulations (anterior, middle, and posterior) is triplanar and capable of pronation and supination.^® Pronation can be described as a combination of foot eversion, abduction, and slight dorsiflexion. Supination refers to inversion, adduction, and slight plantarflexion. Figure 2-3 shows a foot with normal forefoot and rearfoot alignment, while Figures 2-4 and 2-5 show a forefoot varus deformity and a rearfoot varus deformity, respectively. The normal range of supination and pronation during walking gait is subject to controversy. Brown and Yavorsky^ state that between 6° and 10° of combined supination and pronation are necessary for normal walking. Subotnick''® reported a necessary 4-6° of pronation and 8-12° of supination for normal gait, and Wright and colleagues®^ note that an average of 10° of pronation is necessary in the first 8% of the stance phase. Other authors have presented calcaneal eversion requirements for normal gait.® ^ Close and colleagues® report that a peak of 4-6° of eversion is reached by 14% gait cycle, while Moseley and associates^" report a maximum eversion of 7.3° at approximately 57% gait cycle. There remains much debate regarding the best method of describing subtalar joint motion.

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17 Recent reports have demonstrated that measurement of calcaneal eversion/inversion is not representative of the complex three-dimensional motion of the subtalar joint. ^^^^ Role of the Foot During Gait The foot has four primary roles during the stance phase of gait:"^ accommodation to uneven ten-ain,"^ attenuation of the impact load of the ground reaction,"^ ^° ®® fonnation of a rigid lever for effective toe-off,'*^ and transferal of axial rotation of the leg to pronation/supination of the foot during stance.^®' ^ Each of these functions is accomplished by a combination of actions at the midtarsal and subtalar joints. The specific concept of transferal of axial rotation is often called "coupling behavior" and is also used to describe the effect of the rearfoot on the proximal structures of the lower extremity^^. Rodgers^^ stated, "The body requires a flexible foot to accommodate the variations in the external environment, a semirigid foot that can act as a spring and lever arm for the push-off during gait, and a rigid foot to enable BW (body weight) to be carried with adequate stability." At initial contact the heel touches the ground lateral to the ankle joint center. The ground reaction force creates a pronatory moment at the STJ that may stress the Figure 2-3. Normal foot alignment. Adapted from McPoil and Brocato.^'

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18 structures of the medial arch.^^ During pronation the talus rotates medially about the STJ axis and forces the calcaneus into a more valgus position. Talar rotation increases the freedom of motion at the transverse tarsal joint and creates a more flexible foot distal to the talus and navicular.®* A flexible foot, as mentioned above, is more able to comply with the ground beneath. During mid stance the foot is fixed on the ground and the lower leg rotates laterally placing the talocrural joint in a tightly approximated position. Once the talus is closepacked in the mortise, the foot begins to supinate, increasing the stability of the transverse tarsal joint and the longitudinal arch of the foot.®* This stability is further increased by the firm fit of the talus on the navicular.^®' "^ ''^ Figure 2-4. Uncompensated forefoot varus deformity. Adapted from McPoil 46 Figure 2-5. Uncompensated rearfoot varus deformity. Adapted McPoil and Brocato.''® and Brocato. As the limb enters terminal stance, the ankle reverts from dorsiflexion to plantarflexion at heel rise and forces the metatarsophalangeal joints to dorsiflex. Dorsiflexion of the phalanges lengthens the plantar aponeurosis due to its distal insertion on the metatarsal heads, thereby increasing tension and providing greater stability throughout the longitudinal arch. This is known as the "windlass effect" (see Figure 1-3). Thus, the

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19 combination of greater weight bearing, supination, and the windlass effect contribute to the formation of the rigid lever needed for maximum stability at toe-off." Should any, or all, of these mechanisms fail, the foot would be unable to form the same effective lever for pushoff. Unpublished observations^ indicate that while Individuals after SCI tend to present with a cavus foot in non-weight-bearing situations, they have a rigid rearfoot and an excessively mobile forefoot during weight bearing. This combination results in a foot that lacks rearfoot compliance at initial contact and the inability to form a rigid lever during terminal stance, or push-off. Figure 2-7 shows the result of an uncompensated forefoot and rearfoot varus. Figure 2-6. The windlass effect. Adapted from Norkin and Levangie.^^ Decreased rearfoot compliance results in a proximal transfer of ground reaction forces.^®' " Excessive motion in the forefoot decreases the ambulatory stride length, because there is little or no formation of a rigid lever. Placement of orthotics, therefore, is intended to reposition, or support, the bony structures of the foot in a more normal mechanical orientation. For example, an individual that presents with a rigid rearfoot inversion and a hypermobile forefoot varus would benefit from rearfoot "wedging", or posting, that brings the floor up to the medial rearfoot, and forefoot posting that prevents the forefoot from collapsing mediallly during weight bearing (see Figure 2-8). These additions would provide stability as well as increase the foot's capacity to create a rigid lever for push-off.

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20 The Ankle The ankle joint, or talocrural joint, is defined as the junction between the tibia, the fibula, and the talus. The junction of these three bones forms the tibiotalar, fibulotalar, and tibiofibular joints. It is also the point at which body weight is transferred from the leg to the foot. Motion at the talocrural joint has 1° of freedom, allowing flexion and extension in the sagittal plane 58 Figure 2-7. Uncompensated varus In weight-bearing. Adapted from McPoil and Brocato. 46 Figure 2-8. Use of posting to compensate for varus deformity. Adapted from McPoil and Brocato.'^ Role of the Ankle During Gait When the foot comes into contact with the surface, a ground reaction force (GRF) results. The geometric centroid of instantaneous applied force distribution is termed the center of pressure (COP).®^ During initial contact the GRF vector is directed posterior to the ankle joint center. The result is a moment, or torque, about the ankle joint called a plantar flexion moment, because it is trying to force the ankle into plantar flexion. The ankle, in response to this moment, plantar flexes until the foot is flat on the floor at an average angle of seven degrees.®^ Once the foot is flat, the limb progresses forward

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21 throughout the remainder of stance. In late mid stance the GRF moves anterior to the ankle, creating a dorsiflexion moment. To stabilize the ankle joint and continue progression of the limb, the plantar flexor muscles Increase their activity, first allowing an eccentric increase in dorsiflexion range of motion to a maximum of 10° and then concentrically plantar flexing to a maximum of 30° by the end of the stance phase.®^ Once the foot leaves the ground, the ankle rapidly dorsiflexes to neutral (0°) in order to clear the foot as the limb swings through. There is often an increase in plantar flexion during terminal swing (3-5°) that prepares the foot for initial contact.^^ In reference to the example above, a hypermobile forefoot that collapses during mid and terminal stance will theoretically cause a greater dorsiflexion range of motion secondary to the delayed push-off that results from the lack of a rigid foot formation. Placement of orthotics, therefore, should decrease the dorsiflexion range of motion by promoting a rigid forefoot for earlier push-off. The Knee The knee joint is the junction of the femur proximally and the tibia distally. Due to the general lack of congruency of the tibiofemoral articulation, motion in all three planes is possible. The primary arcs of motion are sagittal plane flexion and extension, while the lesser rotations are transverse plane internal and external rotation and frontal plane abduction and adduction. The knee provides stability during the stance phase of gait and mobility during the swing phase^. Role of the Knee During Gait Sagittal plane knee motion during the gait cycle ranges from 0-70° of flexion. At initial contact the knee is close to full extension (2° hyperextension to 5° flexion)^^ and is experiencing an extensor moment secondary to the anterior position of the GRF vector. During loading response, there is a rapid rate of flexion as the body progresses anteriorly

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22 and the limb is loaded. Approximately 18° of flexion is necessary to provide shock absorption and is achieved at rates as high as 300°/sec. This is the time at which maximal weight bearing occurs at the knee joint. After this loading, the knee extends once more to a maximum of 3° flexion by terminal stance, after which it rapidly flexes to a maximum of 60°-70° during the swing phase.^^^^ The peak flexion angle is achieved by mid swing, at which point there is a brief pause followed by a reverse in direction toward full extension by late terminal swing and initial contact. Transverse plane rotation has been described using electrogoniometers and 3-D videography with bone pins and surface markers.^' ^ At initial contact the femur is slightly externally rotated on a fully rotated tibia.^^ During loading response the tibia and, at a slightly slower rate, the femur internally rotate approximately 7° to their maximum angle by the beginning of mid stance.^ Electrogoniometric data presented by Kettelkamp and colleagues^^ indicate that tibial internal rotation persists through mid stance until the knee fully extends in terminal stance. By the time the knee is fully extended, the tibia and femur are, once again, fully externally rotated. This rotation remains until body weight is shifted to the contralateral limb and the ipsilateral limb prepares for swing. At this point the tibia and femur again internally rotate. In opposition to each of these reports, Lafortune et al^, using bone pins to quantify tibiofemoral rotation, report that internal rotation measuring less than 5° occurs at initial contact, while 0° rotation is the mean during the remainder of stance. A maximum external rotation of 9.4° was recorded at toe-off until 75 ms prior to initial contact.^ Coronal plane motion of abduction and adduction has been described by Kettelkamp et aP^ and Lafortune and colleagues.^ There are, again, conflicting results for these values. Both groups of researchers agree that the knee abducts during the stance phase (mean peak value of 6.4° according to Lafortune et al.^ During the swing phase, however.

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23 Kettelkamp and colleagues^^ reported a maximum value of 8° adduction, while Lafortune et al^ reported abduction and not adduction in four out of five subjects. The Hip The hip joint is composed of the head of the femur distally and the acetabulum of the pelvis proximally. These two structures form a ball-and-socket joint with 3° of freedom. Flexion and extension occur in the sagittal plane, abduction and adduction in the coronal plane, and internal and external rotation in the transverse plane. The primary role of the hip joint is to support the head, arms, and trunk in both static and dynamic situations.^ Role of the Hip During Gait The hip joint serves three major functions during gait.^^ First, it provides the junction between the lower extremities (the locomotor units) and the trunk (the passenger unit). Secondly, it provides stability during stance; and thirdly, it controls the limbs during advancement in swing phase. The hip joint is designed to provide more three-dimensional motion than the knee or ankle and its muscles are divided into stabilizers and prime movers. Sagittal plane motion consists of flexion and extension of the thigh segment relative to the vertical frontal plane or the pelvic segment. The rotational angle can be defined as the thigh angle (between the thigh segment and the vertical) or as the hip angle (between the thigh and the pelvis segment). Hip angle measures have been used in clinical literature, because they account for the degree of pelvic tilt present during gait. The normal range of motion of the hip is approximately 40° during gait.^°' At initial contact the hip is flexed 30° and remains relatively steady during loading response. At 38% gait cycle a neutral hip angle (0°) is achieved, and as mid stance is approached the hip continues its progression toward extension. A peak hip extension of 10° is reached when the contralateral foot contacts the ground at 50% gait cycle, signaling the reversal of hip rotation into flexion by

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24 initial swing. At 60% gait cycle the hip passes through neutral once more and continues to its peak near 35° during terminal swing.®^ Small rotations of coronal plane adduction and abduction occur at the hip during gait. At initial contact the hip is adducted approximately 10° and increases another five degrees during loading response.^" During mid stance there is a reversal in rotation and the hip reaches a neutral angle (0°) by the end of terminal stance. This rotation toward abduction continues into initial swing, reaching 5° of abduction before altering direction toward adduction once more. Measurements of transverse plane rotation at the hip have been highly variable when comparing results across various studies.^ The total arc of motion averages 8°, with peak internal rotation occurring during loading response and peak external rotation at the end of pre-swing. Electromyography During Normal Gait Just as there is a definable pattern of motion during human locomotion, there is also a structured and repetitive sequence of muscle activation. The timing of muscle activation is normally described in terms of the phase of gait during which it is active. Magnitude of muscle activation, on the other hand, is defined as a percentage of the maximal EMG produced by a maximal voluntary isometric contraction (MVIC) during a manual muscle test (MMT) muscle activity is presented according to the joint upon which it has the greatest effect. Only those muscles tested for this study will be described in detail here. Ankle musculature The muscles crossing the ankle joint are grossly divided into those that cross the anterior aspect of the ankle joint, termed dorsiflexors, and those that cross the posterior aspect of the joint, termed plantarflexors. The dorsiflexor muscles predominate during the swing period of gait secondary to their role in lifting the foot for ground clearance.

