1 COMPOSITE SCAFFOLDS FOR BONE TISSUE ENGINEERING By TITILAYO MOLOYE A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF DOCTOR OF PHILOS OPHY UNIVERSITY OF FLORIDA 2012
2 2012 Titilayo Moloye
3 To my mom, dad, and sisters
4 ACKNOWLEDGMENTS I would like to thank my family. Through the ups and down you guys have always been there for me, and for that, I am eternally grateful. I would like to thank my lab mates who have been so kind and generous throughout this journey. I would especially like to his guidance. I would like to thank Dr. Angela Lindner who was always to there to lend a thank you to my committee members Dr. Banks, Dr. Ghivizzani and Dr. Keselowsky. Thank you for providing your time, exper tise and advice for this project. A special thank you to Dr. Batich for his guidance, advice and encouragement throughout my time at UF.
5 TABLE OF CONTENTS page ACKNOWLEDGMENTS ................................ ................................ ................................ .. 4 LIST OF TABLES ................................ ................................ ................................ ............ 9 LIST OF FIGURES ................................ ................................ ................................ ........ 10 LIST OF DEFINITIONS ................................ ................................ ................................ 13 ABSTRACT ................................ ................................ ................................ ................... 14 CHAPTER 1 IN T ROD U C T I O N ................................ ................................ ................................ .... 16 2 T I S SUE EN G INE E RI N G ................................ ................................ ......................... 18 Intro d ucto r y R e m a r k s ................................ ................................ .............................. 18 Cell S e lection ................................ ................................ ................................ .......... 20 M aterial Se l ection ................................ ................................ ................................ .... 22 Natural Po l y m e rs ................................ ................................ .............................. 23 S y n t h e tic Po l y m e rs ................................ ................................ ........................... 23 S ca f f old F a br i ca tion M e t h o d s ................................ ................................ .................. 25 Ba ck grou n d ................................ ................................ ................................ ...... 25 Rapid P rot o t y ping/ Fus e d Depos i tion M o d e l i ng ................................ ................ 26 Sol v e nt C as ting/ P orogen Leaching M e t h o d ................................ ..................... 27 3 BONE T I S SUE ENGI N EE R ING ................................ ................................ .............. 30 Bo n e P h y s iolo g y ................................ ................................ ................................ ..... 30 The C ells o f t h e Bone ................................ ................................ ....................... 31 Bo n e T y pe ................................ ................................ ................................ ........ 32 Bo n e Mo d eling and R emodel i n g : Osteo b l a st and Oste o cl a st Interaction ......... 34 Bo n e S pe c if i c Ma r k e r s ................................ ................................ ...................... 35 Bo n e Morph o gen i c P roteins ................................ ................................ ............. 38 Fr ac t u re H e a l i ng a nd Bo n e R e pa i r ................................ ................................ .......... 39 Bo n e G r a f t s a nd B o n e G r a ft S ubstit u tes ................................ ................................ 42 Biologicall y Based G raft Replac e men t s ................................ ............................ 45 Po l y mer and Po l y mer Based G raft Repl a c e men t s ................................ ............ 47 Ceramic and Ce r amic Based Graft Rep l ac e men t s ................................ ........... 48 F D A A p pr o v al ................................ ................................ ................................ .......... 49 Summa r y ................................ ................................ ................................ ................. 50 4 M A T ERI A L S AND ME T HO D S ................................ ................................ ................ 51
6 C2C 1 2 c e l l s ................................ ................................ ................................ ............. 51 Demin e r a l iz e d Bo n e M a tr i x ................................ ................................ ..................... 52 Polycaprolactone ................................ ................................ ................................ .... 57 Solvent Casting/ Porogen Le aching Method ................................ ........................... 60 5 PRE L IMINARY DAT A S CAFFO L D O P TIMIZA T I ON A ND S P ECIF I C A I M S .......... 62 O s teoinducti v i t y a nd D B M G r a de ................................ ................................ ............ 62 Results ................................ ................................ ................................ .................... 63 S ca f f old P r e p a r a tion a nd Optimiz a tion ................................ ................................ .... 65 Was hing T e mp e r a t u re M odifi c a tion ................................ ................................ .. 71 Results ................................ ................................ ................................ .............. 72 A g gr ess i v e L eac hing u s ing S oni c a tion ................................ ............................. 73 Res ults ................................ ................................ ................................ .............. 74 Di sc us s ion ................................ ................................ ................................ ............... 76 6 CO M PO S I T E S CAF F O L DS: S Y N T H E SI S DE G RAD A TI ON A ND C A L CI U M PHOSPH A T E D E PO S I T I O N ................................ ................................ ................... 78 Intro d ucto r y R e m a r k s ................................ ................................ .............................. 78 M a ter ia ls a nd M e t h o d s ................................ ................................ ............................ 79 S ca f f old P r e p a r a tion ................................ ................................ ......................... 79 Degr a dation Media ................................ ................................ ........................... 81 S ca f f old We i ght Ch a nge ................................ ................................ ................... 81 Calcium Ion Release ................................ ................................ ........................ 81 Von K o ss a Stain i ng ................................ ................................ .......................... 82 Phosphorous Ion Release ................................ ................................ ................ 82 pH of S imul a ted Bo d y Flu i d ................................ ................................ .............. 83 Scanning Electron Microscopy with Energy Dispersive Spectroscopy ............. 83 Statistics ................................ ................................ ................................ ........... 83 Re s ults ................................ ................................ ................................ .................... 84 M as s of S ca f f olds ................................ ................................ ............................. 8 4 Calcium Ion Release ................................ ................................ .......................... 86 Von K o ss a Stain i ng ................................ ................................ .......................... 88 Phosphorous Ion C o nc e n t r a tion ................................ ................................ ....... 89 pH of Simulated Body Fluid ................................ ................................ .............. 91 Scanning Electron Microcopy w/ Elemental Dispersive Spectroscopy ............. 92 Di sc us s ion ................................ ................................ ................................ ........ 93 7 IN VITRO EVALUATION OF THE EFFECTS OF CALCIUM SULFATE, HYDROXYAPATITE, AND DEMINERALIZED BONE MATRIX ADDITION TO POLYCAPROLACTONE SCAFFOLDS ................................ ................................ 101 Introductory Remarks ................................ ................................ ............................ 101 Materials and Methods ................................ ................................ .......................... 103 Scaffold Preparation ................................ ................................ ....................... 103 Degradation Media ................................ ................................ ......................... 104
7 Water uptake and Mass Loss of Scaffolds ................................ ...................... 105 Bioactivity Analysis ................................ ................................ ......................... 105 X ray Diffraction ................................ ................................ .............................. 107 Mechanical Testing ................................ ................................ ......................... 107 Contact angle ................................ ................................ ................................ 108 Viability Analysis ................................ ................................ ............................. 108 Stati s tics ................................ ................................ ................................ ......... 108 Results ................................ ................................ ................................ .................. 109 Mass Loss and Water Absorption ................................ ................................ ... 109 Mineralization ................................ ................................ ................................ 111 Contact Angle Analysis ................................ ................................ ................... 115 Viability Analy sis ................................ ................................ ............................. 116 Mechanical Compression ................................ ................................ ............... 117 Discussion ................................ ................................ ................................ ............. 119 8 ANALYSIS OF OSTEOINDUCTIVE ACTIVITY OF POROUS POLYCAPROLACTONE DEMINERALIZED BONE MATRIX SCAFFOLDS FOR BONE REPAIR: A PRELIMINARY STUDY ................................ ........................... 124 Introductory Remarks ................................ ................................ ............................ 124 Scaffold Preparation ................................ ................................ ....................... 124 Sample Preparation ................................ ................................ .............................. 125 Cell Culture ................................ ................................ ................................ ..... 125 Statistics ................................ ................................ ................................ ......... 126 Alkaline Phosphatase Activity (ALP) ................................ ............................... 126 Alizarin Red S Staining ................................ ................................ ................... 127 Calcium Assay ................................ ................................ ................................ 127 SEM Analysis ................................ ................................ ................................ 128 Results ................................ ................................ ................................ .................. 128 Alkaline Phosphatase Expression ................................ ................................ .. 129 Alizarin Red S Staining ................................ ................................ ................... 130 Quantification of mineralization ................................ ................................ ....... 132 Scanning Electron Microscopy ................................ ................................ ....... 133 Discussion ................................ ................................ ................................ ............. 134 9 SUMMARY AND FUTURE DIRECTIONS ................................ ............................. 139 Introductory Remarks ................................ ................................ ............................ 139 Synthesis, Characterization and In Vitro Evaluation of Composite Scaffolds 140 In Vitro Evaluation of the Effects of adding common bone void fillers to a Polycaprolactone scaffold ................................ ................................ ............ 141 C2C12 Cell Proliferation and Differentiation on Composite Scaffolds ............ 143 Final Conclusions and Future Directions ................................ ............................... 144 APPENDIX: PROTOCOLS ................................ ................................ ......................... 145 L I S T O F REFERENCES ................................ ................................ ............................. 147
8 BIOGRAPHICAL SKETCH ................................ ................................ .......................... 164
9 LIST OF TABLES Table page 2 1 C o m m o n f a b r i c a ti o n me t h o d s f o r p o l y me r i c sc a f f o lds. ................................ ....... 26 3 1 Cur r ent B on e Gr a f t substi t ute s ................................ ................................ .......... 45 5 1 D i str i buti o n o f NaCl particle size ................................ ................................ ......... 67 5 2 V o lume o f P CL to N a Cl in 25 % ( w / v ) PCL So lu t io n ................................ ........... 68 6 1 Ion concentratio n of Human Plasma and SBF according to Kokubo. ................. 81
10 LIST OF FIGURES Figure page 2 1 T he ess e n t ials o f tiss u e e n g ineer i n g ................................ ................................ ... 19 2 2 Sc h e m atic o f T issue En g ineer i ng ................................ ................................ ....... 19 2 3 M esen c h y mal s t em cel l s in d uced to d i ffere n t iate in v it r o ................................ ... 21 2 4 Fabr i cat i on m e th o d f or s ol v ent c a stin g / particula t e l e ach i n g ............................... 29 2 5 Sc h e m atic o f scaffold fabr i ca t e d t hrou g h t h e sol v ent c a s t in g / particula t e leaching met h od. ................................ ................................ ................................ 29 3 1 Compact b one and sp o n g y bone ................................ ................................ ........ 33 3 2 T he sta g es o f se c o n dary heal i n g ................................ ................................ ........ 42 3 3 Al l o g ro ................................ ................................ ................................ .............. 46 3 4 I mm ix E x t ende rs ................................ ................................ .............................. 48 3 5 Os t eo s e t 2 ................................ ................................ ................................ ....... 49 4 1 The demineralization of bone ................................ ................................ ............. 55 4 2 C o m merc i al l y a v a i lable DBM pr o ducts ................................ ............................... 56 4 3 P ercen t a g e o f DBM i n c om merc i al l y a v a i lable DBM pr o ducts ............................ 56 4 4 D egradation of PCL thro ugh bulk erosion and surface erosion ......................... 58 4 5 PCL degradation via hydrolysis ................................ ................................ ......... 59 4 6 C rystalline fragmentation of PCL ................................ ................................ ....... 59 5 1 Alkaline phosphatase activity ................................ ................................ .............. 64 5 2 Osteopontin Production ................................ ................................ .................... 64 5 3 Scanning electron micrograph of PCL scaffold ................................ ................... 66 5 4 Porosity of PCL scaffolds as a function of NaCl content ................................ ..... 70 5 5 SEM microgr aphs of scaffolds washed at various temperatures ........................ 72 5 6 Porosity of PCL scaffolds as a function of washing temperature ....................... 73
11 5 7 S EM micrographs of scaffolds leached at various sonication times .................... 74 5 8 Porosity of the PCL scaffolds as a function of sonication ................................ ... 75 6 1 S EM M i c r o g ra ph s o f De m i ne ral i z e d Bon e Ma t r i x ................................ .............. 80 6 2 Schematic of solvent casting/ salt leaching technique ................................ ........ 80 6 3 Mass l oss of scaffolds over 7 weeks ................................ ................................ ... 85 6 4 Calcium ion concentration in simulated body fluid ................................ .............. 87 6 5 Von Kossa staining o f scaffolds ................................ ................................ .......... 88 6 6 Calcium and phosphorous ion concentrations of SBF of pure PCL scaffolds. and 25% DBM ................................ ................................ ................................ ........... 89 6 7 P ho s p h o ro u s i o n c on c en tra t i o n in si mu l a t e d b o d y f luid ................................ ...... 90 6 8 pH of immersion liquid (SBF) for the various scaffolds. ................................ ...... 92 6 9 S c ann i n g e lec t ron m ic r o g ra ph s o f pur e PCL and composite scaffolds ............... 92 6 10 Heterogeneous nu cl e a ti o n o f c a lc i u m pho s pha t e o n de m in e r a l i z e d bon e ma t r i x 95 7 1 Dissolu tion and Reprecipitation of amorphous calcium phosphate on Hydroxyapatite ................................ ................................ ................................ 102 7 2 Molecular Structure of Alizarin Red S dye. ................................ ....................... 106 7 3 Mass loss of scaffolds immersed in SBF for 20 days. ................................ ....... 110 7 4 Water absorption of scaffolds immersed in SBF for 20 days. ........................... 111 7 5 Alizarin Red S staining ................................ ................................ ...................... 112 7 6 Absorbance values for stained scaffolds. ................................ ......................... 113 7 7 SEM w/EDS analysis of scaffolds ................................ ................................ ..... 114 7 8 XRD of fillers before and after immersion in SBF ................................ ............. 115 7 9 C ontact angle results ................................ ................................ ........................ 116 7 10 T he viability of cells on PCL, 35% DBM, 35% HA, 35% CS scaffolds and Tissue Culture Plate at 4 days. ................................ ................................ ................... 117 7 11 Before and After Stress Strain Analysis of Composite Scaff olds Immersed in SBF ................................ ................................ ................................ ................... 118
12 7 12 Ultimate Compressive Strength of Scaffolds before (B) and After(A) immersion in SBF. ................................ ................................ ................................ .............. 121 7 13 XRD of bone demineralized for various hours ................................ ................. 123 8 1 Differentiation of C2C12 cells cultured on tissue culture plate .......................... 129 8 2 Alkalin e Phosphatase activity on scaffolds ................................ ....................... 130 8 3 Cell seeded scaffolds at 14 days. ................................ ................................ ..... 131 8 4 Quantification of mineralization on scaffo lds ................................ ..................... 132 8 5 Surface of cell seeded 35% DBM scaffold. ................................ ....................... 133 8 6 Surface of cell seeded 35% DBM scaffold. ................................ ....................... 134 8 7 C omposite picture of modified protocol for both ALP expression and mineralization quantiification ................................ ................................ ............. 136
13 LIST OF DEFINITIONS B I O ABS O RB AB LE M a t e r i a ls o r d e v ices t h a t c a n d issol v e in b od y f l u ids w it hou t an y po l y me r ch a in cl ea v a g e o r mo lec u l a r m a ss d e c r ea s e B I O DE G RADA B LE M a t e r i a ls t h a t b r e a k do w n du e t o macromol e c u lar de g ra d a ti o n w ith d isp e rs i on b u t no t ne c e s s a r i ly to t a l r e mo v a l f rom t h e b od y B I O E R O DI B LE M a t e r i a ls t h a t s h o w s u r f a c e d e g ra da ti o n a n d res o rb f u r t he r in v i v o B I O E R O S I O N M a t e r i a ls a n d t h e ir b y p ro du cts t ha t a re e l i m i n a t e d c o mp le t e ly f r o m t h e b o d y with n o r e sid ua l si d e e f fe ct s B I O M A T E RI A LS M a t e r i a ls t h a t c o m e in c on t a c t w ith b iol o g ical s y st em s o r are u s e d i n m e d ical de v ices. B I O RE S O RB AB LE M a t e r i a ls t h a t s h o w bu lk de g ra da t i o n an d are completely removed from the body. B I O RE S O RP T I O N A c on c ep t o f t o t a l e l i m i na ti o n o f initi a l f o r e i g n m a t e r i a l a n d o f bu lk de g ra da ti o n b y p ro du c t s ( l o w m o lec u lar w e i g h t c o m p o und s) w ith n o resi d ua l si d e e f fe cts
14 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 CO M PO S I T E S CAF F O L DS FOR B O N E T I S SUE EN GINEERING By T iti l a y o Mol o y e A u g u st 2 01 2 Ch a i r : Chr i st op h e r B a tich M a jor: Bi o med ical E n g in ee r i n g D i s ea s e s, a cc i de n ts, a n d a g ing a re t h e l ead i n g c au s e s o f b on e d e f e cts t h a t c a n w ea k e n th e b o ne na t u ral a n a t o m y and a s a res u lt, i mpa ir its f un ct i on 5 10 % o f t h e 6 m i l l i o n d e f e cts o c c u r r ing ea ch y ea r r e q u i r e f u rt h e r tre a t m e n t f o r c o m p ro m i s e d hea l i n g d u e to l o ss o f bo n e me t abo l i c d ist u r ban c e s, a n d i m p a i r m e n t o f b l o o d s upp l y Cur r en tl y t h e s t an d a rd s o lu t i o n f o r f i x ing s u ch d e f e cts r e q u i r e s multiple s u r g e r i e s and the removal of healthy tissue. An e f fe cti v e a n d v ia b le t issu e en g in e e red b on e sc a f fo ld w ou ld p r o v ide a p la t f o rm f o r b o n e f o r m a ti o n a s w e ll a s a n i n trinsic si g na l f o r pl u r ip o t e n t c e l l s t o d i f fe r en ti a te i n to bo n e f o r m ing c e l l s. In t h is st ud y a c o m po s i t e sc a f f o ld o f po l y c ap rol a ct on e (PCL) a n d dem i ne ral i z e d b o n e m a trix (DBM) is p ro po s ed u si n g a s o l v en t c a sti n g a n d pa rticula t e l e a c h ing t e c hn i q ue The broad hypothesis behind the construct of this composite scaffold was simple; the combination of an osteo inductive/osteoconductive material coupled with a biocompatible polymer would yield a scaffold that 1) delivers growth factors to a targeted site, 2) encourages and enhances the formation of mineral deposits on its surface, and 3) allows the proliferation and differentiation of undifferentiated cells on its surface.