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25 The plantarflexor muscles are more active during stance as they are responsible for controlling progression of the body.®^ The primary movers during dorsiflexion are the tibialis anterior (TA), the extensor digitorum longus (EDL), and the extensor hallucis longus (EHL). The TA has the largest cross-sectional area and mass, has a similar lever arm length when compared to the EDL and EHL, and therefore produces the greatest amount of torque. The onset of dorsiflexor activity begins during pre-swing with the EHL firing first and is quickly followed by the TA and EDL. Tibialis anterior activity peaks initially during swing at approximately 35% MVIC, drops to 10% during mid swing, and rises sharply to 45% by eariy loading response.^^ The double peaks of the TA and EHL designate the pretibial muscles as biphasic. The plantarflexor muscle group is composed of seven primary movers that are subdivided into the triceps surae and the perimalleolar muscles. The triceps surae is composed of the soleus and gastrocnemius. The perimalleolar muscles are made up of the tibialis posterior (TP), flexor digitorum longus (FDL), flexor hallucis longus (FHL), peroneus brevis (PB), and peroneus longus (PL). The primary role of the triceps surae is production of plantar flexor torque, accounting for 93% of the theoretical total. The perimalleolar muscles are concerned with control of the subtalar joint and other foot articulations during gait and account for only 7% of the theoretical torque.®^ The soleus and gastrocnemius muscle activity begins during the loading response phase of gait. The medial head of the gastrocnemius and the soleus activate in near unison, while the lateral head of the gastrocnemius is delayed until mid stance.^® The increase in gastrocnemius activity follows a more linear path until late mid stance when it rises rapidly, culminating at a peak of 60% MVIC values in terminal stance. It returns to zero by the initiation of pre-swing.

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26 Knee musculature The muscles that cross the knee are grossly divided into those that extend the knee joint and those that flex the knee joint. Knee extensors cross the joint anteriorly while knee flexors generally cross the joint posteriorly. Muscle activity is present to assist with the three primary functions of the knee: shock absorption, extensor stability, and limb advancement.^^ The knee extensors are comprised mainly of the quadriceps muscle. The vastus medialis, vastus lateralis, vastus intermedius, and rectus femoris form the quadriceps. The vasti muscles cross only the knee joint, while the rectus femoris crosses the knee and hip joints. The vasti muscles act in synchrony, extending the knee during terminal swing (90% gait cycle) to position the limb for initial contact and rapidly increasing activity that peaks at 25% MVIC during eariy loading response (5% gait cycle). At the onset of mid stance these muscles drastically reduce their level of activity and are quiet by 1 5% of the gait cycle. The rectus femoris has a very different activation pattern due to its dual role of extending the knee and flexing the hip. It is active between late pre-swing (56% gait cycle) and eariy initial swing (64% gait cycle) at approximately 20% MVIC. This muscle is seldom active with the other vasti muscles during loading response.®^ Knee flexor muscles can be divided into those that cross one or two joints, as was seen in the knee extensor musculature. The popliteus and the biceps femoris short head (BF-SH) comprise the knee flexors that cross only the knee joint, while there are three knee flexors that cross both the hip and knee joints.®^ These muscles, in addition to BFSH, make up the hamstrings. They are the biceps femoris long head (BF-LH), semimembranosus (SM), and semitendinosus(ST). Although each of these is a primary hip extensor, they have dual roles as knee flexors. They share an onset time during late mid swing (75% gait cycle) and continue through terminal swing. The SM and ST often continue their activity throughout mid stance. Biceps femoris long head and ST reach

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27 peak activities near 20% MVIC during late mid swing and terminal swing, while SM peaks near 30% MVIC during early terminal swing. ®^ There are other muscles that add to the knee flexion effort even though their primary responsibilities lie at the hip and ankle. The gastrocnemius, mentioned previously, is a primary ankle plantarflexor that assists knee flexion from 15% to 50% of the gait cycle and peaks at 75% MVIC during mid terminal stance.®^ Hip musculature The muscles of the hip joint can be divided into extensors, abductors, flexors, and adductors. The roles of these muscles include stabilization of the trunk on the lower limb and limb advancement. The hip extensors and abductors are active during stance to stabilize the limb, the flexors are active during swing to advance the limb, and the adductors are active during the intervals between stance and swing to assist transition from one role to another. Due to limitations in electromyographic techniques there is no reliable information present regarding the activity of the deep external rotators.^^ The hip extensor muscles include the two-joint hamstring muscles, the adductor magnus, and the lower fibers of the gluteus maximus. The semimembranosus, semitendinosus, and biceps femoris long head act to eccentrically slow the fonward progression of the limb during late mid swing (80% gait cycle) and throughout terminal swing. They continue their activity until mid loading response (8% gait cycle). The adductor magnus and lower fibers of the gluteus maximus have a slightly later onset, initiating in mid terminal swing and continuing through most of loading response.®^ The biceps femoris long head and semimembranosus peak during eariy terminal swing at 20% and 30% MVIC, respectively, and taper off to 10% by initial contact. The adductor magnus and lower fibers of the gluteus maximus do not peak until initial contact, at which time they have been recorded at 40% and 25%, respectively.®^

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28 The hip flexor muscles are composed of the adductor longus, rectus femoris, gracilis, sartorius, and iliacus. According to Perry^\ activity of the flexor muscles is most significant during the first stride of gait at a normal speed. After this initial acceleration the muscle activity subsides greatly (<5% MVIC). During faster walking speeds, however, greater muscle activity is seen. Flexor muscle activity begins during pre-swing and continues through initial swing and into early mid swing. There appears to be a definable sequence of muscle activity dunng this time. The adductor longus is the first muscle to become active (during terminal stance) and also remains active the longest (into early mid swing). The adductor brevis may have a similar pattern of activity, but unreliable EMG techniques preclude its accurate collection. Rectus femoris, as stated earlier, becomes active during pre-swing and remains active into early initial swing. At self-selected speeds for walking, the rectus femoris showed little to no activation in half of the subjects tested.®^ Hip flexor activity is not typically witnessed during mid swing. This is largely due to the release of potential energy that is stored in the muscles when the limb moves into hip extension.®^ Foot Orthotics Foot orthotics are orthopaedic devices that are used to provide support under the metatarsal shafts and calcaneus in order to position the foot in a neutral position and prevent compensatory motion."® These inserts can be described in lay terms as "shims" that transpose the floor to the plantar surface of the foot while maintaining subtalar neutral (see Figure 2-4)."® Functional orthotics should not be confused with arch supports that create a weight-bearing surface within the longitudinal arch. Orthotics are intended to support the load-bearing structures of the foot while leaving the longitudinal arch to function in its normal mechanical role of suspension. A suspension bridge is designed with vertical supports at each end of an expanse and cables to support the central section. The foot, on the other hand, has a talus, a calcaneus, and metatarsals that are present

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29 proximally and distally for support, with longitudinal ligaments and fascia that act as the cables. Central support is not necessary and can be detrimental to the structural integrity of suspension-based structures. Orthotics are Type I medical devices that can be classified according to the density of materials and processes used during fabrication. Soft orthotics are constructed of lowdensity materials that deform easily and are intended for accommodative support"® and shock absorption.^® These can be bought off-the-shelf or can be custom-made but are not used in conjunction with posting (wedging) to correct biomechanical alignment.''® Semirigid orttiotics incorporate softness for shock absorption and rigidity for biomechanical control.'*® Unlike soft orthotics, they use extrinsic posting to control forefoot and/or rearfoot motion. Extrinsic posting refers to the addition of materials to a negative cast of the foot. This method of posting requires casting of the foot while it is maintained in neutral position, followed by formation of a "negative cast", or model, of the foot. Once the neutral negative cast is made, materials are added to the plantar surface until the desired amount of posting/control is achieved. Rigid ortliotics are the least forgiving of all inserts. They are typically made of a single layer of heat-moldable plastic in order to minimize material and provide the highest degree of biomechanical control. Intrinsic posting is used in these orthotics. This method requires careful addition of plaster to the neutral foot cast in order to post the foot without further addition of materials. Inaccuracies in this technique will result in an orthotic with an unacceptable fit for the client. Temporary orthotic posting falls into the semi-rigid category. The materials are of moderate density and are intended to biomechanically support the foot, without the preparation of a neutral cast. Soft inserts are cut to the size of the individual's insoles and posting is applied to the appropriate areas. Temporary orthotics are often manufactured to determine the effectiveness of this intervention and to modify the degree of posting necessary before a custom pair is fabricated for the client.

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30 Relevant Orthotic Research As stated above, the fundamental goal of orthotic intervention is to support the foot during the stance phase of gait such that the need for the foot to compensate for structural malalignment is reduced.^^ Many researchers have attempted to quantify rearfoot motion. A wide variety of analyses have led a large number of authors to report changes in kinematic variables as a direct result of orthotic intervention.^"' Some researchers specifically noted reductions in the maximum pronation of the foot, maximum pronation velocity, time to maximum pronation, total rearfoot movement,"^^^^^ rearfoot pronation,"^^^' and calcaneal eversion.^^^^^^'^^ Others noted alterations in the amount of internal tibial rotation as a result of orthotic use,^°' ^' ''^ while still others focused on the electromyographic changes that resulted from the use of orthotics.^^^ However, because other authors found no significant difference in many of the same variables," ^ ''^ results remain inconclusive®^ and further analyses of the methodologies must be performed. One of the many discrepancies noted among experiments is the type of kinematic analysis used. It is well accepted that proper human biomechanical analyses must be performed using three-dimensional techniques. Human segmental rotations and translations are not purely angular or linear, but a combination of the two. Therefore, the use of two-dimensional analyses to describe three-dimensional motions are suspect. Some authors have adhered to three-dimensional analyses,"^ ^^^^ ^® and others have In addition to questionable videographic methodology, there are concerns regarding the placement of rearfoot markers when attempting to measure calcaneal motion. Superficial markers are generally used in preference to intracortical bone markers, because they are noninvasive, cheaper, easy to apply, and do not typically hinder performance of the activity of interest. They are, however, more susceptible to

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31 measurement error. Calcaneal motion would best be measured noninvasively by marker placement directly over the heel, but this is not possible with activities that require shoes. Attempts have been made to represent calcaneal motion more accurately within the shoe. Nigg and associates^ cut windows in the heel counters of running shoes and placed markers directly over the calcaneus, while Geneva and colleagues^" affixed a piece of molded plastic to the heel and bent it in a "U" fashion over the heel counter of the shoe. Both methods are good attempts at measuring calcaneal motion more accurately, but window cutouts in shoes raise the question of lost stability for the rearfoot within the shoe, and molded plastic inserts may move on the heel and misrepresent calcaneal motion. Sandals have also been used in place of running shoes when quantifying rearfoot motion.^^ While providing adequate access to the heel for rearfoot markers, they lack the stability of a full lace-up running shoe. Altering shoe stability in any way may cause an individual to change the mechanics of their stride. The most accurate method of measuring true skeletal calcaneal motion to date is through the use of intracortical bone pins. Stacoff and colleagues^^ used bone pins to demonstrate the differences between skeletal motion and shoe movement during running. Their findings demonstrated an average 5.8°-7.3° difference between calcaneal and shoe eversion. A critique of the methodology must also be performed here. A small cutout had to be made in the heel counter of the shoe to allow attachment of the marker set to the bone. In addition, all testing was performed with a local anesthetic. Both of these procedures, though necessary, may have altered normal shoe motion or rearfoot biomechanics. Perhaps realizing the difficulties associated with accurate quantification of rearfoot motion, and in attempting to learn more about the impact of the ankle/foot complex on the structures higher in the kinematic chain, some authors have attempted to determine alterations in tibial rotations.^' Here again, the results remain inconclusive. Cornwall

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32 and McPoir° demonstrated a high correlation between rearfoot motion and tibial rotation (r = .953), and noted a decrease in internal tibial rotation velocity and acceleration with orthotic use compared with a barefoot condition. Nawoczenski and colleagues^^ noted a decrease in tibial rotation during the first 50% of stance when semi-rigid orthotics were used, but stated that no there was no significant difference in mean tibial rotation throughout stance between orthotic and non-orthotic conditions. Given the drawbacks and difficulties listed above, it does not seem prudent to attempt measurement of rearfoot motion In a population with abnormal gait characteristics when that of able-bodied "nomnals" is still under scrutiny. More appropriate global measures consist of general gait characteristics, joint angular measures, joint moments, joint powers, and average EMG amplitudes. The purpose of this research is to determine if orthotic intervention has an impact on mechanical components of gait affecting overall function. Because the functional variability between subjects would limit the usefulness of the data, a within-subject comparison of orthotic and shoe-only conditions will be employed.