15 Initially, composite scaffolds of three different DBM concentrations, 25, 35, and 50 wt %, were prepared to evaluate scaffold degradation and surface mineralization. Scaffolds containing 35% DBM both loss and gained mass over the 7 week study. This gain was likely due to mineral being deposited on its surface. Moreover, analysis of the X ray diffraction patterns of the DBM particles before and after immersion in SBF revealed reprecipitation on its surface after 20 day However, despite the additional ions released from the DBM into the simulated body fluid available for reprecipitation, we report no correlation in DBM concentration and surface mineralization. Lastly, analytical assays of alkalin e phosphatase and total calcium content revealed high values of expression on composite scaffolds when compared to the pure PCL scaffolds. The work described herein provides strong support for a composite scaffold with the potential of bonding to bone in v ivo and inducing the differentiation of undifferentiated cells into an osteoblastic lineage.
16 C H AP T ER 1 IN T ROD U C T I O N It is estimated that by 2030 there will be about 72.1 million individuals over the age of 65 in the United States, which is more than doub le the 2000 estimate . With the increasing percentage of an aging population, the need for viable bone substitutes is also increasing. While major strides have been made in the development of functional bone substitutes in particular bone grafting and metal implants they are not without drawbacks. Many patients undergo surgeries in which autologous bone harvested from donor sites, such as the iliac crest, is used to replace the bone defect. Although these grafts offer a nominal replacement for the bo ne defect, they fall short on a number of individual, viable host tissue may not be available. Moreover, the site from which the a term coined for any 5]. While the bone is capable of self repair, the ability for new bone to form usually dwindles with age, and peaks at the age of 30. Aft er 30, the only way for new bone to a term that refers to the process by which osteoclasts break down old bone, thereby releasing its mineral content into the blood stream [6 8]. Osteoblasts then use the minerals (calcium) in the blood stream to form new bone. Essentially, osteoclasts break down the house to get the raw materials for the osteoblasts to build a new house [9, 10]. Another solution to bone replacement is the use of bone received from donors, otherwise known as al logenic bone. In this case, the cadaver bone is thoroughly cleansed, free from surface proteins, cells, or any biological entities that might precipitate
17 an immunogenic response . The cleansing does, however, expose the influencing growth factors (bone morphogenic proteins) within the bone . Although this approach is considered a viable alternative to autografts, it also falls short of being a first rate method, primarily because allogenic bone hinges on the availability of cadaveric bones. In addit ion, according to Haimi et al. , the caustic ingredients (harsh chemicals) that obliterate the biological contaminants on the bone produce an adverse effect on the mechanical integrity of the bone, leaving it in a weakened state  For the above rea sons, an approach that does not use host tissue and is not dependent on the availability of cadaveric bones is greatly needed. Several new methods are currently being explored, including the combination of growth factors, stem cells, and synthetic material s to develop a novel scaffold for repairing bone defects. The engineering of bone tissue requires an effective and productive relationship between the osteoinductive growth factors, the osteoblastic or osteoprogenitor cells, and the osteoconductive scaffo ld. The osteoconductive matrix/scaffold could serve as a mechanical support system and as a platform for the attachment and growth of the cellular component, thus guiding the newly formed bone tissue in addition to operating as a delivery vehicle for the c ells or growth factors. Growth factors are important because they can recruit, differentiate, and aid in the proliferation of the appropriate cell types at the defect site. The cells, donated or autologous, are directly responsible for helping form the new ly engineered bone tissue [15 18]. Hence, these three factors (osteoconductive matrices, osteoinductive growth factors, an d osteoblastic/osteoprogenitor cells) and their subsequent combination constitute the foundation to a successful tissue engineered bon e substitute.
18 C H AP T ER 2 T I S SUE EN G INE E RI N G Intro d ucto r y R e m a r k s Tissue Engineering applies the principles of engineering and life sciences to the development of biological substitutes that restore, maintain, or repair tissue function . The emergence o f tissue engineering can be directly attributed to the problems associated with donated tissue. These problems include severe rejection of the donated tissues, inflammation of the area in which the donated tissue was implanted, and the critical shortage of organ donation . Millions of people suffer from end stage organ failure or tissue loss each year. An average of 17 people die each day from the lack of available organs for transplant . The field of tissue engineering seeks to remedy these proble ms by engineering compatible tissues capable of replacing damaged tissue. Tissue Engineering combines living cells with a support structure that degrades in vivo, leaving behind a 3D assembly that is functionally, structurally, and mechanically equal to or better than the native tissue it replaced . Cell selection and material selection are two key components for a successful tissue engineered replacement. Material selection can be further divided to include signaling molecules within the material a nd fabrication method by which the material is used to make the scaffold. Figure 2 1 illustrates the components of tissue engineering. Figure 2 2 details the overall concept of tissue engineering. It also highlights the main components of tissue engineerin g
19 Fi g ure 2 1. T he ess e n t ials o f tiss u e e n g ineer i n g The ma i n c o m p o n ents o f cell s election and m a ter i al sele c tion, w hich can be f ur t her d i v i ded into si g nal i ng molecules w ithin matrix and f a br i cation m e th o d. Fi g ure 2 2. Sc h e m atic o f T issue En g in eer i ng process us i ng au t olo g ous cel l s. START END
20 Cell S e lection The source autologous or donated of cells has been extensively researched and debated in the scientific community for years. Proponents of donated cells argue that this method will pave the way for off the shelf products. On the other hand, proponents of autologous cells argue that these cells exhibit little to no immune rejection, have a higher proliferation rate, and result in little to no donor site morbidity . Many argue that for academic se ttings and non profit work autologous cells would be ideal, while through a purely economic prism allogenic cells are preferred . Autologous cells are also at a disadvantage when placed in a severely damaged area, since in this case the pool of hea lthy cells is small and can therefore limit the regeneration of the damaged organ or tissue. While autologous cells ar e not always available, and donated cells can cause an immune rejection, allogenic mesenchymal stem cells are readily available and are le ss prone to immune rejection than donated differentiated cells. Mesenchymal stem cells exert powerful immunomodulatory effects which include the inhibition of proliferation and function of T cells, B cells, and natural killer cells hence making them vir tually immune from causing an undesirable effect. These cells, derived from bone marrow, have the capacity to proliferate rapidly and differentiate into various pathways . The path from a mesenchymal stem cell to an osteoblast (a bone forming cell) is detailed in Figure 2 3.The cells are dependent upon the stimulants and its environment. Stimulated, mesenchymal stem cells can form fat cells (adipocytes), bone cells (osteoblasts) and cartilage producing cells (chondrocytes). However the differentiation o f stem cells is not permanent. Therefore, it is possible for the differentiated cell to revert back into an undifferentiated cell.
21 Fi g ure 2 3. M esen c h y mal s t em cel l s in d uced to d i ffere n t iate in v it r o. Stimulat e d w ith v ar i ous biolo g ical a g e n ts, MSCs can f o r m f a t cel l s (adipoc y tes); b one forming cel l s (os t eobl a sts); or cartila g e pr o ducing (cho n droc y te s ) .
22 M aterial Se l ection One of the major components of a viable tissue engineering alternative is the material used to make the scaffold. Henceforth, th is material will be referred to as a biomaterial. The biomaterial should be non carcinogenic, e licit little to no inflammatory response, be compatible and, preferably, be biodegradable and bioresorbable . The biomaterial should also be accommodating to certain parameters warranted for each scaffold, such as pore size and scaffold shape ability. The scaffold plays a critical role in the overall effectiveness of the tissue engineered product. The function of the scaffold is to direct the growth of cells s eeded either within the porous structure of the scaffold or migrating from surrounding tissue . Scaffolding materials must also be able to serve as a conduit for the consistent release of growth factors, antibiotics and other therapeutic agents . S uch materials can be natural polymers, synthetic polymers, ceramics, or a combination of polymer/ceramic. Grafting the process of moving tissue from one area of the body to another is the standard operating procedure for replacing damaged tissues. Ther e are three types of grafting procedures: autografting, allografting, and xenografting . Autografts are tissue that comes directly from the host. In the replacement of bone, tissue is removed from the iliac crest and transferred to the diseased site [2 9]. Consequently, this creates a new defect site, and for many patients this site causes more pain than the site in which the tissue is to be transplanted. Secondly, this procedure increases hospital time and because it is essentially two surgeries inc reases costs. Allografting uses donated cadaveric tissue to replace or repair a damaged tissue. One of the drawbacks of allografts is that by using donated tissue, it increases the likelihood of an immunological rejection and/or an infection A significant amount of research has gone
23 into sterilization methods that decrease adverse reactions in the recipient [3 0 32 ]. Xenografting the process of using tissue from one species to an unlike species is a far less common grafting procedure. The most prevalent problem with xenograft, as with allograft, is immunological rejection. Moreover, there are a number of ethical issues surrounding xenotransplantation [3 3 ]. Possibly due to the above mentioned reasons, 4 ]. The general idea behind tissue engineering is to create an autograft, thereby eliminating all of its related problems. There are a number of biomaterials that have been used for tissue engineering purposes: metals, ceramics, and polymers (natural and sy nthetic). For the purpose of this project, we will focus on natural and synthetic polymers. Natural Po l y m e rs There are a number of naturally occurring polymers that are used for tissue engineering. Examples of natural occurring polymers include starches, p roteins, the extracellular matrix (ECM), alginate, gelatin, chitosan, dextran, silk, chondroitin 6 sulfate and albumin. These polymers have good biocompatibility and offer fewer regulatory constraints compared to synthetic polymers [35 36]. Natural polymer s possess the ability to present receptor binding ligands to cells, are bioactive, and remodel naturally. One of the few drawbacks of natural polymers is that, unlike synthetic polymers, they are not as easily able to conform to certain parameters such as porosity or attain a level of mechanical strength needed for the implant area. S y n t h e tic Po l y m e rs Natural polymers are relatively inexpensive. However, they can vary widely depending on processing, and are more likely to be contaminated from unknown virus es or diseases, in comparison to synthetic polymers . Synthetic polymers can be easily
24 manufactured, generally have no immunogenicity, and their physical attributes such as molecular weight and molecular structure can be easily controlled . Th eir processing also reduces variability amongst batches and can be supplied in large quantities . In designing scaffolds for tissue engineering, the material must be both biodegradable and bioresorbable. These terms, as well as other tissue engineered related terms, are defined Most of the widely used polymers currently utilized in tissue engineered products are synthetic biodegradable polymers because, through their degradation, they lead to the formation of natural tissue and they curtail the adverse chronic foreign body reaction some natural and synthetic non biodegradable polymers produce . The properties of the ideal material are as follows: Biocompatible Mechanical properties that similar/consistent with the tissue it is replacing Bioresorbable Degrade at a pace that matches new tissue formation Possess the ability to change and manipulate structure properties such as pore size, porosity, pore connectivity Despite the progress made on polymers and their use in tiss ue engineered products, many problems still remain. Concerns over migration, encapsulation by fibrous tissue and distortion are well documented in synthetic polymers used in cardiovascular treatments . Additionally, the effectiveness of the material se lected also depends on whether the material possesses signaling molecules, either native to the material or physically and/or chemically embedded in the polymer. Signaling molecules such as growth factors, morphogens and adhesins promote and/or prevent cell adhesion, proliferation,
25 migration and differentiation by upregulating or downregulating the synthesis of protein, growth factors, and receptors [4 1 ]. Moreover, the adhesiveness of the material is important. The majority of mammalian cell types are anchorage dependent, meaning that they will die if an adhesive surface is not provided [4 2 ]. The method of scaffold fabrication is also important in how well the material behaves. Certain cell types thrive at different pore sizes. For example, studies hav e demonstrated the varying pore sizes in which osteoblasts (380 (380 the scaffold is another important aspect that needs to be taken into consideration. The scaffold with a high surface to volume ratio typically favo rs cell adhesion, proliferation, and differentiation . S ca f f old F a br i ca tion M e t h o d s Ba ck grou n d Perhaps no factor of tissue engineering is as important as the scaffolding of the matrix. While the type of material is very significant, how the scaffold is fabricated is of equal importance. Scaffolds designed to regenerate body tissue by seeding cells into porous matrices allow the growth of cells and tissue modeling through a combination of structure and signaling. The construction of pore size and pore di stance can be very useful in letting nutrients and cells to seep into the matrix; these can also function as a buffer by preventing waste and bacteria from flowing into the scaffold . Compared to 2D constructs, culturing cells in 3D constructs results in a microenvironment that more closely resembles that found in vivo. This is mainly due to the fact that seeded cells respond to mechanical and biological cues from the 3D locations to promote cell adhesion, proliferation, and differentiation. The ability to engineer
26 a scaffold and dictate parameters (porosity, pore distance, pore size, hydrophilicity) enhances the functionalities of cells and tissues to support adhesion and growth. Porous structures provide space, permit cell suspension, promote ECM produ ction, transport nutrients, and eliminate waste products . While there are a number of techniques (Table 2 1) that achieve the goal of a porous structure, for the purpose of this study we will focus on those that were specifically considered. T a b le 2 1 C o m m o n f a b r i c a ti o n me t h o d s f o r p o l y me r i c sc a f f o lds. M e c han i s m M e t ho d A d v a n t a g e s/Disa d v an t a g e s Lea c h ing Me th o d S o l v en t Cas t in g /s a lt l e a c h ing me t ho d (e g Ice pa rtic l e le a c h ing me t h od g a s f o a m in g s a lt l e a c h in g g e l p ressing m e t h o d ) A d v : E a si l y c on tr olla b l e po rosity an d ge o m e tr y D i s: Lo ss o f w a t e r sol ub l e b iomolec u les o r c y t o k ine s du r i n g po ro g e n l ea c h i n g s a lt ret en ti o n P ha se S e pa rat i o n M e t ho d T h e r ma l l y i ndu c e d p h a se s epa rat i on c en tr i f u g a t i o n me t ho d A d v : lo ad ing o f h y d r o p h i l ic o r h y d ro phob ic b io a ct i v e mo le c u les. D i s: S m a ll po re si z e an d d i f f iculty in c on troll i n g s t ruct u r e P r i n ti n g an d Prot o t y p ing I n kjet p r i n ti n g p roc e ss, m e lt ba se r a p id p rot o t y p ing A d v : S pe c i f ic pa r a me t e rs (p o re si z e d ist an c e ) c a n b e a c h ie v ed r e p ro du c i b le Dis: Cos t l y hea t ma y c au s e de n a t u r i n g o f p rot e i n s M ic r o s phe re M e t ho d Bio de g ra d a b le m ic r o s p he re, ma cr o po ro u s b e ad p a rtic l e a gg re g a t e d sc a f f o l d A d v : H ig h m e c h a n ical st ab i l ity D i s: Lo w po rosity Rapid P rot o t y ping/ Fus e d Depos i tion M o d e l i ng Rapid prototyping (or, solid free form fabrication) is a technique that produces 3D shape specific scaffolds. This technique is rapidly gaining popularity with tissue engineers. It consists of using a 2D image to make a 3D product, with parameters set and customized by the user. There are various specific methods within the rapid prototyping umbrella, some of which employ UV curable photopolymers, lasers, or printers . Since the technology is relatively new, there are both advantages and disadvantages.
27 In the related literature, 3D prin ting is perhaps the most ubiquitous concept. Founded at MIT, it involves a layer by layer process in which the sliced 2D profile of a computer model is printed on a fresh layer of powder via deposition of a suitable binder [48 49]. Successive layers are de posited on top of the previous layer until the model is complete. One of the disadvantages of this method is that it does not produce a strong bond between the particles on the model which can produce an ill fitted scaffold and collapse in vivo. In addit ion, the cost of fabricating a scaffold using a 3D machine is fairly expensive. Fused Deposition Modeling (FDM) operates along the same layer by layer pathway as 3D printing. A plastic filament or metal wire is unwound from a coil and it supplies material to an extrusion nozzle, which can turn the flow on and off. The nozzle is then heated to melt the material and can be moved along both horizontal and vertical directions by a numerically controlled mechanism . One of the drawbacks of FDM is the large a mount of material needed to make a scaffold, which can make for an expensive and unrealistic alternative to bone grafts. Sol v e nt C as ting/ P orogen Leaching M e t h od As a result of the large amount of material needed, the solvent casting/ salt leaching method was chosen. This technique can produce polymeric scaffolds with a high porosity and varying pore sizes. A suitable porogen is initially combined with a polymer solution in a mold for the assembly of solid polymer scaffolds. The porogen is then subsequently leached out to form highly porous sponges for the cultivation of cells. One of the advantages of this technique is that the porogen (sodium chloride, sodium bicarbonate) is not soluble in the organic solvents typically used to dissolve the biodegradable p olymer [5 1 ]. Nevertheless, an accumulation of un dissolvable porogen could lead to an uneven
28 distribution of pores on the scaffold. Care must be taken to make sure the porogen is evenly dispersed within the polymer solution. After casting the polymer poro gen mixture, the porogen can be further purged by a washing series using distilled water. Once the washing step is completed, what is left is a highly porous, water insoluble scaffold. The purpose o f this re s earch w as to de v elop a hi g hly porous sc a ffo l d f o r t he culti v ation of cells. Si n ce t h e p oro g en l e achi n g techni q ue a c hie v es t he e nd re s ult w ithout the use o f a la r g e am o unt o f pol y mer, w e decided t o fabr i ca t e our s c a f f o l ds using this te c hni q ue. Moreover, porogen leaching is mainly useful for biomater ials in the developmental stage. Figure 2 4 and Figure 2 5 represent the preparation and processing of the polymeric scaffold using the solvent casting/ porogen leaching method.
29 Figure 2 4. Fabrication method for solvent casting/ particulate leaching. The p olymer and particulate (sodium chloride) are dissolved in an organic solvent and cast in a drying mold. The sodium chloride is then leached out of the scaffold using a washing series. The porous scaffold is freeze dried and then placed in a vacuum oven F i g ure 2 5. Sc h e m atic o f scaffold fabr i ca t e d t hrou g h t h e sol v ent c a s t in g / particula t e leaching met h od.
30 C H AP T ER 3 BONE T I S SUE ENGI N EE R ING Bo n e P h y s iolo g y Bone is a dynamic organ. Formation, growth, maintenance, healing, and repair of the skeleton require t hat bone be formed throughout the life of an individual . Bone serves five distinct purposes in the human body: it offers mechanical support, it anchors muscle attachments, it functions as a safeguard for internal organs and marrow, and it provides sto rage for certain elements (mostly calcium) . Bone can be generally classified into two types: cortical and cancellous (trabecular) bone. They are classified in part on their porosity, rigidity, and microstructure. Cortical bone accounts for 80% of bone and cancellous bone accounts for the other 20%. On a microscopic level, bone is composed of several different cell types, each possessing a function and extracellular matrix (ECM) [54, 55]. The ECM is mostly composed of Type I collagen, along with small traces of types III and V collagens at different stages of bone formation. Proteoglycans (5%) and noncollagenous proteins are also included in the ECM. The noncollagenous proteins can be divided into several categories. These include proteoglycans, glycosy lated proteins, and glycosylated proteins with potential cell attachment activities [56, 57]. Two thirds of bone matrix are made up of calcium phosphate (Ca3 (PO4)2). Calcium phosphate reacts with calcium hydroxide (Ca (OH)2) to form crystals of hydroxyapa tite (Ca10(PO4) 6 (OH)2). The remainder of the bone matrix (approximately one third) is collagen. Bone cells only make about 2% o f b o ne ma s s [58 ]
31 The C ells of t h e Bone The cell types associated with bone development include osteoblasts, osteocytes, bone lining cells and osteoclasts. Below is a list of key bone cells and their definitions. Osteoblasts : Immature bone cells that secrete the matrix by the process of osteogenesis (secretion of proteins and other inorganic compounds of the matrix). The role of these secreted proteins is not clearly understood. Many have attributed their function to include the regulation of bone mineral deposition and turnover, and the regulation of bone cell activity [59 ]. Osteoblasts originate from bone marrow stromal cells or mesenchymal stem cells, proliferating and differentiating first into preosteoblasts, and then to mature osteoblasts. They can be found on the lining areas of newly formed and unmineralized tissue (mostly type I collagen) [6 0 ]. Osteocytes: When immature bo ne cells (osteoblasts) are surrounded by bone they matrix. This is during both synthesis and resorption of the matrix. Osteocytes live in lacunae between layers of the mi neralized matrix; they have extensive filopodial processes that lie within the canaliculi in mineralized bone and allow them to maintain connection with each other and with the bone surface [6 1 ]. Osteocytes do not turn over and may live for decades in the human bone. It has also been suggested that the presence of empty lacunae in aging bone means that osteocytes undergo apoptosis [6 2 6 3 ]. Bone lining cells: Perhaps the least understood of the cell types are Bone Lining Cells (BLCs). The surface of bones, which is not under remodeling or formation, is covered by BLCs. Characteristic of the BLCs is their elongated, thin, and flat appearance. At the heart of the confusion surrounding BLCs is their function. Some studies have
32 speculated that BLCs play an integ ral part in the marrow stromal system and have important functions in hematopoiesis, while others suggest that BLCs help in the maintenance of bone fluids and in the fluxes of ions between the bone fluid and the interstitial fluid compartments of mineral h omeostasis. BLCs have also been marked as inactivated osteoblasts [6 4 66 ]. Osteoclasts: Differentiated from hematopoietic stem cells found in the circulating blood, osteoclasts are the cells responsible for the resorption of bone. Histologically, they are roughly 20 100 m in size. Osteoclasts can have a number of different shapes, ranging from flat to rounded shaped cells, depending on the phase of the resorption cycle they are in [6 7 ]. They possess two distinct plasma membrane areas: a ruffled border and a clear zone (also known as a sealing zone). The clear zone surrounds the ruffled border area facing the resorption lacunae. Through the ruffle border, the osteoclasts secrete hydrochloric acid and collagenase to dissolve the inorganic matrix. This slightl y acidic, pH 4.0 4.5 secretion dissolves the hydroxyapatite crystals that form the mineralized extracellular bone matrix. Next, the various enzymes and collagenase s released decimate the collagen matrix, thereby securing the hydroxyapatite crystals to the bone. The degradation products are then removed from the resorption lacunae and released into the extracellular matrix [6 8 7 1 ]. Bo n e T y pe Due to t h e ri g idity and hard n ess o f b o ne, t h ere is a perce p tion th a t b one is not i v Quite t h e co n trar y bone is m a de o f i v o nl i v ing e l e m e C ortical bo n e is ma d e u p o f os t eons, w hich are cel l s c om p os e d o f f rail la y ers o f t he m e mbr a ne cal l ed l a mel l ae. The m emb r ane pro t ects t he blood a n d n er v es o f t h e inter n al core o f t he com p act bo n e. All ost e ons in l ong b o nes r u n t he l e n g th o f the bone, strengthening the
33 bone in that direction [7 2 ]. Cortical bone is quite dense and has a porosity of 5 10% . It is 80 90% mineralized. Cortical bone is very strong w ith the elastic m o dulus and y ield str e n g th o f t h e cortical b o n e ran g ing bet w een a reported 10 40 GPa and 90 140 MPa respectively [7 3 ] and its main function is to maintain the mechanical and protective requirements of the skeleton Cancel l ous, or tr a becu l ar, b o ne is less d e nse and has a p orosity o f b et w een 5 0 80%. Bec a use o f its hi g h por o sit y it is also k no w n as sp o n g y bone. It is f o u n d o n t h e inter i or o f mo s t b one s and only 1 5 25% o f its cont e nts are mineral i z ed C a ncel l ous bone has no blood vessels in its trabeculae. The space between trabeculae is filled w ith ano t her tiss u e, t he r e d b o n e mar r o w w hich has bl o od v essels a nd s uppl i es nutri e nts t o the ost e oc y tes. A schematic showing the relationship between cortical and cancellous bone is shown in figure 3 1 [7 4 75]. Fi g ure 3 1. Compact b one and s p o n g y bone (Taken from http://training.seer.cancer.gov/anatomy/skeletal/tissue.html )
34 Bo n e Mo d eling and R emodel i n g : Osteo b l a st and Oste o cl a st Interaction Bone Model i ng r e fers t o al ter a tions i n b one s ha p e thr o u g h t h e in de pe n dent actions of osteoblasts and osteoclasts. In skeletal growth, removal (osteoclasts) and replacement (osteoblasts) occur at a fairly rapid pace. The rate of turnover of the skeleton approaches roughly 100% in th e first year of life, and dramatically decreases to about 10% per year in late childhood to early teens. This rate is further reduced as time progresses. As a matter of fact, by the age of 25, most women hit their maximum bone peak mass, and by the age of 30 most men hit this peak as well. Therefore, most but not all of the turnover of bone during growth occurs from bone modeling. Bone remodeling accounts for skeletal growth from adulthood to death [7 6 7 8 ]. While bone modeling is based on the independen t, coordinated actions of osteoclasts and osteoblasts, bone remodeling is possible through a dependent and coordinated interaction between osteoclasts and osteoblasts. In an adult skeleton, osteoclasts and osteoblasts belong to an assembly recognized as a basic multicellular unit or a BMU. The BMU is composed of osteoclasts (located in the front), osteoblasts (located in the rear), a nerve supply, a central vascular capillary, and connective tissue. The BMU, beginning at a certain origin, progresses towards its target for bone replacement and continues a distance thereafter until it self terminates. There is a difference in bone remodeling in cortical and cancellous bone. The BMU travels through the bone in the cortical bone, hollowing out a tunnel. However, in cancellous bone, the BMU is largely a superficial surface process m ore of a trench than a tunnel [79 82 ]. More specifically, the activation of preosteoclastic cells in the bone marrow initiates the bone remodeling process. Interluekin 1 (IL 1), Parat hyroid Hormone (PTH), and a few other cytokines are released from inactive bone lining cells. The
35 preosteoclastic cells then migrate to the surface of the bone. These cells mature into osteoclasts which then develop a ruffle border. The osteoclasts cover a space on the matrix surface and secrete hydrogen ions and cathepsin into that space. They in turn resorb the mineral (Ca 2+) on the surface of cancellous (trench) and cortical bone (tunnel). The formation or reconstruction of bone begins when PTH stimulat es the osteoblasts to produce IL 6 and other cytokines that stimulate osteoclasts to resorb bone. This occurs alongside the activation of the preosteoclastic cell by IL 1. This stimulation of osteoclasts produces channels in which blood capillaries grow. The osteoblasts, in turn, line the tunnels so that blood may flow to the bone. The osteoblasts secrete type 1 collagen and various matrix proteins [8 3 8 6 ]. The collagen, now polymerized, forms triple stranded fibers to create an osteoid. When the osteoid i s about 6 microns thick, it begins to mineralize [8 7 ]. Eventually, the mineralization of calcium and phosphate ions precipitates to create crystals of hydroxyapatite (HA). Lastly, the HA traps the osteoblasts into this newly formed bone matrix where they m ature into osteocytes . Bo n e S pe c if i c Ma r k e r s While there are no specific markers that are exclusively produced by bone in both its modeling and remodeling processes, there are a number of proteins in varying concentrations that exist only in bone. A b rief description of the most common markers is included below. Alkaline phosphatase (ALP): Perhaps the most common bone marker, alkaline phosphate is an enzyme that is characteristic of bone formation osteoblasts. It is a cell surface protein that is bound to the plasma membrane through phosphatidylinositol phospholipid complexes. An increase in blood ALP expression could reflect increased
36 bone formation. However ALP expression is not limited to bone. The synthesis of ALP in the liver, kidney, placenta, int estine, serum and other tissues makes it commonplace in the body, which subsequently makes it fairly difficult to take measurements and interpret phosphatase activity arises from the liver and the remaining 50% stems from the bone [8 8 ]. Although the exact function of ALP in bone is not completely understood, there are a number of theories that speculate on its possible role. Primarily, ALP expression has been recognized as pl aying an important part in skeletal mineralization. It is suggested that ALP hydrolyzes an unknow n phosphate ester that increases the local concentration of inorganic phosphate. This substantial increase in available phosphate is purported to facilitate th e formation of the bone mineral hydroxyapatite. Moreover, some studies have suggested that ALP hydrolyzes a calcification inhibitor identified as inorganic pyrophosphate. Others state that ALP may act as a transporter for inorganic phosphate to bind to Ca 2+ to help accelerate ca lcium phosphate precipitation [89 92 ]. Osteocalcin, bone gla protein (BGP): a mid ma r ker for bone formation. It is one of the proteins that make up the bone matrix along with collagen. It is only synthesized by bone and dentin. Yet, similar to ALP, the exact function of BGP is not entirely known. BGP is one of the most abundant non collagenous proteins found in the bone, and has a very high affinity for Ca2+. BGP appears during the latter stages of osteobla st differentiation [93 ]. Os teopontin (OPN): a phosphorylated glycoprotein that is found throughout the body, including but not limited to chondrocytes, skin, brain cells, and kidneys. Synthesized by a cadre of preosteoblasts, osteoblasts and osteocytes, OPN is secreted into the bone and the osteoid ECM [9 4 ]. OPN promotes the attachment and spread of
37 osteoblasts to and through the ECM. Some researchers have found that OPN production occurs before osteocalcin, which would suggest it is a part of bone resorption. In studies that res earched bone growth, OPN was found in large amounts at the borders of growth plates in growing bone. In studies that dealt with bone healing, OPN was found in the borders of a defect between new and native bone. Finally, in studies that focused on bone gro wth between a foreign object and nati v e bo n e, OPN w as detected bet w een t h e imp l ant (foreign object) a n d t h e e x isting bone [ 9 5, 9 6]. Bone sialoprotein (BSP ): a phosphorylated and glycosylated protein that mediates cell attachment to the ECM. Similar to othe r markers, it is not specific to bone. It can be found in chondrocytes, hypertrophic cartilage, and osteoclasts [9 7 ]. It can also be found in mature osteoblasts and osteocytes during the beginning stage of mineralization. Furthermore, it has high affinity for calcium ions [9 8 ]. Osteonectin: (SPARC), osteonectin is a glycoprotein that is expressed in osteoblasts, vascular smooth muscle cells, endothelial cells, megakaryocytes, chondrocytes, steroid ogenic cell s, and placental trophoblasts [99 ]. Similar to the other bone markers, it can bind to calcium ions, hydroxyapatite, and collagen. Studies have shown that osteonectin regulates cellular progression through the cell cycle. It has also been documen ted that osteonectin may be involved in the binding of growth factors to cells [100 101 ]. Collagen type I: Collagen type I is the most abundant protein found in bone, comprising 95% of the extracellular matrix of bone. Its triple helix structure gives bon e its high strength. It is not specific to bone and can be found in t endons, ligaments, and skin [102 ].
38 Bo n e Morph o gen i c P roteins Growth factors are defined as polypeptides that provide two functions: they stimulate cell proliferation and provide major gro wth regulatory molecules for cells in culture and in vivo [10 3 ]. Bone contains many growth factors including but not limited to transforming growth factor beta (TGF like growth factor, fibroblast growth factor and bone morphogenic proteins (BMPS) [10 4 ]. For the purpose of this research, we will focus on BMPs. BMPs are a subgroup of the TGF available when bone has been demineralized. Initially, BMPS were identified by Urist as molecules that indu ce bone when implanted in an ectopic site in a rat model [10 5 ]. A number of ensuing studies conducted on BMPs have found them to be able to recruit and stimulate mesenchymal cell progenitors to differentiate into osteoblasts, therefore suggesting that BMPs are also indirectly involved in bone remodeling and healing [10 6 1 08 ]. BMPs are extracellularly regulated by Noggin, Chordin and DAN, two BMP binding proteins. Through signaling transduction, BMPs bind to a heterodimeric complex of two transmembrane serin e threonine kinase receptors: BMP receptor type (BMPR) I and II. Type I and II receptor kinases phosphorylate the transcription factors S m ad 1, 5 a nd 8. T h ese pho s p h or y lated Sm a ds t hen f o r m a he t erodi m er i c co m plex w i t h S mad 4 in t h e nucleus a n d acti v ate t h e e x pression o f v ar i o u s g enes [ 109 1 13 ] T h ere ha v e b e e n at least 1 5 B M Ps i de n t i f i e d in m a m m als [ 1 14 ] B M P2 a n d B M P 4 ha v e be e n t h e most studi e d a n d c a talo g u e d. T he y are also t he B M Ps w hich have been shown to induce the m ost b o ne form a tion in an ec t opic s ite [ 1 15 1 1 6].