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CHAPTER 3 MATERIALS AND METHODS Because of larger subject pools in major cities, the data collection for this project was conducted in the Georgia State University Biomechanics and Ergonomics Laboratory in Atlanta, Georgia. Subjects Eight males and one female were recruited for this research project (Table 3-1). Prospective subjects were located via a Shepherd Center (Atlanta, GA) database search, and a mailing describing this research was sent to those who resided within a 250-mile radius of Atlanta, Georgia. Individuals were included in the mailing if they were between the ages of 18 and 65 years and classified by the American Spinal Injury Association (ASIA) as Impairment Scale category D^° as of discharge from Shepherd Center, interested individuals were asked to return a reply card that included their name, address, and phone number. A phone interview was conducted with each respondent prior to scheduling. Inclusion criteria required that each subject was: (a) medically diagnosed with a first-time spinal cord injury from trauma, or vascular or orthopaedic pathology at cervical, thoracic, or lumbar levels; (b) at least six months post-SCI prior to participation in this research and had completed all spinal cord rehabilitation; (c) was categorized by ASIA as Impairment Scale category Df° (d) a community ambulator (50 m) with or without an assistive device; (e) a household ambulator (15 m) without an assistive device; and (f) able to give informed consent. Subjects were excluded if any of the above criteria were not met, or they were symptomatic for pain or other significant medical complications that would prohibit or interfere with testing of walking function. 33

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34 Table 3-1 . Subject data with level of injury and degrees of orthotic posting. (years) (cm) (kg) (level) (lower extremity) (Left, Right) (Left, Right) BS Male 62 185.4 97.5 Left 6° ,6° 6° , 6° CO Female 54 165.1 63.1 \ y-iu Right 6°, 4° 6°, 0° Gr Male 1 O •170 7 Dv5.0 Right go 2° 4°,0° DS Male 50 190.5 104.5 T10-11 Right 6°, 6° 6°, 6° EB Male 50 175.3 68.2 C4-5 Right 6°, 6° 4°, 6° JD Male 45 188.0 78.0 C5-6 Left 5°, 3° 6°, 6° MS Male 22 172.7 59.0 C3-5 Right 6°, 6° go 2° MT Male 63 185.4 90.9 C6-7 Right 3°, 5° 6°, 6° RM Male 34 177.8 78.9 T12 Left 6°, 4° 4°, 3° placed medially. Methodology Prior to the arrival of each subject, all experimental equipment was inspected and calibrated by the investigator. The AMTI (model BP400600) force plate (Advanced Medical Technology, Inc., Watertown, MA) was allowed at least 15 minutes to warm up per manufacturer guidelines. Six 120-Hz optical digital cameras were placed as shown in Figure 3-1 and a Peak Performance© calibration frame (Peak Performance Technologies, Inc.) was used to define a 2m x 2m x 2m (length x width x height) three-dimensional space in which all coordinate data would be gathered. Successful calibration was determined when the true length of a 0.914 m calibration rod was accurately estimated by the Peak Motus© 2000 video system and had a standard deviation from its true length of less than 0.003 m. When a subject arrived at the Georgia State University Biomechanics and Ergonomics Laboratory, he/she was provided with a consent form that was approved by the Georgia State University (GSU) Institutional Review Board (Appendix A). After reading and signing the consent form, the subject was assessed during a single four-hour period. The subject first changed into appropriate clothing (dark colored shorts and shirts) and was then weighed and measured. Measurements of anthropometric data (Figure 3-2) from the

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35 Camera 1 Camera 2 Camera 3 Camera 6 ^^^^^^ 5 Camera 4 Figure 3-1. Experimental setup. lower extremities and pelvis were recorded for use in determining segment inertia! properties. A manual muscle test and sensory evaluation were performed as defined by ASIA to confirm category of impairment. Moreand less-involved lower extremities were defined by summing the muscle grades for each leg. The leg with the lower score was labeled as the more-involved lower extremity. Further bilateral measures were taken to define the available rearfoot and forefoot motion. All measurements were recorded on a subject evaluation form (Appendix B). The method described by McPoil and Brocato"® was used to measure available forefoot and rearfoot motion. Briefly, this method required the subject to lie prone while two lines were drawn on the back of each lower extremity, one bisecting the calf and the other bisecting the calcaneus. These lines were used to determine the position of subtalar neutral and to measure the available calcaneal inversion and eversion and forefoot varus

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36 or valgus using a hand-held goniometer. The only variation from the procedure described by McPoil and Brocato"^ was the use of overpressure when measuring forefoot varus. Overpressure was used to obtain a more realistic measure of forefoot motion in a weightbearing position. Foot Length Figure 3-2. Anthropometric measurements. Adapted from Vaughn. All measures were used to determine the amount of forefoot and/or rearfoot posting necessary for optimal positioning for each foot. Temporary orthotics were then fabricated from semi-rigid posting material (EVA 70 shore A; Orthofeet, Inc., North Vale, NJ) and affixed to a semi-rigid felt backing that was cut to the size and shape of the subject's own shoe insoles. A hand-held electric rotary grinder (Black and Decker RTX™, The Black and Decker Corporation, Hampstead, MD) was used to smooth the edges of the posting material to reduce plantar surface irritation. Rearfoot posting material was placed under the calcaneus and forefoot posting was placed immediately posterior to the metatarsal heads in the forefoot. All posting materials were measured and placed according to the methods described by McPoil and Brocato,"® with a maximum of 6° allowed at the forefoot

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37 or rearfoot. The 6° limit was deemed necessary to protect individuals from injury to the fifth metatarsal caused by shifting the center of pressure laterally. Medial forefoot posting was used to support an excessive forefoot varus and medial rearfoot posting was used either to stabilize a greater than normal calcaneal inversion (rearfoot varus) or to limit an extreme calcaneal eversion (rearfoot valgus) during stance. The goal of the orthotics was to place the subtalar joint as close to neutral as possible during functional weight bearing tasks. Once the physical measures were recorded and orthotics were manufactured, electromyography (EMG) electrodes were placed at predetenmined sites (Figures 3-3, 3-4, and 3-5) to record the average and maximum amplitudes of selected muscles during walking. Electrode placement was determined using descriptions and diagrams provided by Cram.^^ Each electrode site was shaved and vigorously cleansed with alcohol prior to placement. Blue Sensor (Medicotest Inc., Rolling Meadows, Illinois) 2-cm diameter surface electrodes and a telemetered Noraxon Telemyo 900 (Noraxon USA, Inc., Scottsdale, AZ) EMG system were used to detect muscle activity, while the Peak Motus 2000© (Peak Performance Technologies, Inc.) displayed and recorded the raw signals. The order of testing was randomly assigned. Initial testing condition (orthotic or shoeonly) and leg (right or left) were logged on the subject's evaluation form. Reflective mari
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Figure 3-4. Electrode placement for rectus femoris. Adapted from Cram.^^

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39 practice passes were required for this leg also prior to collecting the test trials. After successful practice passes, five successful test trials were recorded. An unsuccessful pass or trial occurred when one foot was not fully and solely on the plate during a single pass, or the subject obviously altered his/her stride length (targeted) to step on the plate. Unsuccessful trials that occun^ed during the recording sessions were discarded. A 15minute rest and accommodation period was provided after the first condition. Additional rest periods were provided at any time the subject requested. Upon completion of all testing, electrodes and reflective mariners were removed and subjects were paid for mileage to and from the laboratory. Data Reduction Raw analog data from the force plate and telemetric EMG were saved directly onto the hard drive of the Peak Motus© 2000 workstation computer. Digital optical data from each camera were recorded and used to describe the three-dimensional path of motion of each reflective marker during each walking trial. Each path was defined manually as that pertaining to a specific reflective marker and then automatically tracked by the Peak Motus© 2000 system throughout the remainder of the trial. Occasional aberrant light would interrupt the automatic processing and trigger a query regarding the identity of one or more markers. The marker(s) of interest was (were) redefined manually at that time and automatic tracking was resumed. This process was followed for each subject's 20 walking trials. After pathways were identified in each trial, lower extremity coordinate, raw analog, and event data were extracted from the Peak Motus© 2000 system and loaded into a LabVIEW 6.0© (National Instruments Corporation, Austin, TX) kinematic and kinetic data calculation program. This program used the inverse dynamic approach described by

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40 Robertson and Winter^ to calculate three-dimensional kinennatic and kinetic data, in addition to providing clinical gait measures. ,R. ASIS L.ASIS r~ R. Femoral wand 6 CfJ ,^ R. Femoral j epicondyle^o R. Tibial wan• Heel head II Global reference frame , 1 5 Sacrum L.ASIS 14 1 p,A L. Femoral wand 13 1>^ L. Metatarsal head 11 L. Femoral epicondyie U2.y^ A L. Tibial wand 11 t~-\ J L. Malleolus \ ^"^19!-^. ARHeel L. Metatarsal 8 v^^^S , r. ,"' 7 ^ fcL2 head II \ --VqL. Meei Figure 3-6. Modified Helen Hayes marker system.^^ Clinical Gait IMeasures Clinical gait variables were analyzed because they are measures commonly used in rehabilitation to quantify the effect of specific interventions. Gait values are typically easy and cost-effective to measure and provide valuable information regarding functional capacity. The clinical gait variables of interest are defined below. For each subject, the

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41 average value over five trials for each measure and each condition was used in subsequent analysis. • Stride length (m) the distance either foot travels during one stride (foot contact to ipsilateral foot contact) (Figure 3-7) • Gait velocity (m/s) walking speed (measured as the stride length divided by the time taken to complete that stride) • Step length (m) the distance one foot moves ahead of the other foot during the gait cycle (Figure 3-7) • Step width (m) the distance between the heels of the feet during gait (Figure 3-7) • Toe clearance (m) the distance from the toe of the foot to the floor • Toe-out angle (°) the angle of the foot segment from the line of progression; also known as foot angle (Figure 3-7) • Step duration (s) the time for the test foot to take one step length • Stance duration (s) the time that the test foot is in contact with the ground • Swing duration (s) the time that the test foot is not in contact with the ground • Ratio of stance (%) the percentage of the gait cycle during which the test foot is in contact with the ground • Ratio of swing (%) the percentage of the gait cycle during which the test foot is not in contact with the ground • Double limb support (%) the percentage of the gait cycle during which both feet are in contact with the floor (Figure 3-8) • Single limb support (%) the percentage of the gait cycle during which the test foot only is in contact with the floor (Figure 3-8) I Left step length 1 Right step length I Stride length Figure 3-7. Selected gait variables.

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42 Initial Single Limb Terminal Swing Double Limb Double Limb Stance Double Limb Stance Stance Stance (0%-10%) (10% -50%) (50% -60%) (60% -100%) (0%-10%) Figure 3-8. Perry's phases of gait. Frontal Plane Joint Mechanics Measures Frontal plane joint mechanics were chosen for analysis because the orthotic posting material acts primarily by altering frontal plane position of the bones of the foot. Frontal plane kinematic and kinetic measures were computed using reference frames embedded at each segment's center of mass (Figure 3-9). Segment reference frames are used to define the position and orientation of each segment with respect to a global reference frame. ®^ Changes in position and/or orientation over time, in conjunction with the application of external forces, allow computation of joint angles, joint moments, and joint powers. The following briefly describes the placement of segment reference axes and the computation of frontal plane angles, moments, and powers. Each segment reference frame is composed of an x-, y-, and z-axis. The origin of each reference frame is at the center of mass of each segment, and the global reference frame lies outside the body within the laboratory setting (Figure 3-9). The x-axis of each

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43 segment reference frame is generally directed from the distal joint center to the proximal joint center. In each thigh, for example, the x-axis is directed toward the hip on a line between the knee and hip joint centers. Each leg's x-axis is directed toward the knee on a line between the knee and ankle joint centers, and each foot's x-axis Is directed from the foot's center of mass toward the heel marker on a line between the toe and heel markers. The pelvis must also have an embedded reference frame for hip computations to be performed. The pelvis's x-axis is simply directed from the pelvic center of mass vertically and is parallel to the global Z-axis. Once the x axis is defined, external markers (Figure 3-6) are used to define a plane that bisects the segment. In the hip, the xz-plane is defined by the x-axis and the thigh wand marker; and in the leg, the xz-plane is formed by the x-axis and the calf wand marker. The xy-plane is described in the foot by the ankle joint center, the heel marker, and the toe marker. Lastly, the xz-plane in the pelvis is formed by the x-axis and an axis parallel to the line between bilateral anterior superior iliac spines. Using the hght hand rule, the remaining axis, which is perpendicular to the previously defined plane, is determined. Joint angles in the frontal plane are defined as the rotation of the distal segment relative to the proximal segment about the anterior-posterior (AP) axis. Positive angles represent hip and knee adduction and foot pronation. Negative values represent hip and knee abduction and foot supination. For each subject, the average value over five trials for each measure and each condition was used in subsequent analysis. • Frontal plane hip angle (°) the rotation of the thigh reference frame relative to the pelvic reference frame about the AP axis (y axis) • Frontal plane knee angle (°) the rotation of the leg reference frame relative to the thigh reference frame about the AP axis (y axis) • Frontal plane ankle angle (°) the rotation of the foot reference frame relative to the leg reference frame about the AP axis (x axis)

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44 Moments are the cross product of forces applied at known distances from a point of rotation and are a measure of how much force is required to produce a rotation about the axis. Moments, lil
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45 resistance about the ankle joint center. This moment, along with frontal plane moments at the knee and hip, were used in the analysis. Joint power is the product of the resultant joint moment and the joint angular velocity. The sign of the product determines whether power is being generated (+) or absorbed (-). Maximum values of frontal plane power generation and absorption at the hip, knee, and ankle were analyzed here. For each subject, the average value over five trials for each measure and each condition was used in subsequent analysis. • Frontal plane hip power (W) the product of the frontal plane hip joint moment and the frontal plane angular velocity of the hip joint • Frontal plane knee power (W) the product of the frontal plane knee joint moment and the frontal plane angular velocity of the knee joint • Frontal plane ankle power (W) the product of the frontal plane ankle joint moment and the frontal plane angular velocity of the ankle joint EMG Measures Raw EMG data were extracted from the Peak Motus© 2000 system and processed using Microsoft© Excel (The Microsoft Corporation, Redmond, WA). Data were smoothed using a sliding averaging with a window of 1 3 data points. Once all channels were smoothed, the mean and maximum values during one stride were calculated for each trial and then averaged by side and condition. Because of technical problems encountered during data collection, complete EMG data were only available in five subjects. Statistical analyses, other than descriptive measures, were not performed on EMG values. Trends noted in mean and maximum EMG amplitudes are discussed. Statistical Analysis A single 2x2 (Side x Intervention) doubly multivariate repeated measures multivariate analysis of variance (DM MANOVA) was used to determine significant differences between side conditions and intervention conditions (Figure 3-6). In the presence of a main effect, univariate tests were used to identify the level of significance of each