39 Fr ac t u re H e a l i ng a nd Bo n e R e pa i r Bone is n o t a s tatic or g an; ra t her, it is con s ta n tly re m o d el i ng itse l f Ev en w hen da m a g ed, b o ne c an r e pair a n d h eal itse l f. T h e h e al i ng o f a f rac t ured bo n e oc c urs v ia a primary or secondary healing, with the latter process consisting of several stages [11 7 ,118 ] Pr i mary heal i n g also k no w n as di r ect h e al i n g is only initiated u n der t he most optimal of biological conditions. In this type of healing, the cortex seeks to re establish itself a f t er d am a g e. Al s o, u n l i ke sec o n d ary heal i n g no e x ter n al cal lu s is for m ed; ho w e v er, it re q ui r es an en v i r onm e nt o f r ig id stab i l i z ation [1 1 9]. Y e t s ince most fractures require some sort of motion (sling or cast immobilization, external or intra m e dul l ary f i x ation) t o h eal, pr i ma r y heal i ng is at y pical [ 1 20]. W e can f urt h er di v ide pr im ary heal i ng into t w o g roups: g ap h eal i ng and co n tact h ea l i n g In g ap heal i n g t he ent i re g ap is f i l led w ith di re ct b o ne for m ation. Init i al l y f or m ation is acco m pl i sh e d w ith wo v en bo n e, fol l o w ed by par al l e l f iber e d a nd / or la m el l ar bo n e for sup p ort. W i thin the g ap, t h ere a r e n o traces o f f ibro c arti l a g e or c o n necti v e tissue. T he or i entat i on o f t his new bo n e f o r mati o n is tr a n sv erse to the or ig inal (n ati v e) la m el l ar b one or i entat i on. After several weeks, lo ngitudinal Harvesian remodeling rebuilds the necrotic fracture ends and replaces the woven/lamellar bone with osteons of the original orientation. Gap healing results in a bone that is fairly sim ilar to the bone it replaced [121 124 ]. In contact healing, h ealing occurs where fragments are in direct apposition to each other and osteons grow across the fracture site and parallel to the long axis of the bone. This is accomplished without the intermediary step of new bone being formed transverse to the original bone (via gap healing). First, the osteoclasts resorb irreparably damaged/
40 necrotic bone on one side of the fracture site noted a s tunneling resorptive response The tunneling allows the penetration of capillaries and leads to new Ha ve r sian systems Along with these blood vessels are endothelial cells and osteoprogenitor cells for osteoblasts. Finally, this leads to the production of osteons across the fracture site. The end result of gap healing and contact healing is regeneration of native, pre fr actured bone architecture [122, 12 5 12 7 ]. Secondary bone healing: Secondary healing, far more likely than primary healing, occurs when there is no rigid fixation of the fractured bone ends. Secondary healing leads to an external callus formation. Due to it s complexity, there are three distinct stages of secondary healing: an initial inflammatory phase, an intermediary reparative phase, and a remodeling phase [12 8 ]. There are a number of growth factors involved, some which were detailed in Sec 3.1.4. For the purpose of this research, we will briefly highlight the key points in secondary healing. Stage I Inflammation: Before the inflammation stage can begin, a fracture must occur. Fracturing of the bone occurs when the bone absorbs energy beyond its modul us of elasticity (MOE) [1 29 ]. The impact damages the local bone marrow, the periosteum, neighboring soft and hard tissue, and also disrupts blood vessels. Due to the disruption of blood vessels, a hematoma occurs [13 0 13 1 ]. The ends of the vessels undergo th rombosis, formation of a blood clot inside the vessel. Once this occurs, lysosomal enzymes are released. A steep drop in pH results in an acidic environment. Since this is a largely cellular event, a group of cells respond to the fracture site. Macrophages platelets, monocytes and other inflammatory cells begin the cleanup and repair process .
41 Stage 2 Reparative: In secondary healing, unlike primary healing, an external callus is formed. A callus is a layer of fibrous tissue that forms around the fr acture. Initially, in the primary soft callus formation stage, the aforementioned cells are stimulated to produce the soft callus a few days after bone fracturing. Granulation tissue fibroblasts begin to form cartilage and fibrocartilage around the edges o f the callus and within the fracture site. However, this callus is quite weak to external stresses and can be so for roughly 6 weeks [13 2 ]. After this period, the cartilage and fibrocartilage are converted into woven bone through an endochondral ossificati on process. At the same time, bone is formed through the recruitment and proliferation of mesenchymal stem cells and preosteoblasts within the callus. Depending on the type of bone and age of patient, this stage can take anywhere from 4 to 16 weeks [13 3 ]. Stage 3 Remodeling : The final stage replaces the callus with bone. In remodeling, the environment, pH, and vascularization revert back to normal. Via osteoclast remodeling, the size of the callus is decreased and eventually replaced with woven bone. Even tually, the woven bone is replaced with lamellar bone. This is done through the aforementioned BMU process involving osteoclasts and osteoblasts [13 4 ]. Complete remodeling can take anywhere from 4 months to 4 years, depending on the size of the fracture [1 3 5 ]. Similarly to the reparative stage, the time it takes bone to completely remodel is hugely dependent on the age of the patient. Children can remodel bone at a much faster rate than adults. A schematic of secondary healing is shown in Figure 3 2 In bot h primary and secondary healing, the end result of a newly formed bone comparable to pre fractured bone is achieved.
42 Fi g ure 3 2. T he sta g es o f se c o n dary heal i n g I m p act, I n f l a m mati o n, Repai r a nd Re m o del. Bo n e G r a f t s a nd B o n e G r a ft S ubstit u tes Bo n e graf t s: We briefly mentioned the various types of grafting procedures. Autografting, the gold standard, is the most common in bone replacement surgery. The transplanted bone is usually harvested from the iliac crest [13 6 ]. If obtaining bone from the iliac cre st is not feasible, bone can be taken from the leg, in particular the distal femur and proximal tibial [13 7 ]. Autologous bone is advantageous because of the low risk of disease transmission. It also provides the least potential for rejection, since it cont ains Autologous bone also possesses a scaffold for new bone to grow on and into. This makes autologous bone osteoinductive, oste oconductive, and osteogenic [138 ]. Osteoinductive: Osteoinduction is the process by which mesenchymal stem cells from the graft or the host tissue are induced to differentiate into osteoblasts.
43 Osteoconductive: Osteoconduction is the process by which the graft provides a three dimensional structure that facilitates the ingrowth of capillaries a nd mesenchymal stem cells to support new bone formation on the graft. Osteogenic: Osteogenesis is the process of new bone formation by cells from the host or the graft. While it is considered the gold standard, autologous bone grafting has several major dr awbacks, the biggest being donor site morbidity. Joshi and Kostakis  found 10% of the patients who underwent iliac crest bone grafts for intra oral augmentation had pain for 16 weeks post operatively. This group included those who experienced gait dif ficulties and required the use of a walking stick, deformity at the hip site, hematomas, infection, and stress fracturing. While the 10% who suffered post operative pain is a fairly significant percentage, autografting offers the best and most effective ch ance of replacement. Allografting: While not optimal, is another source for replacement bone. Allograft bone comes from donor tissue, usually from bone banks that harvest cadaver bone. These bones are cleansed and sterilized to reduce the possibility of di sease transmission from donor to recipient. While allografts possess a scaffold for cells to grow into and onto, it may or may not provide an incentive for mesenchymal stem cells to differentiate into osteoblasts. So, while osteoconductive, allograft bone is not necessarily osteoinductive nor osteogenic. While the cells and proteins may be present in an allograft, they are not native and could possibly cause an immune rejection. Moreover, the mechanical strength of allografts is significantly decreased due t o the sterilization process, especially if gamma radiation is used. The purpose of the radiation is to cleanse the graft of bacteria and fungi, but in heavier doses it can significantly
44 weaken the graft by degrading the collagen in the bone matrix [140 ]. Clearly, the elimination of harvesting donor bone and all of its associated problems is a key advantage in allografting. However, the threat of a possible disease being transmitted from donor to recipient is a big reason allografting is not as common as au tografting. Disease transmission in allografting is rare, but can occur. Established screening methods and sterilization protocols have greatly reduced the risk of transmission of diseases in donated cadaveric tissue The risk of a person contracting HIV f rom the use of donated bone allografts is roughly 1 in 1.67 million. The probability that a bone allograft might contain HIV has been es timated at 1 in 2.8 billion [141 ]. Bo n e graft s ubstit u te s : Since allograft, autogenous grafts, and the less common xenog rafts, have significant downsides; the field of bone tissue engineering has gained prominence. Bone tissue engineering seeks to create substitutes that are biologically and mechanically comparable to native bone. While most of those currently on the marke t provide some of these attributes, there are none that can regenerate bone results comparable to those seen in autografting. There are a number of different materials that are used in orthopedic surgeries to repair bone fractures or defects. These substit utes can be naturally or synthetically derived. Many substitutes combine both types of materials in order to improve the effectiveness of the substitute. Synthetic materials that are currently on the market include metals, alloys, certain polymers and cera mics. These can be further categorized into biodegradable and non biodegradable, with each category having distinct advantages and disadvantages. Naturally derived materials that are currently on the market include h ydroxyapatite, d emineralized b one m atrix and collagen. In order to be an effective substitute, these materials must be osteoconductive, osteoinducti ve, and
45 induce osteogenesis [142 ]. A review of current bone substitute materials is detailed in Table 3 1 T a ble 3 1. Cur r ent B on e Gr a f t substi t ut e s Fi l ler T ra d e N a me / M an u f a c t u rer Ad v an t a g e s D i s ad v an t a g e s Res o rpti o n rate Cost P o l y ( l a cti c c o g l y c o l i de ) form ed by particulate leaching/phase inversion Im m i x B on e Tec Os te c o ndu c ti v e No in t r i n sic o st eo i ndu cti v e p ro pe rties 1 2 20 w ee ks $3 7 8 /lc c C olla g en ba s e d ma t e r i a ls Colla g ra f t ( Z i mme r) He a los ( D e p u y Ac r om e d ) Os te c o ndu c ti v e No struct u ral stre n g th Colla g e n f a i r ly f a st, HA slo w e r res o rpti o n $5 4 0 / 2 5 cc B on e M o rp ho g en ic Prot e ins B M P 7 B M P 2 Os teo i n du ctive No struct u ral stre n g th N/A $1 5 0 0 70 0 0 p e r k i t Calcium S u l f a t e Os teo s e t ( W r ig h t M ed i c a l) I ne x pen si v e Res o rpti o n r a te f a st e r t ha n bo n e f o r ma ti o n 4 1 2 w ee ks $4 3 5 / 5 c c St e m C e l l s Osir i s, I n c. Os teo i n du ctive No struct u ral stre n g th N/A H ig h De m in e ral i z e d B on e Ma t r i x ba s e d ma t e r i a ls R T I Al l o g r a f t P a ste ( m i x e d w / g e la t in) R T IX Os teo i n du ctiv e D i f f e r e n t lo t /Diff e re n t p ro pe rties N/A $1 5 4 / 1 c c We can further divide these groups into biologically based grafts, polymer and polymer composite based grafts, and ceramic and ceramic based graft s. Biologic all y Based G raft Replac e men t s Biologically based graft replacements are substitutes that contain a biological aspect, whether in growth factors, cells, or donated tissue. The donated tissue comes in the form of demineralized bone matrix (DBM). We will furt her elaborate on DBM in Materials and Methods The purpose of this section is to briefly introduce the various bone graft substitutes on the market. DBM is demineralized (decalcified) ground bone. It is has been shown to be osteoinductive and induce osteog enesis in rat calvarial bone defects [14 3 14 5 ]. Many products on the market currently contain either DBM or DBM
46 with a carrier material. The carrier material is needed because DBM is quite flaky and can be problematic when handling. The image below (Figure 3 3 ) is Allogro, a DBM product produced by Wright Medical, Inc. and Allosource, Inc. Fi g ure 3 3. Al l o g ro ( W r i g ht m e dical) i s a 1 0 0% DBM pro d uct. Picture taken from wmt.com (wright medical.com) A number of other products from various companies also c ontain DBM. Some of the biological based products include but are not limited to: Osteofil (Sofamor Danek, Inc.): A DBM injectable paste is a mixture containing 24% demineralized human bone matrix and 17% gelatin (of porcine origin) in aqueous solution G rafton (Osteotech, Inc.): Injectable DBM paste that contains 17% DBM with a glycerol carrier DBX Putty (Synthes, Inc.): DBM putty that contains 32% DBM with a sodium hyaluronate carrier Dynagraft Orthoblast (GenSci Inc.): Collagen polymer mixture contai ning DBM based grafts, such as Osteocel Osteocel Plus is an allograft cellular matrix that contains viable stem cells that promote fusion in cervical, thoracic, and lumbar procedures. There are also a number of products that take advantage of the BMPs located within DBM. Recombinant human
47 Po l y mer and Po l y mer Based G raft Repl a c e men t s Polymer replacements can be ca tegorized as natural and synthetic and can be further split into degradable or nondegradable. Some of the positive attributes to using polymeric replacements over grafting include: no donor site morbidity and virtually no chance of contracting a disease. A n effective polymeric replacement would be biocompatible, biodegradable, and bioresorbable. It is important that we clarify this point because a material that is biodegradable is not necessarily bioresorbable as well. As explained earlier, the ability of a material to resorb means that it is completely and naturally removed from the body. Polymer based replacements can be combined with other materials to strengthen the substitute or be inserted with biologics to make the material osteoinductive. Some of th e current polymer based products include but are not limited to: OPLA (THM Biomedical, Inc.) Polylactide based material Immix (Osteobiologics, Inc.) Poly(lactide co glycolide) based material (shown below) BoneTec (BoneTec, Inc.) Poly(lactide co glycolid e) based material OsteoScaf (Tissue Regeneration Therapeutics) poly(lactic co glycolic acid) foam matrix Cortoss (Orthovita) injectable resin based product with applications for load bearing sites
48 Fi g u re 3 4 I mm ix E x t ende rs (Os t e o b iol o g ics, In c) is a pa rticu l a te l ea c h e d p o ly (lacti c c o g l y c o l i c a cid) p ro d u ct w h ich is u s e d a s a g raft e x t en d e r. Picture taken from Smith and Nephew ( http://global.smith nephew.com/master/6600.htm ) Ceramic and Ce r amic Based Graft Rep l ac e men t s There has been a lot of interest in ceramic based materials, mostly due to their high mechanical strength. Calcium phosphate, and calcium sulfate or its derivatives are some ceramics that can be used alone or in combina tion with other type of grafts. The biological rationale for using calcium phosphate is that hydroxyapatite, a subgroup under calcium phosphate, accounts for 70% of bone [14 6 ]. Some of the current ceramic based materials include but are not limited to: Ost eograf (Ceramed): Synthetic hydroxyapatite particles and blocks Norian SRS (Skeletal Repair System) (Syntheses, Inc.): Ceramic based Calcium phosphate cement Actifuse (ApaTech Limited): Silicate substituted calcium phosphate As stated earlier, there h as been an increase in ceramic based bone graft
49 protocol for wounded warriors who have suffered open fracture wounds in battle involves the use of Osteoset (Wright M edical), a calcium sulfate based product. Calcium sulfate, more commonly known as Plaster of Paris, has been used as a bone void filler for well over 2 centuries [14 7 ]. It has gained wide acceptance as a viable substitute due to its biocompatibility, bioa ctivity and resorbability. However, under closer scrutiny, calcium sulfate has shown a faster resorption rate than reported, and had a less than appealing healing in vivo [1 48 150 ]. The manufacturer, Wright Medical, has currently started incorporating DBM into its flagship calcium sulfate based product, Osteoset In Osteoset 2 DBM graft, the percentage of DBM in the bone void filler increases from 0 % seen in Osteoset to 53% seen in Osteoset 2 DBM In the brochu re provided (shown in Figure 3 5 ), Wri ght Medical touts the benefit of F D A A p pr o v al The U. S Food and Drug Administration (FDA) currently classifies most orthobiologicals as Class II (those that do not require premarket approval) devices. Fi g ure 3 5. Os t eo s e t 2. Brochu r e f r o m W r i g ht M edical o f Ost e o s et 2 hi g hl ig hts. (Picture taken from http://www.wmt.com )
50 Several are cleared through the 510 (K) process. The DBM and cell based substitutes are categorized as human tissue and do not need clinical trials or FDA clearance for marketing [15 1 152 ]. Summa r y The purpose of bone graft substitutes is to offer an alternative to the use of bone grafts. Because of the wide array of materials available, they are able to repair, restore or regenerate bone. Currently, none of the substitutes mentioned provide a comparable result to autografting. Autografts provide an osteoconductive surface, contain osteoinductive factors, and induce osteogenesis. Because of this, it would take an array
51 CHAPTER 4 M A T ERI A L S AND ME T HO D S C2C 1 2 c e l l s The C2C12 cell line is a murine myoblastic cell line which, under appropri ate culture conditions, differentiates rapidly, forming contractile skeletal myotubes and producing characteristic muscle proteins. Although C2C12 cells are considered to be a myogenic lineage committed cell line, it is known that treatment of C2C12 cells with bone morphogenic proteins (BMP) ( such as BMP 2 BMP 7 ) induces their transdifferentiation into osteoblasts (bone forming cells) [15 3 ]. Although they possess the ability to differentiate into a number of tissues like mesenchymal stem cells, they cannot be passed forever and eventually stop differentiating. In a n informative study by Cheng et al. [15 4 ], fourteen of the human BMPs were isolated and introduced to the different cell lines to determine the osteogenic activity of each BMP. The cell lines used were the pluripotent mesenchymal progenitor cell line (C2H10T1/2), the preosteoblastic cell line C2C12, and osteoblastic TE 85 cell line. Alkaline phosphatase activity, osteocalcin production and matrix mineralization were potential. As expected, the committed osteoblastic cell line (TE 85) tested positive for alkaline phosphatase activity produced by each bone morphogenic protein. The progenitor cell line, C2H10T1/2, was much more discriminating, and elicited a detectable a mount of A LP activity with BMP 2, 6, and 9. C2C12 cells tested positive for alkaline phosphatase activity in five out of the fourteen BMPs (BMP 2, 4, 6, 7, and 9). In part, C2C12 cells are very attractive because although they can alter their pathway to an osteoblastic pathway in the presence o f DBM
52 they do so only if a high enough volume of specific BMPs are present within the sample. Demin e r a l iz e d Bo n e M a tr i x Demineralized bone matrix (DBM) is derived from cortical bone (human or bovine) and is prepar ed by removing the mineralized components of bone. This removal leaves behind a network of organic and protein elements. While the demineralization process reduces bone to powder, the ground bone possesses a structure that provides a platform for bone form atio n~ making it osteoconductive [155 ]. Perhaps the most important component of DBM is the bone morphogenic proteins (BMPs) located within the matrix. A member of transforming growth factor beta (T GF glycoproteins, BMPs provide an osteoinductive trigger for cell differentiation. While more than 12 isoforms of BMPs have been identified, BMP 2, 3 and 7 have been shown to play a critical role in bone healing [15 6 15 7 ]. The osteoinduc tive potential of DBM was first discovered by Marshall Urist in 1965. By implanting demineralized bone in rat muscle, Urist discovered ectopic bone formation, through the formation of a new ossicle in the implanted area. This further led Urist to investiga te the cause of bone formation, which led him to discover BMPs 8 ]. A number of researchers have confirmed and expanded on the work initiated by Urist. Wozney and colleagues [1 59 ] determined the genetic sequence of BMPs, which led to the identification of its various forms. Utilizing recombinant gene technology, it is now possible to mass produce the individual BMPs. While the production of individual BMPs using recombinant technology added to the significant body of work on DBM, there still exists no universal method of measuring
53 osteoinductive potential in each DBM lot. It would be impractical and costly for osteoinductivity to be measured purely, as Urist had done, through the excision of im planted tissue from the subdermal or muscle pouch of athymic mice or rats. Histological evaluation requires the expertise and knowledge of a skilled histologist or pathologist whose judgment would be exclusively based on the explanted material. Therefore, a great number of sections must be explanted in order to provide a thorough and accurate representation of the results. The more practical method of quantifying osteoinductive potential is to assess the in vitro activity of alkaline phosphatase (ALP). Paga ni et al. [16 0 ] discovered increased levels of ALP activity in areas of greater osteogenesis, a direct result of elevated osteoblast activity. Cellular proliferation cannot be used as a determinant of osteoinductivity because this is a result of mitogenic factors, not intrinsic factors such as BMPs. The demineralization of DBM is a surface mediated acid solubulization reaction that advances from the outermost mineralized surface of the bone particle towards its center. As available surface decreases, so doe s the amount of mineral available to be solubilized [16 1 ]. Theoretically, during this process, bone would be fully demineralized. process of theoretical and actual groun d bone is depicted in Figure 4 1. The residual calcium in DBM could act as a site of h eterogeneous nucleation for the redeposition of apatite, thereby increasing the likelihood of bonding to native bone in vivo [16 2 ]. In this regard, it can be assumed that the larger the DBM particle size, the higher the residual calcium content and, therefore, the more available sites for nucleation. This line of reasoning, however, is not applicable to osteoinductive potential. In a
54 groundbreaking study by Zhang et al., [16 3 ] the osteoinductive potential of ground DBM particles of sizes <250 microns, 250 355 microns, 355 500, 500 710 microns and 710 850 microns were analyzed using ALP activity of human periosteal cells. It was determ ined that the particles in the range of 500 710 microns instead of the larger 710 850 microns provided the highest level of Alkaline Phosphatase activity. While the use of DBM for bone related surgeries has skyrocketed in the past decade, initial usage was nearly impractical due to its handling properties. Commercially available ground DBM comes in a number of sizes, none above 710 microns. Even at its 4 ]. In vivo, th e delivery and migration of DBM from the implanted could cause the defect site The first commercially available DBM product with a carrier was Grafton Gel (Oste otech Inc., Eatonton, NJ) This first generation p roduct used glycerol as the carrier system to effectively deliver DBM to the implant site [16 5 ]. Since the launch of Grafton in the early 1990s, a number of companies have developed and manufactured DBM ca rrier products. Figure 4 2 and Figure 4 3 detail the various DBM carriers, percent by weight of DBM in the carriers, and the current cost of each product. While the potential of pure DBM can be measured in terms of osteoinductivity and osteoconductivity us ing various assays, it would be inaccurate to ascribe these properties to products that contain DBM. A study published in the Journal of Orthopaedic Research in 2003 [16 6 ] examined the ALP activity of C2C12 cells of varying amounts of activated and inactiv ated DBM.
55 DBM was inactivated using a Guanidine Hydrochloric Acid extraction. We define inactivated as the removal of all growth factors from DBM. A linear relationship was seen between activated DBM and ALP activity. This proportional osteoinduction canno t be ascribed to DBM containing products. In other words, the behavior of 100 micrograms of pure DBM is not unequivocally the behavior of a product containing 100 micrograms of DBM. Each DBM containing product must be independently analyzed for its own pro perties. Figure 4 1. The demineralization of DBM is a surface mediated acid solubulization reaction that advances from the outermost mineralized surface of the bone particle towards its center. Since DBM is of irregular sha pe, it cannot be completely demineralized
56 Fi g ure 4 2. C o m merc i al l y a v a i lable DBM pr o ducts Fi g ure 4 3. P ercen t a g e o f DBM i n c om merc i al l y a v a i lable DBM pr o ducts
57 Polycaprolactone Polycaprolactone (PCL) is a semi crystalline linear resorbable aliph atic thermoplastic synthesized by the ring opening polymerization of the cyclic monomer e caprolactone [1 67 ]. PCL has a low glas s transition temperature (~ 60 C) and a low me lting point between 59 and 64 C [ 168 ]. PCL has generated much interest because of its relatively long term degradation rate (>24 mths to lose total mass), its ability to blend well both polymers and non polymers and its unique properties which is biocompatible for both soft and hard tissue replacement [ 169 ]. PCL has been approved fo r use by the FDA since the 1970s and can be found in a number of biomedical products such as sutures, artificial skin grafts and implantable carriers for drug deliver [ 170 ]. The commercial appeal of PCL also extends to its nontoxic degradation byproducts PCL undergoes both bulk hydrolysis degradation and surface erosion degradation, with the former being the chief route of degradation and the latter being fairly rare [ 171 ]. A distinction should be made on the diffusion reaction conditions for bulk or su rface degradation to occur. In order for a polymer to undergo bulk degradation, water must penetrate the entire polymer bulk, which causes hydrolysis throughout the entir e matrix. Surface degradation requires the surface only hydrolytic cleavage of the pol ymer backbone. Moreover, the rate of hydrolytic chain scission and the production of oligomers and monomers must be faster than the rate of water intrusion into the polymer bulk [172 174]. In surface erosion, because you are essentially thinning the polyme r from the inside out, the molecular weight of the polymer is unaffected. Fig 4 4 depicts how bo th bulk and surface degradation advance.
5 8 Fig ure 4 4. Advancing degradation of PCL for bulk erosion and surface erosion. Polymers that undergo bulk erosion d egrade from the outside inward Polymers that undergo surface erosion degrade from the surface Surface erosion is akin to the degradation of a bar of soap over time. Bulk hydrolysis, initially, breaks the ester linkages, which creates fragments and releas es oligomeric species. These low molecular fragments are ultimately engrossed by macrophages and giant cells. The byproduct of the hydrolysis of PCL, 6 hydroxylcaproic acid, an intermediate of omega oxidation and the beta oxidation to acetyl coA (acetyl c oenzyme A), is metabolized via the tricarboxylic acid (TCA) cycle (citric acid cycle) or removed by direct renal secretion. Fig 4 5 depicts the hydrolysis of PCL via hydrolysis intermediates 6 hydroxylcaproic acid and acetyl coA. Fig 4 6 is a schematic o f how the fragmentation of the crystalline regions could occur. While PCL can be biodegraded by bacteria and fungi, they are not biodegradable in animal and human bodies because there are no comparable and suitable enzymes [ 175 ].This does not mean that PC L is not bioresorbable; rather it takes a longer period, as the lengthy process above demonstrates.
59 Fi gure 4 5. The degradation of PCL via hydrolysis intermediates 6 hydroxyl caproic acid and acetyl coenzyme A, which are then eliminated f rom the body via the citric acid cycle. Figure 4 6. Schematic visualization of how crystalline fragmentation could have taken place.
60 We mentioned previously the attractive characteristics of PCL that has given rise to extensive research into its pot ential application in the biomedical field; however, PCL was not always seen in such favorable light compared to other resorbable polymers. While synthesized in the 1930s, PCL became commercially available only when it was discovered that it could be degra ded by microorganisms. During this time, PCL was used in a number of drug delivery devices due to its aforementioned advantages and a few others. These included tailorable degradation kinetics, controlled release of drugs, shape ability, and the manufactu re of specific surface characteristics, such as pore size, that would enhance tissue ingrowth [176, 177] The degradation rate made PCL an ideal polymer to be used for drug delivery devices that needed to remain in vivo for over 1 year. Yet, it was the f airly long degradation rate that kept PCL on the backburner for many years. Resorbable polymers such as polylactides (PLA) and polyglycolides (PGA), that had a faster resorption rate, grew in prominence because the medical field insisted upon the polymer r eleasing the encapsulated drugs within the polymer in 1/6 th of the time PCL required [1 78 ,1 79 ] Also, the medical device industry was interested in replacing metal devices that contributed nothing biologically, with biodegradable polymers that had compara ble strength PCL was not one of them. But in a twist, with the emergence of tissue engineering and biomaterials, PCL resurface because of its long term degradation rate as well as its viscoelastic and superior rheological properties when compared with PLA or PGA. Solvent Casting/ Porogen Leaching Method We previously mentioned the solvent casting/porogen leaching technique. Briefly, PCL pellets are dissolved in glacial acetic acid at a concentration of 25% (w/v) in a beaker and stirred at 65C. Once the PC L is completely dissolved (1 2 hrs), sodium chloride (NaCl)
61 particles at 90% w/w of PCL are added. The polymer solution is allowed to cool at room temperature for 30 minutes. Once the temperature of the polymer solution reaches 37C, the polymer solution i s cast into well culture microplates (Corning Life Sciences). This allows for evaporation of the solvent, glacial acetic acid. After the evaporation of the solvent at 23C (RT) overnight, the resulting solidified scaffolds are placed in 50 mL conical tubes containing 15 mLs of distilled water, with a ratio of sample surface to DI volume of 0.5 cm2/mL. Residual solvent, contaminants and more importantly, sodium chloride particles are removed in this washing sequence. Thus, leaving behind a porous polymer st ructure. To further remove any remaining solvent and/or distilled water, the scaffolds are freeze dried overnight .
62 CHAPTER 5 PRE L IMINARY DAT A S CAFFO L D O P TIMIZA T I ON A ND S P ECIF I C A I M S O s teoinducti v i t y a nd D B M G r a de Preliminary work was carried out to test: 1) the osteoinductive behavior of DBM on human osteoblasts, 2) test and analyze the efficacy of the scaffold design, and to 3) provide a framework for the work to be undertaken. Analytical assays of early and mid bone marker, alkaline phosphate (ALP ), and osteopontin (OP) were employed to gauge the influence of three grades of DBM on human osteoblasts (hOB). Based on industry standard procedures, using histological techniques, DBM was graded according to the amount of ectopic bone formation in nude m ice. Following analysis of the explanted DBM, the lot from which the implanted DBM originated was then given a (0,1) (2,1) or (4,1) score. (4,1) being the DBM with the most ectopic bone formation in vivo and (0, 1) being the least. (2, 1) DBM is usually re served for commercial use. The grades are known as their osteoinductive (OI) score. 50 mg of sterilized bovine demineralized bone matrix of grades (0,1) (2,1) and (4,1) were placed in 24 well culture plates (Corning Costar Corp.). A day prior to cell seed Essential Medium (DMEM) (Invitrogen, Carlsbad, CA) supplemented with 10% fetal bovine serum (FBS) (Mediatech, Herdon, VA) and 5.0% Antibiotic Antimitotic (Invitrogen)). This supplemented medium will be denoted as basal media. After 24 hours, basal media was aspirated out of the DBM and hOB (CRL 11372, ATCC, VA, USA) cells at a density of 4 x 105 cells/well at passage 4 was added to each well, along with 2 mls of fresh basal media. Basal m edia was devoid of known osteogenic inductors glycerosphosphate and ascorbic acid 2 phosphate to assess the osteoinductive
63 behavior of the various DBM grades only. The well plates were placed in an incubator under standard conditions of 37C and 5% CO2. Fresh basal media was replaced every 3 days for the 17 day study. The metabolic activity of the cells was assessed initially every 4 hours for a period was a routine analysis to observe the viability and proliferation of cells, hence no data is biochemical assay, samples were pi well plate. The absorbance was measured, as directed in the manufacture protocol, at 405 nm using a Microplate reader (Molecular Devices). The ALP activity of the hOB is show in Figure 5 1. Data comparing the scaffolds was examined using two way analysis of variance (ANOVA) with statistical significance at p<0.05 (n=3). Results The ALP activity expressed in each of the three grades of DBM was statistically the same. It was interesting to note that (2,1) DBM had a higher OD (optical density) reading than (4,1) DBM. As a matter of fact, (4,1) DBM had a lower OD reading than (0,1) DBM, though not a statistically significant difference. Yet, this is a significant finding because, as previously mentioned, a D BM grade of (4,1) conveys a better than average level of ectopic bone formation in vivo. A (0, 1) DBM grade represents a less than average level of ectopic bone formation in vivo. Therefore, based on these results, the OI score does not seem to be a foreca ster of ALP activity in vitro. Osteopontin (OPN) expression was analyzed using an enzyme linked immunosorbent assay (ELISA) kit specific for human osteoblasts (R&D Systems Minneapolis, MN). Aliquots of media from appropriate samples were added to the
64 ELISA kit and quantified on a spectrophotometer at 540nm. Results were compared to a standard curve and converted to nanograms (ng) of osteopontin and corrected for cell number according to Alamar Blue proliferation assays run in conjunction. Osteopontin expres sion of the hOB is shown in Figure 5 2 Figure 5 1. Alkaline phosphatase activity. Optical density as a function of DBM grade is shown. Each data point represents the mean of three values and error bars denote S tandard Error Mean Figure 5 2. Osteopo ntin Production at 17 days. OPN production as a function of DBM grade is shown Each data point represents the mean of three values and error bars denote Standard Error Mean
65 However, similarly to ALP production, there was not a correlation between OI and OPN production. Cells seeded on (0, 1) DBM expressed the highest levels of OPN, while 4,1 DBM had the lowest although, based on the two way ANOVA analysis, the difference between the two was not statistically significant. There was a large variability i n OPN production within the grades (0,1) and (2,1) but not (4,1). While variability was expected between the grades, it was not expected within the grades. As previously mentioned, the OI score of a batch of DBM is given by assigning a grade to the explan ted tissue of a nude mouse model using histological techniques. Calcium contents of explants, as indicators of new bone formation, are used in this assessment. This in vivo experiment can take more than 8 weeks, after which the mice are sacrificed. It is p robable that the short duration of the in vitro experiments (>3 weeks) could have a neutralizing effect on the grades of DBM. Meaning that, 50 mg of (0, 1) DBM is osteoinductively equal to 50 mg of (4, 1) DBM at least for ALP production and OPN expression in vitro. If a (4, 1) DBM grade indicates a better than average level of ectopic bone formation in vivo, it could be assumed that it has a higher concentration of BMPs with bone forming potential. Yet, based on these results, the potency (bone forming pot ential) of the BMPs within each DBM grade may not have as great an effect on the cells in vitro as it does in vivo. S ca f f old P r e p a r a tion a nd Optimiz a tion In addition to analyzing the effects of DBM grade on bone markers, we also wanted to test the efficacy of our scaffold. The method chosen for the construction of the scaffolds was the solvent casting/salt leaching technique utilizing polycaprolactone (PCL). Sieved sodium chloride (<100 um) particles were selected as the porogen. PCL (molecular weight 80 kD a) and the solvent, glacial acetic acid, were both purchased from
66 Sigma Aldrich. The size of the sodium chloride particles was confirmed using laser diffraction with a Beckman Coulter LS 1320 (Beckman Coulter Inc., Fullerton, CA). PCL pellets were dissolve d in glacial acetic acid at a concentration of 25% (w/v). This concentration was found to be best for dispensing the fairly viscous slurry suspension. PCL pellets and acetic acid were placed in a 250 ml beaker and stirred at 65C. Once the PCL was complete ly dissolved (1 2 hrs), sodium chloride (NaCl) particles at 90% w/w of PCL were added. 2 mls of the solution was then transferred to 12 well polystyrene culture plates (Corning Costar Corp.). The solution was allowed to air dry overnight at room temperatur e (23C). The resulting solidified scaffolds were washed in distilled water for 24 hours; with DI water exchanges every 6 hours, to remove residual solvent and remaining NaCl crystals. Scanning electron microscopy (SEM) images were taken of the resulting s caffolds (Figure 5 3). Results : The resulting scaffold was fairly porous (Fig 5 4 A). The micrographs, however, confirmed the presence of residual NaCl crystals in the scaffold after passive leaching using distilled water (Fig B). Figure 5 3. Scanning e lectron micrograph of PCL scaffold. Notice in (A) the surface of the scaffold is porous. However, when magnified, there appear to be residual sodium chloride crystal (B and C). The presence of sodium chloride can cause cell death [18 0 ].