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\ Wrx Mrv /' Figure 3-10. Free body diagram of the foot during pre-swing. The forces acting at the foot's center of mass create moments of resistance at the ankle joint center about their respective axes of rotation. FR=resultant force; XA=force in the x plane; YA=force in the y plane; ZA=force in the z plane; WF=force due to the weight of the foot; FAx=force at the ankle in the x plane; FAy=force at the ankle in the y plane; FAz=force at the ankle in the y plane; MRx=moment of resistance about the x-axis; MRy=moment of resistance about the y-axis; MRz=moment of resistance about the z-axis. dependent variable within each condition (Figures 3-7 and 3-8). Pairwise comparisons using Bonfferoni adjustments were used when appropriate as post hoc analyses to determine the effect of one condition level on another (Figure 3-9). Differences between orthotic and shoe-only conditions were used to determine the impact of the orthotics on clinical gait parameters and frontal plane gait mechanics in this sample of individuals. Electromyography data were not included in the statistical analysis. Data from only five subjects were available and were not deemed appropriate for a quantitative

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47 analysis. Mean and maximum values were used to detect trends in EMG outcomes as a result of orthotic intervention. All variables that achieved statistical significance at a=.10 are represented in the results and discussion. This a-level was chosen due to the experimental nature of this research. The functional level of individuals with chronic incomplete SCI required to participate in this research also resulted in recruitment of a small number of subjects. In order to determine valid conclusions about the effectiveness of orthotic intervention among a representative population of individuals with chronic incomplete SCI, a larger number of subjects would be needed. However, experimental research was necessary to justify the time and expense of a large sample and/or long-term study of this nature. The present research was intended to do just that. Orthotic Shoe MoreInvolved LessInvolved N = 9 N = 9 N = 9 N = 9 Figure 3-6. 2x2 RM MANOVA layout.

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Orthotic MoreInvolved LessInvolved Shoe N = 9 N = 9 I N = 9 N = 9 Figure 3-7. Univariate comparisons within the orthotic and shoe-only conditions. Orthotic Shoe MoreInvolved N = 9 N = 9 LessInvolved N = 9 N = 9 Figure 3-8. Univariate comparisons within the moreand less-involved lower extremities.

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Orthotic Shoe MoreInvolved N = 9 N = 9 -Il LessInvolved N = 9 N = 9 Figure 3-9 Pairwise comparisons between sides and conditions.

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CHAPTER 4 RESULTS The results of the MANOVA showed a significant main effect for the intervention condition (Wilk's Lambda=0.010; F=5665.89; error df=1) and no significant main effect for the side condition (Wilk's Lambda=0.642; F=1.052; error df=1). There was also no significant interaction between the side and intervention conditions (Wilk's Lambda=0.143; F=29.035; en-or df=1). The results of univariate analyses for the intervention condition are reviewed below and are followed by pairwise comparisons of intervention results to determine specific effects on bilateral limbs. Descriptive changes in EMG are reviewed last. Effects of Orthotic Intervention The univariate analyses for intervention used analyzed to determine which variables were significantly affected by the orthotic intervention, regardless of limb. These differences depicted the general effect of orthotics on bilateral lower extremities. Sphericity-assumed p-values were used to determine significance because there was no violation of sphericity according to Mauchly's test of sphericity. Clinical Gait Variables Univariate analyses revealed significant differences between the orthotic and shoeonly conditions on gait velocity, step duration, stance duration, ratio of stance, and ratio of swing (Table 4-1). Gait velocity was decreased by 8% during the orthotic condition (F=4.435; df=1 ,8). The step duration was lengthened by 1% when the orthotic was in use (F=6.530; df=1,8). The stance duration (F=5.614; df=1,8) and the ratio of stance (F=3.738; df=1 ,8) were also longer during the orthotic condition (2% and 1%, 50

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51 respectively), and the swing duration was 1% sliorter with the orthotics (F=3.738; df=1,8). Table 4-1. Univariate analysis results for the clinical gait variables. Orthotic Shoe-only Variable RAssn S.E. IVIean S.E. o -value Gait velocitv 1.01 0.082 1.1 0.100 0.068* Stride lenath 1.34 0.055 1.34 0.058 0.923 SteD lenath 1 (vn) 0.670 0.028 0.671 0.029 0.905 Step length 2 (m) 0.667 0.027 0.667 0.030 0.956 Step width (m) 0.093 0.009 0.094 0.010 0.801 Toe clearance (m) 0.028 0.003 0.028 0.003 0.780 Toe-out angle (°) 8.7 1.2 8.2 1.5 0.347 Step duration (s) 1.318 0.092 1.299 0.085 0.034** Stance duration (s) 0.891 0.079 0.872 0.072 0.045** Swing duration (s) 0.427 0.020 0.427 0.020 0.867 Ratio of stance (%) 67.1 1.2 66.7 1.2 0.089* Ratio of swing (%) 32.9 1.2 33.3 1.2 0.089* Double limb support (%) 15.1 1.3 14.7 1.2 0.213 Single limb support (%) 37.0 1.3 37.2 1.3 0.484 S.E. = standard error; * significant at a = .10 level;* * significant at a = 05 level. Gait variable comparisons within the less-involved lower extremity There was a significant difference between the orthotic and shoe-only conditions for step duration, stance duration, ratio of stance, and ratio of swing in the less-involved lower extremity according to pairwise comparison results (Table 4-2). The step duration was 2% longer with the orthotics in place. The stance duration and ratio of stance were both lengthened on the less-involved leg as a result of the orthotic intervention (3% and 1%, respectively). The ratio of swing was reduced by 1% during orthotic trials. Gait velocity and stride length were not included in pairwise comparisons as they are measures of a composite moreand less-involved lower extremity function. Gait variable comparisons within the more-involved lower extremity Orthotics did not have a measurable effect on the more-involved lower extremity gait variables in this sample of individuals. No clinical gait variables were significantly changed in the presence of orthotics (Table 4-3). Though not significantly different, there was a

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52 trend toward an increase in stance duration on the more-involved side during the orthotic condition. Table 4-2. Painwise comparisons between the orthotic and shoe-only conditions for the less-involved leg gait varibles. Orhtotic Shoe-onlv Variable Mean S.E. Mean S.E. p -value Step length 1 (m) 0.663 0.032 0.674 0.032 0.099* Step length 2 (m) 0.659 0.027 0.656 0.032 0.806 Step width (m) 0.091 0.011 0.101 0.010 0.286 Toe clearance (m) 0.025 0.003 0.024 0.003 0.389 Toe-out angle (°) 8.0 1.8 7.7 2.0 0.340 Step duration (s) 1.322 0.093 1.294 0.085 0.034** Stance duration (s) 0.914 0.081 0.887 0.074 0.030** Swing duration (s) 0.408 0.019 0.407 0.019 0.756 Ratio of stance (%) 68.6 1.3 68.0 1.3 0.058* Ratio of swing (%) 31.4 1.3 32.0 1.3 0.058* Double limb support (%) 15.3 1.3 14.9 1.2 0.287 Single limb support (%) 38.0 1.5 38.2 1.2 0.718 S.E. = standard en-or; * significant at a = .10 level; ** significant at a = .05 level. Table 4-3. Painwise comparisons between the orthotic and shoe-only conditions for the more-involved leg gait variables. Orthotic Shoe-onlv Variable Mean S.E. Mean S.E. p -value Step length 1 (m) 0.678 0.030 0.668 0.031 0.329 Step length 2 (m) 0.674 0.031 0.678 0.030 0.581 Step width (m) 0.094 0.009 0.087 0.012 0.292 Toe clearance (m) 0.030 0.004 0.032 0.004 0.214 Toe-out angle (°) 9.4 2.1 8.7 2.3 0.458 Step duration (s) 1,314 0.091 1.304 0.087 0.220 Stance duration (s) 0.869 0.077 0.858 0.071 0.189 Swing duration (s) 0.445 0.022 0.446 0.022 0.922 Ratio of stance (%) 65.6 1.3 65.4 1.2 0.329 Ratio of swing (%) 34.4 1.3 34.6 1.2 0.329 Double limb support (%) 14.8 1.2 14.6 1.2 0.206 Single limb support (%) 36.0 1.3 36.3 1.5 0.359 S.E. = standard error Frontal Plane Joint Mechanics When data for both legs were pooled, significant differences between the orthotic and shoe-only conditions were found in the maximum hip abduction angle, maximum frontal plane ankle power generation, and maximum frontal plane knee power absorption (Table 4-4). When compared to the shoe-only condition, the maximum hip abduction angle decreased by 16% during the orthotic trials (F=4.244; df=1,8), and the maximum

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53 frontal plane ankle power generation was reduced by 22% (F=4.365; df=1,8). The maximum knee power absorption was also decreased with the orthotic intervention (F=3.570; df=1,8). Table 4-4. Univariate results for the intervention kinematic and kinetic variables. Orthotic Shoe-only Variable Mean S.E. Mean S.E. p -value Max ankle pronation angle (°) 14.9 2.2 15.2 2.2 0.746 Max knee adduction angle (°) -4.1 1.2 -3.6 0.9 0.297 Max hip adduction angle (°) 9.2 0.8 9.0 1.0 0.593 Max ankle pronation moment (Nm) 35.45 2.80 42.91 7.24 0.276 Max knee adduction moment fNm^ 18.79 3.57 24.59 6.98 0.170 Max hip adduction moment (Nm) 50.98 6.81 47.53 8.67 0.317 Max ankle power generation (W) 52.84 14.34 61.93 18.30 0.070* Max knee power generation (W) 22.48 5.20 26.22 7.54 0.214 Max hip power generation (W) 40.00 3.04 40.56 6.03 0.920 Max ankle supination angle (°) -3.9 1.6 -4.4 1.8 0.394 Max knee abduction angle (°) -16.9 1.6 -17.1 1.4 0.820 Max hip abduction angle (°) -3.0 0.6 -3.5 0.7 0.073* Max ankle supination moment (Nm) -14.61 2.09 -20.23 5.40 0.182 Max knee abduction moment (Nm) -26.91 2.28 -34.08 6.30 0.299 Max hip abduction moment (Nm) -42.86 9.00 -44.14 10.14 0.728 Max ankle power absorption (W) -53.13 15.90 -77.33 36.33 0.298 Max knee power absorption (W) -19.44 7.50 -22.23 7.30 0.095* Max hip power absorption (W) -35.18 6.86 -32.80 4.17 0.642 S.E. = standard error; * significant at a = .10 level. Joint mechanics comparisons within the less-involved lower extremity There were significant differences in the pairwise comparisons between the orthotic and shoe-only conditions in the less-involved limb for the following variables: maximum hip abduction angle, maximum frontal plane ankle power generation, and maximum frontal plane knee power absorption (Table 4-5). The mean maximum hip abduction angle was decreased by 24% on the less-involved side as a result of orthotic intervention. The mean maximum frontal plane ankle power generation and knee power absorption were also reduced during orthotic trials (23% and 34%, respectively). While not achieving statistical significance, the maximum knee abduction moment was reduced with the orthotic intervention.