67 The presence of Na Cl crystals is quite concerning. An excess of NaCl has been shown to cause cell death and hinder cellular proliferation. NaCl particle size analysis (Table 5 1) clearly indicates that a significant amount of NaCl crystals are less than 5.22 is a source of potential concern because the smaller NaCl crystals may get trapped into the slurry suspension and become difficult to extract. T a ble 5 1 D i str i buti o n o f NaCl particle size P ore Si z e ( m ) % o f P a r t i c l es 5 2 2 < 10 % 15 9 4 < 25 % 37 2 1 < 50 % 64 3 0 < 75 % 92 3 6 < 90 % Since an analysis of the scaffolds revealed salt retention after passive leaching using exchanges of distilled water, we decided to use other means of removing salt from the scaffolds. As mentioned earlier, the higher the concen tration of sodium chloride in the polymer solution, the more porous the scaffold. To test this theory, we reduced the c onc e ntrati o n by v olume o f NaCl in t he P CL sc a ffolds in inc r e m ents o f 1 0% by v olu m e 90%, 8 0 %, 7 0%, 6 0% a nd 50%. Th i s dec r ea s e in s alt c o n c entrati o n by v olu m e w as acc o mp an ied by an inc r ease in p ol y mer solution con c e n tration by v olu m e. In o t her w ords, a 10% reducti o n in v olume o f sodi u m chlor id e w as a ccom p ani e d by a 10% incr e ase in pol y m er ( PCL) soluti o n. T a b l e 5 2 s u m mar i z es the v olu m e o f P CL to NaCl in a 2 5% ( w /v) PCL soluti o n. The relationship between sodium chloride volume and porosity was analyzed using a Quantachrome Ultrapyc 1000 Gas Pycnometer.
68 T a b le 5 2 V o lume o f P CL to N a Cl in 25 % ( w / v ) PCL So lu t io n PC L conce nt r ati o n ( w / v %) V o l ume o f PC L t o N aC l ( w / w % (i.e. 90% PCL to 90% NaCl) 25% 90% 25% 80% 25% 70% 25% 60% 25% 50% The pycnometer measures the true density of a porous structure using gas displacement. The helium gas, which quickly penetrates small pores and is nonreact ive, is used to determine the solid portion of the sample. Porosity, for this purpose, is defined as a measure of the void volume contained within a scaffold. Scaffold porosity calculated by the pycnometer. Equation 5 1 is the formula used to calculate poro sity. Scaffolds (n=3) were cut to approximately 8 mm by 4 mm by 2 m m (l x w x h). Three apparent volume measurements and a single mass measurement were taken for each of the scaffolds. Scaffold mass was divided into average scaffold volume for each scaffold type % (w/w) of NaCl (i.e., 90% NaCl, 80% NaCl, etc.). (5 1)
69 In order to determine true density, the cut scaffold was placed in a sample chamber of known volume in the pycnometer. A reference chamber, also of known volume, was pressurized. The two chambers were then connected, allowing the helium gas to flow from the reference chamber to the sample chamber. The ratio of the initial and final pressures was used to determine the volum e of the sample solid. To calculate this volume, the pycnometer makes use of Boyles Law (Equation 5 2) where (P1) is the init ial pressure, (P2) is the final pressure, (V1) is the initial volume calculated using the equation 5 3 : In Equation 5 measured volume, calculated in equation 5 2 .We can then obtain the porosity of the scaffold using equation 5 1. Results : As expected, the porosity of the scaffold decreased as the volume of sodium chloride was reduced and that of the polymer solution was increased. However, a 90% (w/w) sodium chloride concentration only obtained 43% porosity. This is roughly half the porosity we expected. A 50% (w/w) sodium chloride concentration only obtained 20% porosity, while the other salt concentrations fell in between the 90% and 50%, though not in a straight line. (5 2) (5 3)
70 Figure 5 4. Porosity o f PCL scaffolds as a function of NaCl content. Shown in the figure are the theoretical (expected) and actual porosities of the scaffold as a function of NaCl. Notice that the actual porosity fell far below what we expected. Based on these results, we hypot hesize that the sodium chloride was entrapped within the scaffold. Thus, as a consequence, these trapped particles do not create a pore that would allow for the penetration of water during the leaching process and therefore would not remove the remaining N aCl crystals. Moreover, this reduces the overall porosity of the scaffold. Equally as important, this creates a high sodium chloride environment for the cells. This would validate the residual NaCl crystals we observed in the scanning electron micrographs In an attempt to increase the porosity and remove residual sodium chloride from the scaffolds, we modified the washing temperature and the leaching process. Sonication was included in the leaching process.
71 Was hing T e mp e r a t u re M odific a tion Temperature mod ification of the washing process was conducted in attempt to reduce salt retention. While the air drying process removes the solvent, the washing process removes the residual salt content. Since the standard washing temperature at 23C (Room Temperature) p roduced a scaffold with a high salt concentration, we decided to wash the scaffolds at temperatures of 37C, 45 C, and 55C. Three scaffolds of 90% NaCl (w/w) were each placed in a 37C, 45 C, and 55C in a 50 ml conical tube containing 25 mls of DI wate r. The scaffolds were incubated overnight using gentle rotation at 6 rpm in a Hybridization Incubator (Robbins Scientific, Model 400).After the washing period, the scaffolds were removed and NaCl content was analyzed using SEM with elemental Energy disper sive X ray spectroscopy (EDS). After being cut to approximately 8 mm by 4 mm by 2 mm ( l x w x h) (n=3), the porosity of the scaffold was analyzed using the gas pycnometer. The pycnometer measurements were taken three times for complete accuracy. While the pycnometer gave both density and volume measurements, only volume measurements were recorded.
72 Results A B c Figure 5 5. SEM micrographs of scaffolds washed at various temperatures. Scaffolds washed at 37 C (A) exhibit a slightly porous st ructure on their surface; scaffolds washed at 45 C (B) and 55 C (C) do not. In the former, there appears to be fault/crack lines along the surface. In the latter, the surface of the scaffolds appears to have re liquefied, causing a flatter and less porous surface.
73 Figure 5 6. Porosity of the PCL scaffolds as a function of washing temperature. The high temperature of 55C re liquefied the scaffold, closing the pores. Moreover, the pyncometer was unable to give an accurate and repeatable reading. UR= Unrea dable A g gr ess i v e L eac hing u s ing S oni c a tion Aggressive leaching of the sodium chloride from the scaffolds using sonication was conducted in an attempt to reduce salt retention and increase the interconnectivity of the pores. Three scaffolds of 90% (w/w) NaC l content were each placed in three separate 50 mL conical tubes containing 15 mLs of distilled water, with a ratio of sample surface to DI volume of 0.5 cm2/mL. The scaffolds were sonicated for 1, 3, 8 and 12 hours at 37C. After the sonication period, th e scaffolds were removed and residual salt content was analyzed using SEM with EDS. After being cut to approximately 8 mm by 4 mm by 2 mm (l x w x h) (n=3), the porosity of the scaffold was analyzed using the gas pycnometer
74 Results An analysis by SEM wit h EDS revealed reduced salt retention in all samples when compared with scaffolds washed at 23C. In particular, no salt particles were observed on the surface of the scaffold. Also, the sonicated scaffolds exhibited an interconnectivity not previously enc ountered. Figure 5 7. SEM micrographs of scaffolds leached at various sonication times. Scaffolds sonicated for 1 hour (A) exhibit a porous structure on its surface. These pores were circular in nature and appeared to be widespread across the surface of the scaffold. Scaffolds sonicated for 3 hours (B) revealed a highly porous to was also a n increased level of interconnectivity between the pores. Scaffolds sonicated for 8 hours (C ) displayed a porous structure, but one that lacked depth. Perhaps the increasing sonication time and the consequent slight increase in temperature caused the surface to re liquefy and close in on the pores.
75 Figure 5 8. Porosity of the PCL scaffolds as a function of sonication. The longer sonication time of 8 Hours re liquefied the scaffold, closing the pores. Moreover, the pyncometer was unable to give an accurate and repeatable reading. UR= Unreadable Scaffolds sonicated for 1 and 3 hours exhibited a high porosity of 89% and 90% respectively. Moreover, the sizes of the pores in these sonicated scaffolds ranged source of concern is that the smaller sized salt particles may get trapped in the viscous polymer and hence would not be able to be extracted. Removal of salt particles by ultrasonic leaching eliminates a potential cause of cell death. The highly porous and interconnected scaffolds produced by sonication for 3 hou rs would make an ideal scaffold because this would allow cell nutrients to enter and remove waste from the cells.
76 Sonication for 8 hours caused the scaffolds to re liquefy and become quite gummy. The pores produced by the scaffolds sonicated for 8 hours l acked the depth seen in the scaffolds sonicated for 1 and 3 hours. Similar to the scaffolds washed at 55C, we were unable to obtain the porosity of the scaffold sonicated for 8 hours. Compared to the scaffolds sonicated for 1 and 3 hours, the scaffolds so nicated for 8 hours exhibited higher peaks on the EDS spectrum for both sodium and chlorine. We believe that the longer sonication time increased the temperature of the water bath, which in turn increased the temperature of the distilled water the scaffold s were immersed in. Furthermore, this increase in temperature re liquefied the scaffold causing the pores to close. The water bath temperature was measured with a non mercury thermometer and verified with the digital temperature display of the water bath. Di sc us s ion Visually, SEM micrographs confirmed the absence of sodium chloride particles within scaffolds sonicated for 1 and 3 hours and those washed at 37C and 45C. The EDS spectra confirmed the absence or reduction of sodium chloride particles. The con firmation by EDS is perhaps more important than the SEM micrographs. EDS is a technique that is used for identifying the elemental composition of a particular area of interest on the scaffold. In the EDS analysis, the scaffold was bombarded with an electro n beam inside the SEM microscope. These electrons collide with the scaffold electron occupies the position ejected by the inner shell electron. However, in order to exe cute this, the higher energy electron must give up some of its energy by emitting an X ray. The amount of energy present in the X rays is then used to identify the atom from which the X ray was emitted. This is then translated into an EDS spectrum composed of
77 peaks, which plots how frequently an X ray is received for each energy level. Each peak corresponds to a unique atom and, hence, to a single element. The more concentrated an element is in a scaffold, the higher its peak will be in the spectrum. Based on these results, the 90% w/w NaCl scaffolds sonicated for 3 hours at 37C present the ideal characteristics sought in an ideal scaffold high porosity and interconnectivity between the pores.
78 C H AP T ER 6 CO M PO S I T E S CAF F O L DS: S Y N T H E SI S DE G RAD A TI ON A ND C A L CI U M PHOSPH A T E D E PO S I T I O N Intro d ucto r y R e m a r k s We previously developed a three dimensional interconnected porous PCL scaffold. The thrive and proliferate at a pore size range of 380 n microscopy (SEM) with energy dispersive spectroscopy (EDS), we confirmed the removal of the porogen, sodium chloride. The removal of sodium chloride eliminates a potential cause for cell death. Our overarching goal in this project is to develop a composi te scaffold of polycaprolactone polymer and demineralized bone matrix particles. We hypothesize that the design of this biodegradable polymeric scaffold that incorporates DBM particles will enhance the efficacy of an otherwise purely polymeric scaffold. By enhance, we mean that the incorporated DBM will add nucleation sites for the deposition of calcium phosphate and provide signals for the differentiation of pluripotent cells. Specifically, the goal of this study was to examine both the degradation charact eristics and the deposition of calcium phosphate on the composite scaffold. We assessed the degradation solution. To evaluate whether DBM increases the deposition o f calcium phosphate on the surface of the scaffold, we compared a pure polycaprolactone scaffold to a composite scaffold of 25 % DBM (w/w), 35 % DBM (w/w), and 50 % DBM (w/w). Henceforth, these scaffolds will be denoted as low, mid, and high composite scaf fold respectively. These scaffolds were assessed using the following parameters: calcium ion release into the degradation solution, qualitative assessment of calcium phosphate deposition through SEM w/EDS, and a Von Kossa staining.
79 M a ter ia ls a nd M e t h o d s S c a f f old P r e p a r a tion PCL (molecular weight 80 kDa) and the solvent, glacial acetic acid, were sourced commercially (Sigma erously donated by Regeneration Technologies (RTIX, Alachua, FL).The sizes of the DBM and sodium chloride particles were confirmed using laser diffraction with a Beckman Coulter LS 1320 (Beckman Coulter Inc., Fullerton, CA). PCL/DBM composite scaffolds wer e formed by using a solvent casting/porogen leaching technique. Initially, PCL pellets were dissolved in glacial acetic acid at a concentration of 25% (w/v) in a beaker and stirred at 65C. Once the PCL was completely dissolved (1 2 hrs), sodium chloride (NaCl) particles at 90% w/w of PCL were added. The polymer solution was allowed to cool at room temperature for 30 minutes. Once the temperature of the polymer solution reached 37 C, the polymer solution was cast into well culture microplates (Corning CoS tar Corp.). Demineralized bone matrix was then added to the resulting scaffolds. Since the proteins within DBM are deactivated at high temperatures (65 C and higher), it is important to allow the polymer suspension to cool . Four types of scaffolds were formed: pure polymeric scaffolds, PCL/DBM scaffolds that contained approximately 25% DBM (w/w), PCL/DBM scaffolds that contained approximately 35% DBM (w/w), and PCL/DBM scaffolds that contained approximately 50% DBM (w/w). After the evaporation of th e solvent at 23C (RT) overnight, the resulting solidified scaffolds were placed in 50 mL conical tubes containing 15 mLs of distilled water, with a ratio of sample surface to DI volume of 0.5 cm2/mL. The scaffolds were sonicated for 3 hours at 37C. After they were freeze dried overnight. The overall process is shown in Fig 6 2.
80 Fi g u re 6 1 S EM M i c r o g ra ph s o f De m i ne ral i z e d Bon e Ma t r i x Fi g ur e (A) i l lus t rat e s t h e i r re g u lar s hap e o f t h e pa rtic l e s, w h ich c o n trib u t e s t o t h e resi d ua l m i ne ral ( < 5 % c a lc i u m a n d p h o s p h o r ou s) l e f t i n t h e p a rtic l e T h e m e a n si z e o f t h e pa rtic l e s w a s 68 7 m Fi g u re B i l lus t rat e s t h e s u r f a ce o f e a ch DBM pa rtic l e F igure 6 2. Schematic of solvent casting/ salt leaching technique. The porogen is added to a 25% PCL solution. Once the temperature of t he polymer solution cools to 37C the polymer solution is cast and the polymer solution begins to solidify. Before the solution solidifies, and while it is a slurry like paste, DBM particles are added. DBM is added at this time for two reasons: first, the proteins in DBM are denatured at temperatures above 65 C and second, adding the particles while the solution is still a slurry like paste, allows for uniform distribution of DBM particles throughout the scaffo ld. Scaffolds are then freeze dried to remove residual solvent.
81 Degr a dation Media Scaffolds were incubated in 15 ml of simulated body fluid (SBF) that was maintained at 37C and changed every three days for up to 7 weeks. The sample surface to DI volume w available during the entire test). SBF was prepared according to Kokubo et al.  and contained the following ions and their respective concentrations (Table 6 1 ): Table 6 1. Ion concentration of Human Plasma and SBF according to Kokubo. I o n Hu m a n P la s m a ( m mo l / L ) Si mu l a t e d B o d y Fluid ( mmo l / L ) Na + 142 0 142 0 K+ 5 .0 5 .0 Ca 2 + 2 .5 2 .5 M g 2 + 1 .5 1 .5 Cl 103 3 148 8 H C O 3 27 .0 4 .2 HPO 42 1 .0 1 .0 SO 42 0 .5 0 .5 S ca f f old We i ght Ch a nge Scaffolds were weighed once prior to the degradation study and then every three days for 7 weeks. At each time point, scaffolds were removed from the SBF, rinsed with DI water and blotted dry with a Kimwipe till no liquid was seen on the wipe. Then, the scaffolds were freeze dried for 24 hours to ensure complete removal of water, and weighed. Six samples were weighed for each group (25% DBM, 35% DBM, 50% DBM and pure PCL). Calcium Ion Release The SBF solution in which the s caffolds were immersed was analyzed for calcium ion content at the following time points: 5 days, 10 days, 15 days and 20 days. SBF was analyzed for calcium ion concentration using a colorimetric calcium ion kit (No. 0150, Stanbio Laboratories). The manufa
82 to 1 ml of reagent (0.5 ml Color reagent and 0.5 ml Base reagent). Six samples of each type of scaffold were analy zed for calcium ion concentration. The SBF was then analyzed for optical density using a spectrophotometer (Shimadzu, Japan) at a wavelength of 550 nm. Per manufacturer, optical density was normalized to a known calcium ion concentration and reported as mg /dl. SBF, not exposed to scaffold, was used as the control. Von K o ss a Stain i ng Von Kossa staining was used to detect mineral deposition on the scaffold. Briefly, the scaffolds were washed three times with Phosphate Buffered Saline (PBS), fixed in 3.7% form aldehyde in PBS for 10 minutes, washed again in distilled (DI) water three times, incubated in .5% silver nitrate solution (Sigma Aldrich) for 1 hour. Upon removal, the scaffolds were washed in DI water three times, incubated in 0.3% sodium thiosulfate pen tahydrate (Sigma Aldrich) solution for 3 minutes and, finally, air dried. The stained scaffolds were then observed under light microscopy. Phosphorous Ion Release The SBF solution in which the scaffolds were immersed was analyzed for phosphorous ion conten t at the following time points: 5 days, 10 days, 15 days and 20 days. SBF was analyzed for phosphorous ion concentration using a colorimetric phosphorous ion kit (No. 0830, Stanbio Laboratories). The reaction to determine phosphorous content is based on th e reaction of the phosphate ion with a molybdate compound. This reaction results in a reduced colorless hexavalent molybdenum phosphate complex. Per the unreduced compl type and added to 1 ml of reagent. Six samples of each type of scaffold were analyzed for
83 phosphorous ion con centration. The SBF was then analyzed for optical density using a spectrophotometer (Shimadzu, Japan) at a wavelength of 340 nm. Per manufacturer, optical density was normalized to a known phosphorous ion concentration and reported as mg/dl. SBF, not expos ed to scaffold, was used as the control. pH of S imul a ted Bo d y Flu i d The pH of the SBF in which the liquids were immersed were analyzed at the following time points: 1 hour, 4 hours, 24 hours, 5 days, 15 days, 25 days, 35 days and 45 days. The SBF was analy zed using both pH test strips and a pH meter (Corning, LS 430). Again, six samples from each scaffold type were analyzed. Scanning Electron Microscopy with Energy Dispersive Spectroscopy Calcium phosphate precipitation was qualitatively evaluated at 14 day s post immersion using scanning electron micrographs (JEOL JSM 6335F) at an acceleration voltage of 15kV. The samples were coated with gold prior to analysis. Statistics Six samples were analyzed for each scaffold type at the various time points for mass c hange, calcium ion release, phosphorous ion release and pH change. Three samples from each scaffold type were analyzed for Von Kossa staining. All statistical analysis was conducted using the GraphPad Instat program. The data within each scaffold type was examined between time points and analyzed using a one way analysis of variance (ANOVA) with statistical significance at p<0.05. Further, data comparing the scaffolds was examined using two way analysis of variance (ANOVA) with statistical significance at p<0.05. Post hoc analysis was conducted using the Tukey test. The values were expressed as the mean standard error (n=6).