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54 Joint mechanics comparisons witliin the more-involved lower extremity Orthotic intervention pairwise comparisons only revealed one frontal plane variable that was significantly different in the more-involved limb. The maximum ankle supination angle was significantly different between the orthotic and shoe-only conditions for the more-involved extremity (Table 4-6). This supination angle was reduced by 70% with the use of orthotics. Table 4-5. Pairwise comparisons between the orthotic and shoe-only conditions for the less-involved leg kinematic and kinetic variables. Orhtotic Shoe-only IVIGan S.E. Mean S.E. p -Value Max ankle pronation angle (degrees) 12.9 3.8 12.7 3.6 0.851 Max knee adduction angle (degrees) -4.3 1.4 -3.5 1.1 0.253 Max hip adduction angle (degrees) 9.2 1.6 8.6 1.8 0.258 Max ankle pronation moment (Nm) 36.98 3.54 38.79 3.37 0.309 Max knee adduction moment (Nm) 17.06 3.19 18.50 3.96 0.387 Max hip adduction moment (Nm) 42.51 8.68 43.78 8.68 0.800 Max ankle power generation (W) 41.95 10.07 54.72 12.12 0.004** Max knee power generation (W) 14.44 2.17 17.54 1.97 0.197 Max hip power generation (W) 36.03 4.58 34.41 4.92 0.670 Max ankle supination angle (degrees) 7.1 3.7 6.7 3.4 0.705 Max knee abduction angle (degrees) 14.2 1.6 14.6 1.4 0.757 Max hip abduction angle (degrees) 2.6 1.4 3.4 1.5 0.032** Max ankle supination moment (Nm) 9.27 1.27 9.60 1.57 0.707 Max knee abduction moment (Nm) 27.95 3.20 31.33 3.17 0.117 Max hip abduction moment (Nm) 46.43 10.89 44.20 10.29 0.557 Max ankle power absorption (W) 32.42 7.27 34.96 5.21 0.614 Max knee power absorption (W) 9.32 1.45 14.11 2.93 0.043** Max hip power absorption (W) 29.94 5.90 30.65 2.85 0.900 S.E. = standard en-or; * significant at a = .10 level; ** significant at a = .05 level. Electromyography A qualitative analysis of mean and maximum EMG amplitudes during one stride was performed in five subjects. An examination of mean and maximum EMG values for the tibialis anterior (TA), medial gastrocnemius (MG), medial hamstrings (MH), and rectus femoris (RF) during shoe-only and orthotic conditions revealed changes in muscle activity in both the moreand less-involved lower extremities due to the orthotic intervention.

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55 Table 4-6. Pairwise comparisons between the orthotic and shoe-only conditions in the more-involved leg kinematic and kinetic variables. Orthotic Shoe-only Variable WIean S.E. Mean S.E. p -value Max ankle pronation angle (°) 16.9 3.4 17.7 3.6 0.562 Max knee adduction angle (°) -4.0 1.3 -3.7 1.0 0.514 Max hip adduction angle (°) 9.2 1.4 9.5 1.4 0.398 Max ankle pronation moment (Nm) 33.93 2.95 47.03 13.34 0.333 Max knee adduction moment (Nm) 20.51 4.37 30.67 11.19 0.210 Max hip adduction moment (Nm) 59.46 11.26 51.27 12.26 0.150 Max ankle power generation (W) 63.73 25.68 69.13 32.88 0.537 Max knee power generation (W) 30.51 10.23 34.90 15.59 0.461 Max hip power generation (W) 43.97 7.02 46.72 11.15 0.822 Max ankle supination angle (°) 0.6 3.5 2.1 3.8 0.049** Max knee abduction angle (°) 19.7 2.2 19.6 2.1 0.796 Max hip abduction angle (°) 3.4 1.1 3.7 1.3 0.545 Max ankle supination moment (Nm) 19.95 4.51 30.87 11.39 0.200 Max knee abduction moment (Nm) 25.88 1.69 36.83 12.66 0.408 Max hip abduction moment (Nm) 39.29 10.94 44.08 12.08 0.340 Max ankle power absorption (W) 73.83 30.02 119.69 70.41 0.317 Max knee power absorption (W) 29.56 14.34 30.34 13.80 0.715 Max hip power absorption (W) 40.41 10.63 34.94 7.45 0.634 S.E. = standard error; ** significant at a = .05 level. Electromyographic Changes In the Less-Involved Lower Extremity Mean TA amplitudes were generally higher during the orthotic condition, while mean MG activity showed little change in two subjects, a small increase in two other subjects, and a small decrease in one subject (Figures 4-1 and 4-2). Mean MH amplitudes were generally unaffected by orthotic intervention, while RF activity showed an increase in four subjects and no change in the fifth (Figures 4-3 and 4-4). Of these muscles, only the increases in mean TA and RF activity showed consistent changes in most subjects during the orthotic condition. Changes in TA maximum EMG activity showed a similar trend with orthotic intervention as that seen among mean amplitudes (Figures 4-1 and 4-5). Four of the five subjects demonstrated higher maximum values during the orthotic condition on the lessinvolved side. The remaining subject had a reduction in maximal TA activity (Figure 4-5). Maximum MG activity varied among the five individuals. Three subjects exhibited little or

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56 no change, while the other two subjects showed small decreases in activity on the lessinvolved limb with orthotic intervention (Figure 4-6). Little change was noted in maximal MH activity on the less-involved side in three subjects when orthotics were in use, one subject experienced a mild increase in maximum amplitude, and one subject demonstrated a substantial reduction in maximal activity (Figure 4-7). The use of orthotics had a mixed affect on maximum RF activity. One subject showed little change on the lessinvolved side during orthotic use, two subjects had small reductions, and two subjects showed increases in maximum RF activity (Figure 4-8). Maximum EMG values were more variable than mean amplitudes, with only the TA showing consistent changes (increases) in most subjects on the less-involved extremity due to the orthotic intervention. Electromyographic Changes In the More-Involved Lower Extremity When going from the shoe-only to the orthotic condition, increases in mean TA activity were observed in four out of five subjects (Figure 4-9). Mean MG activity increased in three subjects and showed little change in two subjects during the orthotic condition (Figure 4-10). Increases in mean MH activity were noted in two subjects during the orthotic condition, while two were largely unaffected, and one showed a substantial reduction (Figure 4-11). Mean RF activity in the more-involved limb increased minimally in two subjects, moderately in two subjects, and decreased moderately in one subject when the orthotics were in use (Figure 4-12). These findings are very similar to those on the lessinvolved side, but changes were generally of smaller magnitude. The most marked exception was the variability in mean MH activity between the orthotic and shoe-only conditions. Unlike the minimal change accompanying the orthotic intervention on the lessinvolved side, mean MH amplitudes varied considerably with orthotics on the moreinvolved side. Maximum EMG amplitudes on the more-involved lower extremity also changed during the orthotic intervention (Figure 4-1 3). The maximum TA amplitude increased during

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57 orthotic use in the same four subjects as was seen when observing the mean activity levels. The other subject demonstrated a decrease in maximum TA activity on the moreinvolved side. Maximum MG activity was increased in three subjects during the orthotic condition, while one subject showed a decrease and another remained unchanged (Figure 4-14). The more-involved extremity MG maximum EMG amplitude during the orthotic condition was minimally affected in three subjects, decreased in one subject, and increased in another (Figure 4-1 5). Lastly, the RF maximum EMG amplitude was reduced in four subjects and increased in one (Figure 4-16). la Orthotic Shoe-only | 350 1 BS RM CP JD Subject Figure 4-1 . Mean tibialis anterior EMG activity in the less-involved lower extremity during orthotic and shoe-only conditions.

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80 70 > 60 O 50 S 40 % 30 S 20 10 0 Orthotic Shoe-only i T r Tl BS RM CP Subject JD CO Figure 4-3. Mean medial EMG activity in the less-involved lower extremity during orthotic and shoe-only conditions (n=5).

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Orthotic Shoe-only Subject Figure 4-4. Mean rectus femoris EMG activity in the less-involved lower extremity during orthotic and shoe-only conditions. Orthotic Shoe-only 1800 1600 3. 1400 0 1200 m 1000 1 800 E 600 (0 400 ^ 200 0 BS RM CP Subject JD CO Figure 4-5. Maximum tibialis anterior EMG activity in the less-involved lower extremity during orthotic and shoe-only conditions.

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BS RM CP JD CO Subject Figure 4-6. Maximum medial gastrocnemius EMG activity in the lessInvolved lower extremity during orthotic and shoe-only conditions. Orthotic Shoe-only 1600 1400 2. 1200 Subject Figure 4-7. Maximum medial hamstring EMG activity in the less-involved lower extremity during orthotic and shoe-only conditions.

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Orthotic Shoe-only 1200 > 1000 O 800 600 E 400 X S 200 BS RM CP Subject JO CO Figure 4-8. Maximum rectus femoris EMG activity in ttie less-involved lower extremity during orthotic and shoe-only conditions. Orthotic Shoe-only 250 200 e> 150 s Ui c 100 ra o> = 50 Figure 4-9. Mean tibialis anterior EMG activity in the more-involved lower extremity during orthotic and shoe-only conditions.

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Orthotic Shoe-only 250 ^ 200 O 150 S 111 c 100 ra 0) = 50 0 BS RM CP Subject JD CO Figure 4-10. Mean medial gastrocnemius EMG activity in the more-involved lower extremity during orthotic and shoe-only conditions. Figure 4-1 1 . Mean medial hamstring EMG activity in the more-involved lower extremity during orthotic and shoe-only conditions.

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Figure 4-12. Mean rectus femoris EMG activity in tlie more-involved lower extremity during orthotic and slioe-only conditions. Orthotic Shoe-only 1200 > 1000 3 0 800 m g 600 .1 400 1 200 ll It BS RM CP Subject JD CO Figure 4-13. Maximum tibialis anterior EMG activity in the more-involved lower extremity during orthotic and shoe-only conditions.

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[Orthotic Shoe-only] 1200 > 1000 0 800 s lU g 600 1 400 I 200 BS RM CP JD Subject CO Figure 4-14. Maximum medial gastrocnemius ElVIG activity in the moreinvolved lower extremity during orthotic and shoe-only conditions. Figure 4-15. Maximum medial hamstring EMG activity in the more-involved lower extremity during orthotic and shoe-only conditions.

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Orthotic Shoe-only 1200 1 > 1000 0 800 Ml g 600 1 400 S 200 0 BS RM C Sub P ect JD CO Figure 4-16. Maximum rectus femoris EMG activity in ttie more-involved lower extremity during orthotic and shoe-only conditions.

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CHAPTER 5 DISCUSSION All of the nine subjects who participated In this research had an excessive forefoot varus bilaterally and six had a hypomobile rearfoot valgus bilaterally. Had these individuals presented without neurologic injury to an orthopaedic clinic, it is likely that they would have been measured for and fitted with orthotics to improve their lower extremity mechanical alignment. Because the medical paradigm for prescribing orthotics has not yet shifted to Include those with neurologic injury, these Individuals are typically overlooked. Results of this study indicate that orthotics have an immediate effect on specific clinical gait charactehstlcs and frontal plane joint mechanics of this small subject sample and suggest that ambulatory individuals with chronic incomplete SCI may benefit from foot evaluation and appropriate orthotic placement. This study represents the first known attempt to determine the effects of orthotic intervention among ambulatory individuals with chronic Incomplete SCI. Orthotic Intervention Statistical analyses revealed that orthotics had a measurable effect on specific clinical gait parameters and frontal plane joint mechanics in this group of individuals with chronic incomplete SCI. The following sections discuss the greater impact of orthotics on the lessinvolved leg and the lesser impact on the more-Involved lower extremity. Clinical Gait Variables Five clinical gait measures were significantly affected by the use of orthotics. Step duration, stance duration, and ratio of stance experienced Increases, while gait velocity and ratio of swing were reduced In the orthotic condition. The magnitude of change for 66

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67 each of these variables was minimal and not clinically significantly, but such immediate changes in these gait parameters may be an indicator of future adaptations. The presence of these immediately measurable changes justifies the need for future research that includes follow-up testing after several weeks of orthotic use. Follow-up testing and a larger sample size would help to discern the long-term benefits of orthotic use in this population. Univariate statistics showed that gait velocity was decreased during the orthotic condition. While no orthotic research was found that has studied the effect of orthotics on gait velocity, an increase was anticipated due to the function of the forefoot posting. The medial forefoot posting was placed to provide support for a hypermobile forefoot varus, that had limited ability to form an effective and/or timely lever for push-off during terminal stance. One of the functions of orthotics is to bring the ground up to the foot instead of allowing the foot to become overiy compliant with the ground. Improving the effectiveness of push-off should, theoretically, have led to increases in stride length and gait velocity. In these subjects, however, this was not the case. It is likely that the short period between orthotic placement and testing did not allow each individual enough time for sensory accommodation. The altered sensory stimulus may have caused these subjects to be hesitant during gait. Further testing would be beneficial to determine the effect of orthotics on gait velocity after a longer period of accommodation. Clinical gait comparisons within the less-involved lower extremity Differences in less-involved lower extremity clinical gait characteristics between the orthotic and shoe-only conditions were limited to measures of phase duration. Increases in step duration, stance duration, and ratio of stance all referred to the lengthened period of time that the less-involved foot was in contact with the ground. Ratio of swing, having a direct relationship with ratio of stance (ratio of stance + ratio of swing = 100% gait cycle), showed an expected decrease. The only research report found that discusses changes in

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68 the stance duration of gait as a result of orthotic placement was a gait study among children with Down Syndrome/^ The authors did not report significant changes in stance duration as a result of orthotic use but did find a decrease in the inter-trial variability of stance duration with orthotics. A review of stance duration standard errors in the cun-ent study revealed a very small increase in stance duration variability resulting from orthotic placement (Table 4-2). This finding is in direct opposition to that of Selby-Silverstein and colleagues.''^ Clinical gait comparisons within the more-involved lower extremity Orthotics did not have a measurable impact on the more-involved lower extremity gait variables in this sample of individuals. This result was not entirely unexpected given the neurologic capacity of the more-involved limb versus that of the less-involved limb. The more-involved lower extremity in all subjects demonstrated greater strength and sensory impairments than the less-involved leg. Altered sensation may not have allowed for adequate awareness of the placement of the orthotics and, subsequently, may not have triggered enough of a response to alter the gait characteristics of the limb. Impaired strength, on the other hand, may have reduced the individual's physical ability to make adjustments to their altered mechanical alignment. Frontal Plane Joint Mechanics Selected hip, knee, and ankle kinematic and kinetic characteristics were significantly affected by the orthotic intervention. Maximum hip abduction angles were decreased, and maximum knee power generation and maximum ankle power absorption were reduced when orthotics were used. Pairwise comparison results showed that the less-involved lower extremity was influenced to a greater degree than the more-involved lower extremity.

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Joint mechanics comparisons within the less-involved lower extremity A 24% decrease in the maximum hip abduction angle was noted on the less-involved limb during the orthotic condition. Graphic representation of the average frontal plane hip range of motion of one subject over five trials shows this reduction occurring during the stance phase of gait (Figure 5-1 ). During single limb support, while the foot is fixed in a closed kinetic chain relationship to the trunk segment, frontal plane hip range of motion varies as a function of lateral pelvic tilt and thigh position. A lateral lean toward the stance limb was typically employed by most of the subjects during the swing phase in order to clear the contralateral (swing) foot from the floor. When the trunk laterally flexes and the pelvis tilts laterally to a greater degree over the stance limb, a larger abduction range of motion results. Frontal plane hip angles during gait among able-bodied subjects have a very different pattern than the one found here. During initial stance the hip is adducted an average of 10° in normal gait.^° During the shoe-only condition this group of subjects did not adduct the less-involved limb at all during initial stance. They averaged 1° of abduction until orthotics were added, at which point, they achieved 1-2° of adduction in eariy stance. By mid stance, nonnal range of motion values reach 15° of adduction. In these subjects, however, the hip angle progressed into abduction and peaked in eariy mid stance at 7° during the shoe-only trials and 5° during orthotic trials. While the impact of the orthotics is not large in its magnitude of change, one must remember that the orthotics were in place for only 30-45 minutes prior to testing. Further research is justified to perform a longer-term study that tests this population immediately and after four to eight weeks of orthotic use. In this group, specifically, immediate decreases in hip abduction during orthotic intervention trials had a very positive effect. Due to the lateral lean employed by the majority of the subjects to clear the contralateral limb

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70 from the floor, lower maximum hip abduction angles indicate that there was a decrease in lateral flexion over the less-involved limb during the stance phase. A decrease in lateral flexion could partially be the result of more appropriate contralateral limb toe clearance during swing. Though it did not achieve statistical significance, five of the nine subjects actually demonstrated a decrease in toe clearance height on the more-involved side during the swing phase while using orthotics. This finding is consistent with the decrease in lateral trunk flexion over the less-involved limb during stance. Reducing lateral tnjnk flexion over the less-involved leg brings the more-involved foot closer to the ground, and toe clearance height is thus reduced. If the reduction in the amount of lateral flexion exceeds the corresponding change in toe clearance height, however, another mechanism must be adopted. •^—Orthotic —Shoe 10 I , -10 % Gait Cycle Figure 5-1. Less-involved limb frontal plane hip angles from one subject. A review of sagittal plane hip flexion angles, though not included in this analysis, revealed an increase in maximum flexion angles in six subjects during the swing phase of gait in the more-involved leg. This suggests that the subjects in this study altered their lower extremity compensatory motions in favor of a more normal pattern of sagittal plane hip flexion to clear the foot from the floor. Though the changes in joint angles were small

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71 in magnitude, they represent an immediate and distinct alteration in gait strategies as a result of orthotic intervention. A significant decrease (23%) in frontal plane ankle power generation in the lessinvolved limb was noted during early stance (Figure 5-2). Inspection of the power curves for the orthotic and shoe-only conditions reveals that there was a smaller magnitude of change from peak power generation (~10% gait cycle) to peak power absorption (~16% gait cycle) in the orthotic condition. Joint power is the product of joint angular velocity and the resultant moment of the joint. Angular data indicate that the time rate of change of ankle angle (angular velocity) from supination to pronation is relatively similar for the two conditions. This finding fails to support reports in the literature that pronation velocity decreases with orthotic intervention."^ ^® Orthotic Shoe 100 I 80 ^ 50 A Generation 0 "W " \ y^r>^^^2i^^ > <^ — . — . — -20 ' V / „Absorption -ou ' _j % Gait Cycle Figure 5-2. Less-involved lower extremity frontal plane ankle power from one subject. Data represent the average of five trials per intervention. A graph of the frontal plane ankle joint range of motion from one subject shows that the slopes of both orthotic and shoe-only curves are almost identical at approximately 10% of the gait cycle (Figure 5-3). Because the curves represent the change in joint angle over time, the slope of the line is descriptive of frontal plane ankle joint angular

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72 velocity. Lack of variation between the two conditions indicates that any change in power generation is due to differences in the joint moment. One subject's average frontal plane ankle joint moments were graphed for each intervention to illustrate the change that caused the reduction in power generation (Figure 5-4). The pronation moment was found to be lower for the orthotic condition, suggesting that the smaller moment was primarily responsible for the lower power generation with the orthotic intervention. Changes in frontal plane ankle moment and power as a result of orthotic placement have not been reported previously. Orthotic Shoe % Gait Cycle Figure 5-3. Less-involved lower extremity frontal plane ankle angle from one subject. Data represent the average of five trials per intervention. There was a significant decrease (34%) in maximum knee power absorption as a result of orthotic intervention (Figure 5-5). As with ankle power generation, the two components that determine knee joint power are joint angular velocity and joint moment. During terminal stance (~60% gait cycle), when the maximum knee absorption power was most notably different between conditions, the knee angle was progressing into abduction (Figure 5-6) with an adduction moment of force (Figure 5-7). An absorption power indicates that the knee was being eccentrically controlled late in the stance phase (i.e., slowing the progression of the knee into abduction).

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73 Orthotic —Shoe % Gait Cycle Figure 5-4. Less-involved lower extremity frontal plane ankle moment from one subject. Data represent the average of five trials per intervention. —Orthotic —Shoe % Gait Cycle Figure 5-5. Less-involved lower extremity frontal plane knee power from one subject. Data represent the average of five trials per intervention. During late stance the COP is acting through the forefoot. A GRF that acts lateral to the ankle joint center causes a pronation moment that forces the foot into pronation. Foot pronation has been shown to increase tibial internal rotation angles'°^^ and femoral internal rotation angles.®" A combination of internal tibial and femoral rotation will necessarily increase the knee abduction angle due to the congruency of the tibial and

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74 femoral condyles. From the graph of knee abduction angle (Figure 5-6), it appears that the abduction angular velocity (the slope of the lines) did not change between the orthotic and shoe-only conditions. Thus, the change in knee power absorption must have been the result of different frontal plane moments at the knee joint. The frontal plane knee moment curves (Figure 5-7) reveal that there was a reduction in the knee adduction moment late in the stance phase. This could have been caused by a decrease in the magnitude of the GRF and/or the moment arm length between the knee joint center and the point of application of the GRF. Regardless of the cause, the result was a decrease in the adduction moment that was required to control the knee's progression into abduction during late stance. A reduction in the frontal plane moment should, theoretically, result in less compressive force placed on the lateral condyles of the tibia and femur. A smaller compressive force, in turn, would reduce the stress (force per unit area) on the articular cartilage of the knee. Less stress on the joint's passive soft-tissue structures over time could result in decreased wear and tear of the joint and even reduce pain. Figure 5-6. Less-involved lower extremity frontal plane knee angle from one subject. Data represent the average of five trials per intervention.

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75 Orthotic Shoe %Gait Cycle Figure 5-7. Less-involved lower extremity frontal plane knee moment from one subject. Data represent the average of five trials per intervention. Joint mechanics comparisons within the more-involved lower extremity Orthotic intervention did not have as great an effect on the more-involved limb as it did on the less-involved limb. Only the maximum ankle supination angle was significantly different between the orthotic and shoe-only conditions for the more-involved lower extremity. This supination angle was reduced 70% with the use of the orthotic. As previously mentioned, the relatively limited Impact of the orthotic on the more-involved limb was not altogether surprising considering the type and degree of neuromuscular deficits noted within the extremity. Because the neuromuscular system responsible for the limb function was not as sensitive to or able to adequately respond to minor alterations in mechanical positioning, segments higher in the chain may have been unable to respond to the same degree. Peripheral sensory deficits result in less afferent information that is available to the spinal cord and consequently influences reflexive or voluntary output to the musculature. Weakness in the more-involved musculature also affects outcomes secondary to the inability to increase or decrease muscle activity to

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76 optimal levels. The fact that the orthotic was able to produce an immediate change in frontal plane ankle alignment, however, is quite promising. Symmetry When discussing the effect that orthotics have on specific gait variables, it is also necessary to discuss the bilateral symmetry of these variables with and without orthotic intervention. Gait is a product of symmetric segmental motions and any decrease in asymmetry that results from orthotic intervention should be seen as a positive outcome. There were several variables that did not achieve significance in the statistical analysis that did, nonetheless, demonstrate mean changes between the two intervention conditions. Gait Symmetry Variations In lower extremity gait symmetry between the two interventions can be used to deduce benefits and drawbacks of orthotics in this population. In the following graphs, the position of each line's intersection with the vertical axis can be used to determine the degree of symmetry between orthotic and shoe-only conditions. The closer the two points are together on an axis, the greater the symmetry. Step width and toe clearance height are two clinical gait parameters that demonstrated greater lower extremity symmetry during the orthotic condition. There was a large reduction in step width on the less-involved side and an increase of similar magnitude on the more-involved side when orthotics were In use (Figure 5-8). Step width is the distance between the two heel markers during double limb stance (Figure 3-7). This distance Is calculated during the double limb stance period that follows the test foot contact with the force plate. Though the magnitudes of change between intervention conditions for each lower extremity were small (more-Involved: 0.69cm difference, lessinvolved: 0.94cm difference), the step width difference between bilateral extremities was noticeably smaller during the orthotic condition (orthotic: 0.28cm difference, shoe-only:

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77 1.34cm difference). Alterations in the GRF's, line of progression, path of the COM, and other variables could have contributed to greater step width symmetry. E .096 .094 Q. CO .092 .090 .088 .086 More-Involved Less-Involved Orthotic Shoe Intervention Figure 5-8. Step width maximum values for both sides and interventions. Toe clearance height also showed greater symmetry during the orthotic condition (Figure 5-9). As was seen with step width, there was a reduction in toe clearance height on one side as a result of orthotic intervention and an increase on the other with relatively small magnitudes of change (less-involved: 0.18cm difference, more-involved: 0.25cm difference). Here again, however, there was a greater disparity between extremities within the shoe-only condition (orthotic: 0.45 difference, shoe-only: 0.89 difference). No literature was located that supports or refutes greater toe clearance symmetry resulting from orthotic intervention. While step width and toe clearance were both more symmetrical during the orthotic condition, step duration was actually more asymmetrical after the orthotic placement (Figure 5-10). One possible explanation for a loss of symmetry after orthotic placement

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78 lies within the less-involved limb's ability to better accommodate to changes when compared to the more-involved extremity. Superior muscular and sensory capabilities on the less-involved side may have enabled these subjects to adapt and respond to the placement of the orthotic. .034 More-Involved Less-Involved Orthotic Shoe Intervention Figure 5-9. Maximum toe clearance height for both sides and interventions. In their research of various gait parameters among individuals after CVA, Yavuzer and Ergin^ actually discovered immediate increases in the hemiparetic (more-involved) limb stance duration with an arm-sling intervention. The arm sling could not have altered the sensory or muscular ability of either leg, yet it caused an increase in stance time on the hemiparetic side. These investigators also discovered that the ami sling reduced the medial-lateral excursion of the center of mass (COM) and may have been the key to the longer stance duration on the hemiparetic side.

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79 .92 (0 C O CO .89 .90 .91 Q 88 Side More-Involved .85 Less-Involved Orthotic Shoe Intervention Figure 5-10. Stance duration for both side and intervention conditions. Use of a sling to place the ami across the torso decreases the arm's effective perpendicular distance from midline, thus shifting the COM slightly toward the contralateral side. This shift necessitates a smaller medial-lateral translation from the line of progression toward the less-involved side to maintain upright balance. It also requires a greater mediallateral translation toward the more-involved side for the COM to pass over the foot. Because this distance is greater than that without the ann-sling intervention, more time would be necessary to travel the distance to and from the line of progression (assuming no change in velocity) and would be measured as an increase in stance duration. Center of mass variables were not included in the present study and, therefore, cannot be addressed directly. It is possible, given Yavuzer and Ergin's^° findings, that analysis of medial-lateral COM translation toward the more-involved leg would not reveal any significant changes between orthotic and shoe-only conditions because orthotics were unable to immediately increase the stance time on that limb.