84 Re s ults M as s of S ca f f olds The mass of the pure PCL scaffolds were remarkably stable over the course of the 7 week experiment. As p reviously mentioned, PCL is a hydrophobic polymer and its susceptibility to hydrolytic degradation is diminished by this property. On the contrary, scaffolds containing DBM exhibited a significant mass loss over the duration of the experiment. 25% DBM (Low ) scaffolds mass loss tracked closely with pure PCL scaffolds, but statistical significance was observed between the 20 and 25 day mark. Both 35% DBM (Med) and 50% DBM (High) experienced a statistically significant mass loss in the first five days. In the first five days the scaffolds were immersed in SBF, 35% DBM exhibited a 4% drop in mass while 50% DBM exhibited 18% drop in mass. 35% DBM exhibited a slight increase in mass between 15 and 20 days. This course was slightly reversed later, beginning at day 25 and then throughout the experiment. 50% DBM exhibited another significant drop between the 15 and 20 day time points. This was followed by a statistically significant increase between day 45 and 50. From this data, it appears that the percent content o f DBM in the scaffolds greatly mattered in evaluating mass loss. Scaffolds with a higher DBM content experienced far more mass loss than those with little or none. This was expected for two reasons. The DBM is only physically blended with the PCL and, thus ly, occupies random space within the scaffold. Secondly, DBM is hydrophilic. Therefore, its interaction with the immersion liquid could increase the chances of its dislodgement from the scaffold.
85 Figure 6 3. Mass loss of scaffolds over 7 weeks. Each ba r represents the average of six scaffolds and the error bars denote SEM. Time point 0 is the starting mass prior to immersion in SBF. Scaffolds that contained 35% and 50% DBM showed greater reductions in mass over time compared to pure PCL scaffolds and sc affolds composed of 25% DBM. 25% DBM and pure PCL scaffolds tracked closely in terms of degradation over time.
86 Calcium Ion Release The concentration of free calcium ions in the SBF was largely dependent on scaffold composition and to a certain and, perh aps, lesser degree, the duration of immersion time. Scaffolds composed of 50% DBM showed a statistically significant burst of calcium ion concentration between Day 5 and Day 10. It is important to note that between the time of Day 5 and Day 10, 50% DBM un derwent a significant mass loss. Although there was a slight uptick between Day 5 and Day 10 for both the low (25% DBM) and med (35 % DBM) scaffold, no statistically significant increase or decrease of calcium ion was found at any time point on these scaff olds. The free calcium ion concentration of pure PCL scaffolds was consistent throughout the experiment. Obviously, the PCL scaffold contains no DBM, so an increase in free calcium ion was not expected. Yet, the lack of a decrease in calcium ion suggests v ery little calcium phosphate precipitation on the scaffold. DBM particles incubated in SBF were also assessed for calcium ion concentration. After an initial 5 days, the calcium concentration of the DBM had increased from 10 mg/dl (the calcium concentratio n of the SBF without any scaffold or particles) to 12.98 mg/dl. Furthermore, on Day 15, the concentration of free calcium ions in the simulated body fluid from which the DBM was immersed had decreased to 11.46 mg/dl. This suggests an initial dissolution of residual calcium ions from the DBM, followed by calcium phosphate reprecipitation on the surface of the DBM
87 Figure 6 4 Calcium ion concentration in simulated body fluid containing 25 % DBM scaffolds, 35% DBM scaffolds, 50% DBM scaffolds pure PCL scaf folds and DBM particles. Fi g u re A r e p res e n ts d a ta f o r 25% DBM scaffolds Figure B represents data f o r 35% DBM scaffolds, Figure C represents data for 50% DBM scaffolds Fi g u re D r e p res e n ts d a ta f o r pure DBM particles Notice that the free calcium ion con centration of 25%, 35% and pure PCL scaffolds stayed stable throughout the immersion period. Scaffolds composed of 50% DBM had an initial increased of calcium ions, as compared to the concentration of control SBF, at the first time point of 5 days. This co ncentration gradually decreased by Day 15. The concentration of free calcium ions in control SBF was 10 mg/dl.
88 Von K o ss a Stain i ng Von Kossa staining confirmed that calcium salts were deposited on the scaffolds after immersion in SBF at 14 days (Figure 6 5) The mineral deposits were observed in all scaffolds containing DBM. Figure 6 5. Von Kossa staining of 25 % DBM scaffolds, 35% DBM scaffolds 50% DBM scaffold and pure PCL scaffolds. Clockwise, (A) 25% DBM Scaffolds 6 5 (B) 35% DBM Scaf folds, (C) 50% DB M Scaffolds 6 5 (D) Pure PCL Scaffolds. Mineralized Deposits are indicated by black arrows. Magnification at 40X.
89 We observed no mineral deposits on the pure PCL scaffolds. We did, however, observe a correlation between DBM content and mineral deposition on the scaffold. Deposition of mineral on the scaffold became visually more pronounced as the volume of DBM increased in the scaffold. Phosphorous Ion C o nc e n t r a tion The concentration of phosphorous (P5+) ions in the SBF was not as dependent on scaffold com position as the concentration of calcium (Ca2+) ions. At Day 5, the phosphorous concentration of the SBF incubated with 25% DBM, 35% DBM and pure PCL scaffolds were not statistically different from the control SBF at 4.33 mg/dl. However, 50% DBM scaffold w as. Moreover, between time points of Day 5 and Day 10, the SBF of the 35% and 50% DBM scaffold exhibited a slight increase in phosphorous ion concentration. Meanwhile, the SBF of the 25% DBM scaffold and pure PCL scaffold exhibited a consistent level of ph osphorous concentration throughout the immersion period. The parallel fluctuations of calcium and phosphorous ion concentration of the 25% and pure PCL scaffold are illustrated in figures 6 6 (A, B). Figure 6 6. Calcium and phosphorous ion concentratio ns of SBF of pure PCL scaffolds. (A) Calcium and phosphorous ion concentrations of SBF of 25% DBM (B) scaffolds. The concentrations of calcium and phosphorous ions of both scaffolds track closely to each other
90 Fi g u re 6 7 P ho s p h o ro u s i o n c on c en tra t i o n in si mu l a t e d b o d y f luid c on t a i n ing 2 5 % DBM sc a f f o ld s 3 5 % D BM sc a f f o l d s, 5 0 % D B M sc a f fo lds, pu r e P CL sc a f f o ld s an d DBM pa rtic l e s. Fi g u re A r e p res e n ts d a ta f o r 25% DBM scaffolds, Fi g u re B r e p res e n ts d a ta f o r 35% DBM scaffolds, Fi g u re C r e p res e n ts d a ta f o r 50% DBM scaffolds, Fi g u re D r e p res e n ts d a ta f o r pure DBM particles and Fi g u re E r e p res e n ts d a ta f o r pure PCL scaffolds. N o tice t h a t pho s pho r o u s ion c o n c en t r a ti o n o f 2 5 % an d pu re P CL s c a f fo lds st a y e d s t ab le th r o u g hou t t h e i m m e rs i o n pe r i od Sc a f f o l d s c o m p o s e d o f 35 % an d 50 % DBM h a d a sli g h t an d s t e a d y i n c r ea se o f pho s pho r ou s i o n c o n c en tra t i o n a t b o th D a y 5 an d D a y 10 A t Day 15 t h e c o n c en tra t i o n h a d re v e rs e d c ou rse a n d d e c r ea s e d T h e c on c e n tra t i o n o f p h o s p ho ro u s i on s o f t h e c on t rol S BF (S B F w it hou t a n y sc a f f o ld o r p a rtic l e s im me rs e d ) w a s 4 3 3 m g / d l.
91 The uptick in phosphorous ion concentration seen at Day 10 for 35% and 50 % DBM scaffolds could be a result of dissolution of residual phosphorous ions from the DBM. This uptick is followed by a statistica lly significant decrease at Day 15. This suggests calcium phosphate deposition on the surface of the scaffolds. Once again, the lack of a decrease in phosphorous ion suggests very little calcium phosphate precipitation on the pure PCL scaffold. DBM particl es incubated in SBF were also assessed for phosphorous ion concentration. After an initial increase in phosphorous ion concentration, as compared to the control SBF (4.33 mg/dl), there was a significant decrease in phosphorous ion concentration at Day 10. However, by Day 15, the ion concentration had slightly increased. This is unlike calcium ion concentration, in which we saw a slight decrease from Day 10 to Day 15. Figures 6 7 (A E) illustrate the fluctuations in the phosphorous concentrations of the scaf folds. Notice that the PCL stayed stagnant throughout the entire 15 day experiment. pH of Simulated Body Fluid The pH of the simulated body fluid containing 25% DBM scaffolds, 35% DBM scaffolds, 50% DBM scaffolds, and pure PCL scaffolds were reported as de viations from control SBF which contained no scaffold (pH 7.1). Scaffolds containing DBM showed the greatest deviations from control SBF. In particular, scaffolds containing 50% DBM. Pure PCL scaffolds, and to a lesser extent, 25% DBM scaffolds demonstrate d slight fluctuation in pH from Day 0 to Day 50. Scaffolds containing 50% DBM exhibited a sharp drop in pH in the first five days. This was followed by a moderate decrease in pH from Day 10 to Day 20. At Day 30, a slight increase in pH was observed. The pH of the 50% DBM remained stable until Day 50. Figure 6 8 illustrates the change in pH of the SBF over time.
92 Figure 6 8. pH of immersion liquid (SBF) for the various scaffolds. Scanning Electron Microcopy w/ Elemental Dispersive Spectroscopy Fi g u re 6 9 S c ann i n g e lec t ron m ic r o g ra ph s o f (A) pu re po l y me r i c P CL sc a f fo lds, sc a f f o lds w ith 25 % D B M c on t e n t ( C ), 3 5 % D BM po st 1 4 da y s i m me rs i o n (B), an d 50 % DBM s c a f fo l d s ( D ). Pu re P CL s c a f fo lds s ho w e d n o e v id en c e o f p recipit a ti o n w h i l e th o se c o n t a ini n g DBM c o n t a i n e d v a r i ou s a m o u n ts o f c a lc i u m pho s p h a te p r e cipit a ti o n
93 SEM analysis indicated very little Calcium and Phosphorous deposition on the surface of the PCL scaffold (Figure 6 9 A). Calcium and Phosphorous deposition on the scaffolds steadily increased as DBM content increased. Noticeably, there exists very little difference in the signal lines for Calcium and Phosphorous when comparing 35 % DBM (Figure 6 9 C) and 50% DBM (Figure 6 9 D). Di sc us s ion The purpose of this chapter is to explore the physical and c hemical characteristics of the various scaffolds as well as analyze the deposition of a calcium phosphate layer on its surface. In this chapter, we explore properties such as rate of degradation and pH. The degradation of a scaffold is perhaps the most imp ortant feature in designing a match the properties of the tissue it was restoring. An ideal scaffold would provide a biochemical and mechanical support system until th e damaged tissue was fully regenerated. An ideal scaffold would completely biodegrade at a rate consistent and in line with tissue regeneration. This prevents a pseudo healing, in which the scaffold has completely degraded but the tissue generated is not y et at full functionality. Furthermore, scaffold degradation that is slower than tissue generation could inhibit or disrupt new tissue functionality. Equally as important to the success of an orthopedic scaffold is a layer of calcium phosphate mineralizatio n on its surface. The coating of calcium phosphate on an implant has been shown to increase its ability to bond in vivo with native bone . The success of an orthopedic implant is inherently linked to its ability to bond to native bone in vivo. This bo nding promotes osseointegration, the formation of a direct interface between an implant and bone, without intervening soft tissue [183,184]. The
94 encapsulation of orthopedic implants by fibrous tissue is well documented [185 188] and is one of the leading c auses of implant failure. Implant loosening is an indirect result usually due to mechanical or wear debris causing fibrous tissue to isolate the implant from the native bone. While surface characteristics such as surface roughness, surface chemistry and to pography play an important role in developing good osseointegration, so too does a surface layer of calcium phosphate mineralization. While DBM is by definition mineral free, it does contain less than 5% mineral. We believe that the residual calcium and ph osphorous ions leech out of DBM and reprecipitate on the surface of the DBM to create new mineral. In essence, DBM is able to remineralize. This is not unlike how bone is formed. When the blood stream lacks calcium, the parathyroid gland is stimulated to r elease parathyroid hormone (PTH). The PTH then signals osteoclasts to break down the bone tissue and subsequently, release calcium salts into the bloodstream. Afterward, the osteoblasts form new extracellular matrix using the calcium salts released into th e bloodstream by the osteoclasts. We can liken this sequence to builders demolishing a house and retrieving its raw materials to build a new house. This reprecipitation of calcium phosphate on the scaffold allows for bonding and integration of the scaffold and native bone. However, there are varying theories on how bone (mineralized or not) reprecipitates. It is hypothesized that exposed phosphate (PO43 ) and hydroxyl groups ( OH) from the mineral layer elicit a negative charge [189 191] on the surface of t he material. This attracts the positively charged calcium ions, leeched from the bone, to the surface of the bone. This then attracts the negatively charged phosphate ions, which results in reprecipitation on the surface. This is known as heterogeneous nuc leation. A
95 cartoon diagram is depicted in Figure 6 10. Obviously, immersing DBM in simulated body fluid (SBF) rich with calcium and phosphorous ions would increase the probability of mineral precipitation on the surface of the scaffold. Yet, it is not corr ect to assume that immersing a material in SBF would initiate a mineralization on its surface. osseointegration of the implant and native bone is well documented. Therefore, it is e vident that the ability of a scaffold to form a calcium phosphate layer on its surface is paramount to developing an effective and functional scaffold in vivo. Figure 6 10 Heterogeneous nu cl e a ti o n o f c a lc i u m pho s pha t e o n de m in e r a l i z e d bon e ma t r i x Fi g u re 6 10 A dep icts t h e rel e a se o f c a lc iu m an d pho s p h a te f r o m b on e (DBM is u s e d i n th is f i g u re). Fi g u re 6 10 B il l u str a t e s th e po siti v e ly c ha r g e d c a lc iu m be ing a t t ra c t e d t o t h e n e g a ti v e l y c ha r g e d h y d r o x y l g ro u p (loc a t e d o n t h e apa ti t e ) a n d t he n t h e n e g a ti v e l y c ha r g e d ph o s pha t e g ro u p be ing a tt ract e d t o c a lc i u m (Gro w th o f a p a tit e ). T h e s a t u ra t e d S BF incre a s e s t h e g ro w t h o f apa ti t e o n t h e s u r f a c e For th e pu r p o se o f t h is d ia g ra m t h e DBM w a s re m o v e d f r o m t h e po l y me r s u b stra t e a s to cl e a r l y s ho w th e d issol u t i o n o f t h e io n s. In light of this consideration, it would seem that the greater the amount of DBM a scaffold contains, the greater the potential for calcium phosphate mineralization. To examine this, scaffolds composed of 25% (w/w) DBM, 35% (w/w) DBM, 50% (w/w ) DBM, and pure polycaprolactone were analyzed for mineralization.