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80 Frontal Plane Joint Mechanics Symmetry Improved proprioception and pain perception help the musculoskeletal system to control joint motion through stabilization and acceleration/deceleration. For example, during mid stance when the knee reaches terminal extension, proprioceptors signal the spinal cord to trigger a response from the surrounding musculature to facilitate cocontraction and prevent hyperextension.^^ If an individual has a sensory impairment and lacks appropriate proprioception, the protective response to reduce knee extension is not elicited in time, and hyperextension occurs. Additionally, if an individual has a flexible flat foot, he/she would undergo excessive tibial internal rotation and knee abduction during weight bearing."^' Greater knee internal rotation and abduction angles place both the posterior capsule of the knee and medial collateral ligaments at risk for developing laxity. Knee abduction and internal rotation also place greater compressive and shear forces on the lateral femoral and tibial condyles. Alternatively, if a person has appropriate proprioceptive awareness but not the necessary strength to respond to joint feedback, then a similar dilemma develops. Individuals with quadriceps weakness often lean forward and thrust their knees into hyper-extension, thereby eliminating the need for muscular control of the joint, relying instead on the ligaments of the knee to maintain the extended position. Here again, an additional excessive forefoot varus creates greater tibiofemoral internal rotation and knee abduction, placing even greater stress on the ligamentous structures of the knee. Over time, both poor sensation and poor muscular control increase the risk of joint laxity, articular cartilage damage, and eventual deformity. There were two frontal plane moments that demonstrated greater symmetry between the intervention conditions. The bilateral difference between maximum ankle supination moments in the shoe-only condition was twice that of the orthotic condition (Figure 5-11), and the difference between bilateral maximum ankle pronation moments was three times

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81 greater in the shoe-only condition (Figure 5-12). Interestingly, these figures clearly show that the more-involved leg had greater magnitudes of change as a result of orthotic intervention, yet statistically there were no significant differences found here. A review of descriptive statistics (Tables 4-5 and 4-6) reveals that there was greater variability on the more-involved side than the less-involved side and, therefore, no significant differences were found. In both these examples, there was a reduction in the joint moments as a result of orthotic intervention. Here again, regardless of whether the reduction was due to a decrease in the GRF magnitude and/or a decrease in the moment arm distance, a smaller joint moment is a positive outcome. Decreasing the moment of resistance of the joint will help decrease the stresses on the joint structures. More-Involved Less-Involved Shoe Intervention Figure 5-11. Maximum ankle supination moment for both sides and interventions.

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82 Orthotic Shoe Intervention Figure 5-12 Maximum ankle pronation moment for both sides and interventions. E lectromyog ra phy Due to a small sample size and lack of statistical analysis, no solid conclusions are drawn based on EMG data. Instead, consistent trends between the orthotic and shoeonly conditions are noted. Each of these trends is discussed in light of available research, and the possible impact of each is explored. Qualitative analysis of EMG changes in the moreand less-involved lower extremity during both orthotic and shoe-only conditions revealed consistent differences only in the mean and maximum TA muscle activity. Mean RF activity was the only other muscle to demonstrate a consistent change with orthotic intervention, but it was limited to the lessinvolved lower extremity. Mean MH amplitudes in the less-involved lower extremity, alternatively, were not substantially altered in any subject during orthotic use. The remaining muscles of interest demonstrated variable responses to orthotic intervention.

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83 Mean and maximum TA amplitudes were increased during the orthotic condition in both the more and less-involved lower extremities. An increase in mean TA activity is consistent with findings reported by Nawoczenski and colleagues^ in a small sample of runners with structural malalignment of the foot. These researchers found that orthotic intervention caused an increase in mean TA activity during the first 50% of running stance. It was suggested that the increase in mean activity may have been due, in part, to the TA's role in resupination of the foot just before terminal stance, and that an earlier conversion from eccentric to concentric control may account for the noted increase in mean amplitude.^ Closer inspection of one subject's mean TA activity on the lessinvolved lower extremity indicates that this explanation may not hold true in this sample of ambulatory individuals with chronic incomplete SCI. Anterior tibialis activity from one stride during each of five walking trials was smoothed, rectified, averaged, and graphed for both the orthotic (Figure 5-13) and shoeonly (Figure 5-14) conditions. Results indicate that while the eccentric control during loading response was of similar mean magnitude and duration for both interventions, the onset of concentric activity during terminal stance actually occurred later during the orthotic condition. In this individual, the larger mean TA amplitude during one stride resulted from a larger maximum TA amplitude during initial swing, higher levels of activity throughout the remainder of swing, and a longer stride time recorded during the orthotic condition (1.04 s versus 1.00 s). Possible explanations for increases in maximum TA activity during the orthotic condition include the rigidity of the posting material, eccentric control of the foot, the role of the muscle in early stance, and orientation of the foot in terminal stance. Komi and colleagues^ found that increasing the rigidity of a shoe's heel counter caused an increase in EMG amplitude just prior to and at heel strike. The subject presented here actually had lower TA amplitudes at initial contact, which were followed by a rapid

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84 increase during loading response. Though this subject did not exhibit higher maximum TA amplitudes in early stance, the timing and magnitude of activity in other subjects may have differed somewhat. 700 , 1 i Figure 5-13. Mean tibialis anterior EMG activity of one subject on the less-involved side during the orthotic condition. Data represent the average of five strides. 700 % Gait Cycle Figure 5-14. Mean tibialis anterior EMG activity of one subject on the less-involved side during the shoe-only condition. Data represent the average of five strides. Nawoczenski et al^ and Komi et al^ have suggested that the presence of an increase in TA activity duhng early stance may have been necessary to eccentrically

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85 control the forefoot's rapid descent to the floor when orthotics were in use. Other researchers,'*^' ®^ however, have shown that orthotic intervention reduces foot pronation velocity during early stance (a finding not supported among these subjects). A decrease in pronation velocity should, theoretically, reduce the eccentric TA activity needed to control the foot during early stance. For the subject presented in Figures 5-13 and 5-14, there was an initial decrease in TA activity followed by an increase during loading response to the level measured in the shoe-only condition. The lower mean TA activity was recorded during the first 6% of the gait cycle, while the foot was being lowered to the floor, and the increase in activity occurred when the foot was flat and was loaded with body weight. The spike in TA activity in early stance during the orthotic condition may also have been the transition from an eccentric to a concentric role. A delay in the orthotic condition may represent a more effective push-off from the contralateral leg that delayed the need for concentric TA activity. This suggests that the changes in eariy mean TA activity in this subject may have been caused by a combination of concentric activity to pull the tibia anterioriy and by the altered sensory stimulus resulting from the rigid orthotic. An increase in mean and maximum TA activity was seen during the swing phase in the same subject. Nawoczenski and colleagues^ proposed that this type of result may be due to a change in foot position during terminal stance. Several studies of orthotic intervention and its effect on gait kinematics have demonstrated a reduction in maximum pronation during stance."^' Nawoczenski and associates^ have suggested that a reduction in the maximum pronation angle may place the TA in a position of greater mechanical advantage that leads to an increase in activity. Additionally, extrapolation of Komi and associates'^ work suggests that increases may also be a function of the rigid forefoot posting material. It is possible that intrinsic placement of high-density forefoot posting may

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86 cause an increase in TA activity during terminal stance that carries over into the swing phase. No research has been found that addresses this possibility. Verification of each of these potential explanations would require experimental research focused on each specific issue and remains beyond the scope of this study. Regardless of the cause of increased swing phase TA activity, however, this population of individuals would benefit from greater dorsiflexion range of motion during the swing phase of gart to reduce the pathological compensatory mechanisms employed to clear the foot from the floor. Other joint motions that would affect swing phase mechanics in this group of subjects are hip and knee flexion. Greater hip and knee flexion angles may decrease the need to circumduct the leg or hike the hip (both of which were used by this group of subjects) to provide an adequate toe clearance. Higher mean RF activity was noted on the less-involved side in four of the five subjects while no gross changes were seen in the mean MH activity. The RF is a two-joint muscle that acts as a knee extensor as well as a hip flexor. Graphical representation of one subject's RF EMG (Figures 5-15 and 5-16) demonstrates that the maximum RF activity was higher during eariy stance in the shoe-only condition, but that the mean activity was greater in the orthotic condition due to its more regular activity throughout the stance phase. The same trend was noted during the swing phase of gait. In addition to the more uniform activity seen during stance and swing, the timing of the RF activity varied notably between the two conditions. During the orthotic condition, the RF became most inactive at approximately 40% of the gait cycle and then increased near initiation of the swing phase at 60% of the gait cycle (Figure 5-15). Activity during the shoe-only condition was high during eariy stance, tapered off to a quiet state at approximately 40% of the gait cycle, and did not increase again until neariy 78% of the gait cycle (Figure 5-16).

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87 800 0 10 20 30 40 50 60 70 80 90 100 % Gait Cycle Figure 5-15. Mean rectus femoris EMG activity from one subject on the less-involved side during the orthotic condition. Data represent the average of five strides. Neither the orthotic nor the shoe-only RF EMG graphs have normal phases or amplitudes of activity. Of the two conditions, however, the orthotic condition more closely approximated the normal patterns described by Winter^^ (Figure 5-17). During normal gait, the peak RF activity is reached during loading response as the muscle works to control knee flexion. Activity then tapers throughout the remainder of stance while the muscle assists in extending the knee. Once the knee is extended and the body's center of mass is passing over the foot, the RF activity decreases to minimal levels. This is followed by an increase in pre-swing activity when hip flexion is initiated and then tapers off slowly as the swinging foot and leg are decelerated. The final increase in RF activity occurs during terminal swing as the knee is extended in preparation for initial contact.®^ During the orthotic condition peak RF activity did occur during the stance phase, but more importantly, the timing of activity at swing initiation was more appropriate than during the shoe-only condition. This eariier activation should improve hip flexion motion and decrease the compensatory mechanisms necessary to clear the foot from the floor

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88 during the swing phase. There are no known experimental data that address the effect of orthotics on RF activity to corroborate this finding. 800 700 ^ 600 1-500 S 400 g 300 a> S 200 100 0 10 20 30 40 50 60 % Gait Cycle 70 80 90 100 Figure 5-16. Mean rectus femoris EMG activity from one subject on the less-involved side during the shoe-only condition. Data represent the average of five strides. > 3 o 150 m 100 u. 0£ ID 50 Figure 5-17. Rectus femoris EMG during normal walking (n=16).Adapted from Winter. In an experiment performed by Duysens and Pearson/^ lower limb flexor activity during treadmill walking was decreased in premammillary decerebrated felines as the result of a constant stretch placed on the triceps surae. Maintaining a constant stretch on the triceps surae mimicked continuous loading of the limb. When the nervous system did not interpret unloading of the test hindlimb, all flexor activity in that limb first decreased

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89 and then ceased altogether. When all flexor activity was abolished, the hindlimb remained in a position of full extension while the feline ambulated on three legs. Duysens and Pearson^^ speculated that proprioceptors in the triceps surae inhibited the rhythmic activity created by the CPG. In the current sample of ambulatory individuals with SCI, mean RF hip flexor activity was increased during the orthotic condition primarily because it had an earlier onset during the swing phase of gait than that seen during the shoe-only condition. An increase in TA activity as a result of orthotic intervention, as discussed previously, was also seen when the orthotics were in use. It is possible that the increase in flexor activity was a direct result of altered triceps surae afferentation acting on the flexor burst generators of the CPG. Obviously, the present research was not designed to explore these findings. A study focused specifically on the impact of orthotics on EMG and central pattern generation would need to be performed to adequately address this proposition. Summary and Conclusions Orthotic inserts have played a significant role in the treatment and rehabilitation of orthopaedic dysfunction among able-bodied individuals for years. They have been used both as a preventive measure for correcting poor mechanical alignment and in response to pain caused by dysfunction. Much of the research that has focused on determining the impact of orthotic intervention has primarily been concerned with the quantitative measurement of rearfoot motion. Few studies have looked at the impact of this intervention on frontal plane lower extremity joint mechanics and on the functional parameters of gait. This study represents the first attempt to measure immediate changes in clinical gait parameters and frontal plane mechanics in a sample of ambulatory individuals with chronic incomplete SCI. Despite the small sample size, the

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90 current study demonstrates that several of these parameters can be altered immediately by orthotic intervention in ambulatory individuals with chronic incomplete SCI. The findings presented justify further research in this area. Orthotics had an immediately measurable effect on clinical gait parameters and kinematic and kinetic parameters, as was hypothesized, in a small sample size of this population. Increasing the sample size would increase the internal and external validity of the results. Longitudinal data, on the other hand, would provide necessary information regarding long-term effects on the gait mechanics in this population. Immediate changes that were noted in the less-involved lower extremity in this subject sample were hypothesized to be the result of greater sensory and muscular function in this limb. Post-testing this group after a longer period of time would provide more conclusive results regarding the effect of orthotics on the more-involved lower extremity and the time necessary to see these results. Other future research opportunities include the study of orthotic intervention in other neurologically involved populations. Individuals with multiple sclerosis, Gillian-Barre, cerebrovascular accident, and head injuries are examples. Any intervention that has the potential to improve quality of life should be investigated in detail. Thus far, orthotic intervention has been found to be an inexpensive and efficient method of improving mechanical alignment of the lower extremities. As with any research, there are many limitations to this study. The primary limitations are the sample size and the variability in this subject population. Secondary limitations include those typical to biomechanics, such as marker movement, surface electrode noise caused by skin movement, and accuracy of the calibrated space and the digitizing process. These sources of en-or were controlled to the greatest extent possible in a human experimental setting.

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APPENDIX A CONSENT FORM

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Orthotic Gait Analysis: Adults, 18 years and older Informed Consent to Participate in Research IRBU You are being asked to take part in a research study. This form provides you with information about the study. The Principal Investigator (the person in charge of this research) or a representative of the Principal Investigator will also describe this study to you and answer all of your questions. Before you decide whether or not to take part, read the information below and ask questions about anything you do not understand. Your participation is entirely voluntary. 1. Name of Participant ("Study Subject") 2. Title of Research Study Do Orthotics Have an Immediate Impact on the Gait Mechanics of Individuals with Chronic Incomplete Spinal Cord Injury? 3. Principal Investigator and Telephone Nuniber(s) Kristen Jagger MS, PT (352)392-0580 ext. 1321 (University of Florida) (678) 642-3642 (Georgia State University) 92

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93 4. Source of Funding or Other Material Support University of Florida 5. What is the purpose of this research study? The purpose of this study is to describe in detail how people with chronic incomplete spinal cord injuries walk over ground with modified footwear versus normal footwear. Three different types of shoe inserts or ankle braces will be worn while you walk over ground in order to compare the effects on your walking pattern. No specific studies have been performed that look at these types of inserts/braces and their effect on walking. This study is directly interested in how much improvement is noticed in your walking with each of the different types of inserts. You are being asked to participate in this study because you have a chronic spinal cord injury and you are able to walk at least 150 feet with or without an assistive device. The total number of persons with chronic spinal cord injuries who may take place in this study is ten. 6. What will be done if you take part in this research study? All testing will be completed either in the University of Florida Biomechanics Lab located in Florida Gym or in the Biomechanics and Ergonomics Laboratory located on the Georgia State University campus in Atlanta, GA. Handicapped parking is located directly outside the building and a campus parking pass will be provided. Testing will occur during one day, taking about four hours. You will need to wear dark colored shorts for the testing session. If you do not have any, they will be provided. A dark colored sleeveless shirt will be provided. When you first arrive at the laboratory the Principal Investigator will ask you questions about your current walking ability and what kind of assistive devices you use. You will then be weighed and various measurements will be taken (the length and distance around your feet, legs, and trunk). Next, reflective markers in the shape of small balls will be placed at various points on your shoes, legs, hips, and shoulders. These markers are used to tell the researchers exactly where your joints are so that all of the necessary numbers can be calculated. Electrodes will also be placed on specific muscles to determine how your muscles react to the different types of footwear. In order to get a good reading from the electrodes, your legs will be shaved in small areas with an electric razor and then cleaned with alcohol. You will first be asked to walk on a treadmill at a comfortable speed for 1 0 minutes while muscle signals are recorded. Next, you will be asked to walk on a walkway, approximately 20 feet long, while four video cameras record your movements and a metal plate in the walkway records how much force you step with. You will be asked to

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94 walk along the walkway at your normal pace a total of five times for each different shoe insert. A five-minute rest period will be given between each walking trial and a fifteenminute rest and break-in period will be given when shoe inserts are changed. These rest periods are minimums; you will be allowed more rest upon request. 7. What are the possible discomforts and risks? The risks involved in this study are no greater than they are when you are walking at home or in the community. There is a minimal risk of falling with all walking. A licensed and trained physical therapist will be present during each testing session to decrease this risk by guarding during transfers and walking. Rest periods have been built into the testing session and you will be allowed fiirther rest at any point during testing by request. There is a slight risk of skin irritation due to the use of adhesive tape with the reflective markers and electrodes. For participants who are not accustomed to mild exercise, there is a slight risk of muscle soreness or fatigue after walking. Any soreness or fatigue should disappear quickly. Heart rate and blood pressure are monitored throughout all procedures. A history of autonomic dysreflexia is reviewed for participants and signs of this event are monitored and discussed with each individual. Throughout the study, the researchers will notify you of new information that may become available and might affect your decision to remain in the study. If you wish to discuss the information above or any discomforts you may experience, you may ask questions now or call the Principal Investigator or contact person listed on the fi-ont page of this form. 8. What are the possible benefits to you and others? The information from this research may provide you with proper foot positioning within your own shoes. Proper positioning places the foot in a neutral and more stable position. The information gathered in this study will show if this foot placement improves your walking pattern and may help improve the treatment of chronic spinal cord injuries in the future. The information may also help improve rehabilitation that involves the use of bracing during the recovery of walking in those with spinal cord injuries. 9. If you choose to take part in this research study, will it cost you anything? Testing conducted in the biomechanics lab will be provided at no charge to you. 10. Will you receive compensation for taking part in this research study? There will be no monetary benefit if you take part in this study. You will be paid for mileage to and fi-om the lab, and lunch will be provided to you at no cost. In addition, you

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95 will be asked if you wish to keep the in-shoe orthotic that is custom made for you during this testing session. 11. What if you are injured because of the study? If you experience an injury that is directly caused by this study, only professional medical care professional dental JL. professional consultative care will be provided to you free of charge by Kristen Jagger, MS, PT (352) 392-0580 xl321 (University of Florida) or (678) 642-3642 (Georgia State University). Any other medical costs incurred will be the responsibility of the participant. 12. What other options or treatments are available if you do not want to be in this study? Participation in this study is entirely voluntary. You are free to refiise to be in the study, and your refiisal will not influence current or fiiture health care you receive at this institution. 13a. Can you withdraw from this research study? You are free to withdraw your consent and to stop participating in this research study at any time. If you do withdraw your consent, there will be no penalty, and you will not lose any benefits you are entitled to. If you decide to withdraw your consent to participate in this research study for any reason, you should contact Kristen Jagger at (352) 392-0580 ext. 1321 (University of Florida) or (678) 642-3642 (Georgia State University). If you have any questions regarding your rights as a research subject, you may phone the histitutional Review Board (IRB) office at (352) 846-1494. 13b. If you withdraw, can information about you still be used and/or collected? If you withdraw from the study, information that has already been gathered in the University of Florida Biomechanics Lab or the Georgia State University Biomechanics and Ergonomics Lab may be used as part of the dissertation in progress. 13c. Can the Principal Investigator withdraw you from this research study? You may be withdrawn from the study without your consent for the following reasons: •You need a medical treatment not allowed in this study. •The investigator decides that continuing in the study would be harmful to you. •Study treatments have a bad effect on you.

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96 14. How will your privacy and the confidentiality of your research records be protected? Authorized persons from the University of Florida and the Institutional Review Board have the legal right to review your research records and will protect the confidentiality of them to the extent permitted by law. Otherwise, your research records will not be released without your consent unless required by law or a court order. If the results of this research are published or presented at scientific meetings, your identity will not be disclosed. With your permission, you will be videotaped during this research. Your name or personal information will not be recorded on the videotape and confidentiality will be strictly maintained. However, you should be aware that the showing of these videotapes may result in others being able to recognize you. The videotapes will be kept in a locked cabinet by the Principal Investigator of this study, Kristen Jagger. These videotapes may be shown under Kristen Jagger' s direction to students, researchers, doctors, and other professionals and persons. Please sign one of the following statements that indicates under what conditions Kristen Jagger has your permission to use the videotape. I give my permission to be videotaped solely for this research project under the conditions described. Signature Date I give my permission to be videotaped for this research project and for the purposes of education at the University of Florida Health Science Center under the conditions described. Signature Date I give my permission to be videotaped for this research project, for the purposes of educafion at the University of Florida Health Science Center, and for presentations at scientific meetings under the conditions described. Signature Date 15, How will the researcher(s) benefit from your being in this study? hi general, presenting research results helps the career of a scientist. Therefore, the Principal hivestigator may benefit if the results of this study are presented at scientific meetings or in scientific journals.

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97 16. Signatures As a representative of this study, I have explained to the participant the purpose, the procedures, the possible benefits, and the risks of this research study; the alternatives to being in the study; and how privacy will be protected: Signature of Person Obtaining Consent Date You have been informed about this study's purpose, procedures, possible benefits, and risks; the alternatives to being in the study; and how your privacy will be protected. You have received a copy of this Form. You have been given the opportunity to ask questions before you sign, and you have been told that you can ask other questions at any time. You voluntarily agree to participate in this study. By signing this form, you are not waiving any of your legal rights. Signature of Person Consenting Date

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APPENDIX B SUBJECT INTAKE FORM Birthdate: Date of injury: SCI level: ASIA level: ASIA Motor Eval Left Right Hip flexors (L1 -2) I I Knee extensors (L3-4) (Knee flexors) Dorsiflexion (L4) Plantarflexion (L5-S1) Long toe extensors (L5-S1 ) (Foot eversion) (Foot inversion) ASIA Sensory Eval Left Right (L1) I I (L2) (L3) (L4) iL5) (SI) Body Segment Parameters Left (cm) Right (cm) Total height Total body mass ASIS breadth ~ Thigh length Mid-thigh Circumference Calf length ~ Calf Circumference Knee diameter Foot length Malleolus height Malleolus width Foot breadth Orthotic Evaluation Left Right Calcaneal Inversion PROM Calcaneal Eversion PROM Subtalar Neutral ROM Standing calcaneal ROM Available WB eversion FF:RR PROM (varus/valgus) Ankle dorsiflexion PROM Ankle plantarflexion PROM Knee extension PROM Knee flexion PROM 98

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BIOGRAPHICAL SKETCH Kristen had varied educational experiences in her undergraduate education prior to discovering her love for physical therapy. She initially followed the dream of being an engineer by entering Virginia Polytechnic Institute and State University as a freshman. After two years of disappointment in the lack of true "hands-on" experiences in engineering, she left to take a "time out" to find a career path that would challenge her and satisfy her need to interact with people throughout the day. During her six months away from university life, she worked as an engineering draftsperson in a private company (just to be sure about her decision) and as a physical therapy technician in a local medical center. After only one week in the physical therapy arena she announced that she had found the answer to her career search. The only remaining question involved how to get there. In order to maximize her physical therapy education experience, Kristen decided to complete a Bachelor of Science in biology and then attend a Master of Science program in physical therapy. She attended Georgia State University and then the University of South Florida to complete the biology degree. Her search for a physical therapy program took her to North Georgia College and State University, where she joined the charter class in a three-year adventure. After completion of her Master of Science, Kristen worked in a variety of settings, broadening her skills in the profession and eventually teaching students during their clinical rotations. It was her experiences with students that turned Kristen's thoughts to academic life in a physical therapy program. The search was on again. Now she had to 105

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106 find a field of study that excited her and a school with a reputation that would take her wherever she wished. Enter the University of Florida. Kristen attended the University of Florida for three and one-half years while she pursued her doctorate in biomechanics. She chose to apply her knowledge of biomechanics and her commitment to her profession to a unique problem among persons with spinal cord injury. Kristen will now carry the knowledge that she gained at the University of Florida with her as she begins her new career as a faculty member in the graduate program in physical therapy at Western Carolina University.

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I certify that I have read this study and that in my opinion it confonns to acceptable standards of scholarly presentation and is fully adequate, in scope and quality, as a dissertation for the degree of Doctor of Philosophy. ^ >>hfi W. Chow, CIAssociate Professor of Exercise and Sport Sciences I certify that I have read this study and that in my opinion it conforms to acceptable standards of scholarly presentation and is fully adequate, in scope and quality, as a dissertation for the degree of Doctor of Philosophy ames H. Cauraugh Associate Professor of Exercise ai^d Sport Sciences I certify that I have read this study and that in my opinion it conforms to acceptable standards of scholarly presentation and is fully adequate, in scope and quality, as a dissertation for the degree of Doctor of Philosophy. Mark D. Tillman Assistant Professor of Exercise and Sport Sciences I certify that I have read this study and that in my opinion it confonns to acceptable standards of scholarly presentation and is fully adequate^J^n scope and quality, as a^ dissertation for the degree of Doctor of Philosophy. Ronald A. Siders Associate Professor of Exercise and Sport Sciences I certify that I have read this study and that in my opinion it conforms to acceptable standards of scholarly presentation and is fully adequate, in scope and quality, as a dissertation for the degree of Doctor of Philosophy Robert Vander Griend Associate Professor of Orthopaedics and Rehabilitation

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I certify that I have read this study and that in my opinion it conforms to acceptable standards of scholarly presentation and is fully adequate, in scope and quality, as a dissertation for the degree of Doctor of Philosophy. This dissertation was submitted to the Graduate Faculty of the College of Health and Human Perfonnance and to the Graduate School and was accepted as partial fulfillment of the requirements for the degree of Doctor of Philosophy. December, 2002 Dean, Collegyof Health and Human Perfonnance Dean, Graduate School