96 In our degradation studies, the scaffolds were immersed in SBF solution prepared according to Kokubo. The ionic concentration of the SBF was similar to that of human plasma. Moreover, SBF w as used to assess bioactivity namely, calcium phosphate deposition on the surface of the scaffolds. In particular, we investigated whether phosphate mineralization on its surface. The mass of all scaffold types declined over the 7 week experiment; however, some declined more than others. The mass of the 50% DBM declined by nearly 43%, while those of the pure PCL scaffolds only declined by 7%. Since the DBM was only physica lly blended into the scaffold, we expected a decline in scaffold mass. PCL is hydrophobic and, therefore, its surface precludes liquid penetration into the scaffold. Significant mass loss was experienced in the composite scaffolds due to the interaction be tween the hydrophilic DBM and the SBF. During the period between Days 15, 20, 30, and 35, 35% DBM exhibited a slight increase in mass. This possibly could have been due to reprecipitation on its surface. However, the mass loss of the composite scaffolds ma kes it fairly difficult to ascertain mass gained as a consequence of mineral deposition. The pH of the SBF solution for the composite scaffolds dipped slightly through the 7 week experiment. The SBF solution that contained pure PCL scaffolds remained stabl e throughout. The fluctuations of the pH of the composite scaffolds can be explained in a number of ways. It is entirely possible that residual acidic solvent remained in the scaffold. We were able to achieve a porosity of nearly 90% with the pure PCL scaf folds. However, the larger sized DBM particles made it difficult to obtain
97 such a high porosity. The porosity of the composite scaffolds was as follows: 37% for the 50% DBM, 51% for the 35% DBM, and 56% for the 25% DBM. Pulverizing the DBM to a finer size would have both increased the porosity of the composite scaffold and allowed the penetration of water to remove any residual solvent from within the scaffold. For the purpose of this research, we decided to use DBM particles that provided the most osteoinductive potential, those within a range of 510 to 710 microns. It is also calcium, phosphate, and hydr oxyl ions from the DBM within the scaffold. At Day 10, the calcium ion concentration for the SBF that held 35 and 50% DBM scaffolds as well as pure DBM particles was higher than that of the control SBF, 10 mg/dl. The calcium concentration for 50% DBM scaf folds and pure DBM particles was significantly higher. This suggests that residual calcium ions were released from the DBM at, or at least up until, Day 10. Moreover, it was in these first 10 days that we saw a significant amount of mass loss associated wi th scaffolds containing 35 % and 50% DBM. However, by Day 15, the concentration of calcium ions in the composite scaffolds and pure DBM had decreased and returned to the calcium concentration of control SBF. It was also interesting to note that, at Day 15, the mass loss of the composite scaffolds had begun to subside, save for a temporary drop in the 50% DBM at Day 20. It is plausible that as the calcium ions were released, they proceeded to te layer by binding with the also leeched phosphate This could also explain the small mass increase we saw in the composite scaffolds. This is seen in both the Von Kossa Staining [Figure 6 5]
98 and SEM imaging [Figure 6 9]. It is particularly pronounced in the 35% and 50% DBM scaffolds. Alongside its binding with phosphate ions, the calcium ions could have been binding with hydroxyl ions to form CaOH2 as a precipitate. This would have decreased the amount of OH ions in the SBF and would have resulted in a net increase of H+ ions and a decrease in pH. In the same respect, we noticed an increase in phosphorous ion concentration, as compared to control SBF, starting at, or lasting until, Day 10. Furthermore, by Day 15, the concentration had returned to that of control SBF, 4.33 mg/dl. Again, the availability and leeching of both calcium and phosphorous ions make reprecipitation on the scaffolds plausible. And while both 35% and 50% DBM eluted a significant amount of calcium and phosphorous ions, based on the SE M images, we did not detect a significant difference in mineralization on the 35% and 50% DBM scaffolds. There was, however, a lack of mineralization on both the PCL polymer and the 25% DBM scaffold. The dip in pH lasted for roughly two weeks for the compo site scaffolds, namely the 35 and 50%. Notwithstanding the temporary blip at Day 20 for the 35% DBM scaffold, the pH steadied and began to increase thereafter. Both the decrease and increase can be explained by considering the fact that the mass loss follo wed an identical trend. As mentioned, the decrease in pH can be explained by the dissolution of the residual calcium phosphate within the DBM. Notice that the scaffold that produced the highest increase in calcium and phosphorous ion concentration was the 50% DBM scaffold. Likewise, 50% DBM scaffolds also had the most mass loss over the 7 week experiment. While both the calcium and phosphorous ion concentration data illustrated a difference amongst the various composite scaffolds, the SEM images were not as
99 expressive. Consider, for instance, the EDS analysis of the 35% and 50% DBM scaffold. While the signal was higher for both calcium and phosphorous for 50% DBM scaffold, it was not significantly different from that of the 35% DBM scaffold. One theory that could explain this occurrence is the pH drop found in the SBF solution of the 50% DBM scaffold. It is known that calcium phosphate degrades at an acidic pH. It is possible that the dip in pH caused dissolution of synthesized calcium phosphate. In other w ords, as quickly as calcium phosphate reprecipitated on the 50% DBM scaffold, it began to dissolve due to the low pH. This cycle continued until at least Day 15, when the pH steadied and began to rise. Since the SEM images were only taken at Day 15, it is possible that calcium phosphate began to reprecipitate at a faster clip than that of 35% DBM scaffold after this time point. This is a finding that requires further analysis. As mentioned, post Day 15, the pH of the composite scaffolds began to rise. The r eprecipitation of calcium phosphate on the surface of the scaffold would have had an increasing effect on the pH. Less calcium available would have meant less calcium hydroxide formed and more OH ions in the SBF solution. Since there are more hydroxide io ns than hydrogen ions present, the pH of the solution would increase which we saw in our experiment. The formation of a calcium phosphate layer on an implant is integral to its bonding to native bone in vivo, and thus its osseointegration. Osseointegrati on is defined as the formation of a direct interface between an implant and bone, without intervening soft tissue. The purpose of this study was to analyze whether the amount of DBM contained in a scaffold was proportional to the amount of mineralization d eposited on its surface. The results proved to be mixed.
100 In summation, we set out to develop a composite scaffold of demineralized bone matrix and polycaprolactone that supports and promotes mineralization. We developed three types of composite scaffolds a nd a pure polymeric scaffold. The composite scaffolds contained 25% DBM (w/w), 35% DBM (w/w) and 50% DBM (w/w). We developed these composite scaffolds to determine if mineral reprecipitation was proportional to percent volume of DBM in a scaffold. We repor t the results as mixed because, even though the concentration of calcium and phosphorous ions increased in the SBF solution of the composite scaffolds that contained a higher volume of DBM with respect to PCL in the scaffold, mineralization on the surface of the scaffold was not proportional to DBM content. From this data, we can conclude that both the 35% and 50% DBM scaffolds supported and promoted calcium phosphate mineralization on its surface. Moreover, since no significant difference in mineralizatio n was detected on the surface of the 35 and 50% DBM scaffold, we believe that continuing with the 35% DBM scaffold for this project is appropriate.
101 CHAPTER 7 IN VITRO EVALUATION OF THE EFFECTS OF CA LCIUM SULFATE, HYDROXYAPATITE, AND DEMINERALIZED BONE M AT RIX ADDITION TO POLYCAPROLACTONE SCA FFOLDS Introductory Remarks Bone is a dynamic tissue. An effective bone substitute would need to possess both the requisite biological and mechanical properties of bone. Given that, bone substitutes are rarely single mat erial products. Rather, they usually combine a biodegradable polymer with an additive that provides a biological and/or mechanical supplement. There has been a lot of interest in using ceramic based materials for tissue engineering, mostly due to their hig h mechanical strength or their similarities with bone. Calcium sulfate and calcium phosphate, or their derivatives, are ceramics that have been heavily explored for use in bone tissue engineering. The biological rationale for using calcium phosphate is tha t hydroxyapatite (HA), a subgroup under calcium phosphate, accounts for 70% of bone . An inorganic mineral, hydroxyapatite, is mostly crystalline but is quite brittle. Owing to this, it is mostly used as coating for heavy duty materials such as tita nium alloy. This is typically done by plasma spraying, which produces predominantly amorphous HA coating . Calcium phosphate crystallizes in a number of crystal structures that possess different mechanical, thermal, and chemical stabilities [ 193 ]. Ce rtain structures include amorphous calc ium phosphate (ACP), brushite tricalcium phosphate ( TCP ), octacalcium phosphate and finally, hydroxyapatite ( HA ). Amorphous calcium phosphate is widely associated with being a metastable precursor for the formation of the more thermodynamically stable and advanced stages of calcium phosphate, octacalcium
102 phosphate and hydroxyapatite. The formation for the various subgroups of calcium phosphate is highly dependent on the pH value of the immersion solution. The process from amorphous calcium phosphate to the more stable hydroxyapatite reprecipitation and growth on HAP. A cartoon diagram depicting the release of calcium and phosphate ions from the les s crystalline segments of Hydroxyapatite in simulated body fluid is shown in Figure 7 1 Figure 7 1. Dissolution and Reprecipitation of amorphous calcium phosphate on Hydroxyapatite The orientation of the hydroxyl and phosphate groups in hydroxyapati te lends itself to a negatively charged surface. The negatively charged surface attracts the positively charged calcium ions which then attract the negatively charged phosphate group, making an amorphous calcium phosphate on the layer of the hydroxyapatite The formation of the calcium poor amorphous calcium phosphate increases as the soaking time in SBF increases. Calcium sulfate hemihydrate (CSH), more commonly known as Plaster of Paris, has been used as a bone void filler for well over 2 centuries . It has gained wide acceptance as a viable substitute due to its biocompatibility, bioactivity and resorbability. Mixed with water, CSH reacts to form calcium sulfate dihydrate, which is HA
103 also known as gypsum. This paste, also known as bone cement, is widel y used due to its ability to conform to the size and shape of the defect and its lack of inflammatory response in vivo. suffered open fracture wounds in battle involves the use of Os teoset (Wright Medical), a calcium sulfate based product. However, under closer scrutiny, calcium sulfate has shown a faster resorption rate than reported, and had a less than appealing healing in vivo [148 150]. The manufacturer, Wright Medical, has curr ently started incorporating demineralized bone matrix (DBM) into its flagship calcium sulfate based product, Osteoset In Osteoset 2 DBM graft, the percentage of DBM is increased from 0 % seen in Osteoset to 53% in Osteoset 2. According to the manuf acturer, the addition of DBM was intended to increase the bioactivity of the calcium sulfate based disks. The purpose of this study was to compare the effects of adding calcium sulfate, hydroxyapatite and demineralized bone matrix to a polycaprolactone sca ffold using the following parameters: mineralization, adhesion analysis through contact angle measurement, cellular viability, and mechanical testing. Materials and Methods Scaffold Preparation PCL (molecular weight 80 kDa) and the solvent, glacial acetic acid, were both purchased from Sigma Aldrich. Sieve generously donated by Regeneration Technologies (RTIX, Alachua, FL). Spheroidal hydroxylapatite powder was obtained from BDH Chemicals (Poo le, England). Calcium Sulfate hemihydrate powder was purchased from DAP Inc (Baltimore, MD). A 25 wt%
104 PCL solution was prepared to which 35% by wt of the appropriate filler (DBM, CS, HA) was added to the mixture. The scaffolds were prepared according to th e protocol in Chapter 6. Briefly, PCL pellets were dissolved in glacial acetic acid at a concentration of 25% (w/v) in a beaker and stirred at 65C. Once the PCL was completely dissolved (1 2 hrs), sodium chloride (NaCl) particles at 90% w/w of PCL were a dded. The polymer solution was allowed to cool at room temperature for 30 minutes. Once the temperature of the polymer solution reached 37C, the polymer solution was cast into well culture microplates (Corning Life Sciences). Demineralized bone matrix wa s then added to the resulting scaffolds. Four types of scaffolds were formed; pure polymeric scaffolds, PCL/DBM scaffolds that contained approximately 35% DBM (w/w), PCL/CS scaffolds that contained approximately 35% Calcium Sulfate (w/w), and PCL/HA scaffo lds that contained approximately 35% Hydroxyapatite (w/w). After the evaporation of the solvent at 23C (RT) overnight, the resulting solidified scaffolds were placed in placed in 50 mL conical tubes containing 15 mLs of distilled water, with a ratio of sa mple surface to DI volume of 0.5 cm2/mL. The scaffolds were sonicated for 3 hours at 37C and were freeze dried overnight. Degradation Media Scaffolds were incubated in 15 ml of simulated body fluid (SBF) that was maintained at 37C and changed every thre e days for up to 20 days The sample surface to DI volume was 0.5 cm 2 volume is available during the entire test. SBF w as prepared according to Kokubo
105 Water U ptake and Mass Loss of Scaffolds Scaffold s were weighed once prior to the degradation study and then every five days for 20 days. At each time points, scaffolds were removed from the SBF, weighed to measure water uptake, rinsed with DI water and blotted dry with a Kimwipe till no liquid was seen on the wipe. After, the scaffolds were freeze dried for 24 hour to ensure complete removal of water, and weighed. Six samples were weighed for each group (35% DBM, 35% CS, 35% HA and pure PCL). Water absorption was measured using equation 7 1 (7 1 ) where Ww (wet weight) and Wd (dry weight) are specimen weight after and before immersion in SBF respectively. Mass loss of the scaffold was measured using equation 7 2. (7 2 ) where Wi (initial weight) and Wf (final weight) are specimen weight before and after immersion in SBF, respectively. pH of the scaffolds was also taken. Bioactivity Analysis Mineralization on the scaffolds was monitored systematically by collecting the supernatants and anal yzing the levels of ionized calcium and phosphorus using Stanbio Laboratory, L.P. Kits No. 150 and No. 160, respectively (Boerne, TX). This was done at time periods of 14 and 21 days post immersion in simulated body fluid (SBF) at 37C. The ionic concent Mg2+, 1.5; Cl 148.8; HCO3 4.2; HPO42 1.0; SO42 0.5. Morphological changes on
106 the surface of the scaffolds were characterized using a scanning electron microscope (SEM) equipped wit h energy dispersive X ray spectrometer (EDS) and alizarin red S staining. X Ray diffraction was used to identify and determine the various crystalline compounds on the surface of the scaffolds. Alizarin red S (ARS) staining was used to detect mineralizat ion on the scaffold surfaces [Figure 7 2]. ARS reacts with calcium via its sulfonate and hydroxyl groups. The deposits are stained brick red . Figure 7 2. Molecular Structure of Alizarin Red S dye. The binding of calcium with the dye occurs via it s sulfonate and hydroxyl groups. The specimens were washed three times with phosphate buffered saline (PBS), fixed in 3.7 vol % formaldehyde in PBS for 10 min, and stained with 1 mL of 40 mM ARS staining, with the pH adjusted to 4.1, for 30 min. Excess an d unbound dye was
107 rinsed off the scaffolds using DI water and blotting paper. One set of stained scaffolds was then mounted on cover slides to observe using an optical microscope. The other set of stained scaffolds was transferred into 10 mL conical tube s containing 1.5 mL of .5 M Acetic Acid. The dye bound on the scaffolds was allowed to dissolve in the acetic of 1M NaOH. The absorbance was measured at 550 nm using a spectrophotometer (Shimadzu, Japan). X ray Diffraction X Ray diffraction was used to identify and determine the various crystalline compounds on the surface of the scaffolds, before and after imm ersion in SBF. The sample spectra were collected using a Phillips diffractometer operating in the Bragg Brentano configuration with Co K_ (_ = 1.54056 ) radiation at a current of 20 mA and an accelerating voltage of 40 kV. Spectra were recorded in the ran ge 5
108 Contact angle The surface of the scaffold influences cell adhesion by providing a receiving or inhibiting environment for anchor adhesiveness of a material can be characterized based on the surface tension between a solvent drop at its interface and the homogenous surface. Due to large deviations that can be caused by a non homogen ous surface, non porous scaffolds were used to measure the contact angle. Contact angle was determined by the sessile drop method with a Rame Hart Goniometer. The contact angle is an average of 5 measurements on the scaffold. Viability Analysis Biocompati bility of scaffolds were evaluated using an in vitro viability test with the mouse skeletal muscle myoblast C2C12 cell line. The cells were cultured for 4 days in bovine serum and 5% (v/v) 10 U/ml penicillin/streptomycin. All cell culture reagents were bought from Sigma Aldrich. Scaffolds without cells were used as blanks and cell seeded tissue culture wells served as the controls. After 4 days, AlamarBlue analysis was performed to study the viability and proliferation of the cells on the scaffolds. Scaffolds were incubated with Alamar Blue dye for four hours at 37C in 5% CO2 incubator. The color change was read on a SpectraMax M (Molecular Devices) microplate reader. Absorbance was read at 570 nm wavelength. Results were expressed in terms of optical den sity. Stati s tics Six scaffolds were analyzed for each scaffold type at the various time points for mass change, water absorption, calcium ion release, phosphorous ion release and pH
109 change. Three samples from each scaffold type were analyzed for Alizarin Red S staining and Viability/ Toxicity tests. All statistical analysis was conducted using the GraphPad Instat program. The data within each scaffold type was examined between time points and analyzed using a one way analysis of variance (ANOVA) with stati stical significance at p<0.05. Further, data comparing the scaffolds was examined using two way analysis of variance (ANOVA) with statistical significance at p<0.05. Post hoc analysis was conducted using the Tukey test. The values were expressed as the mean standard error. Results Mass Loss and Water Absorption Scaffold degradation in SBF is a good indicator of how scaffolds will behave in vivo. PCL is a semi crystalline hydrophobic polymer whose complete degradation can take upwards of 2 years. It is then no surprise that the scaffold that lost the least amount of mass was pure PCL with less than 9% loss Scaffolds with Demineralized Bone Matrix (DBM) lost 34% of mass, Hydroxylapatite (HA) lost 41% of mass, and scaffolds with calcium sulfate (CS) lost 5 6% of mass. In preparation of the scaffolds, the dense calcium sulfate tended to sink towards the bottom of the suspension and not mix homogenously with the polycaprolactone been likely faster than that of DBM which had a fairly uniform mix with PCL
110 Figure 7 3. Mass loss of scaffolds immersed in SBF for 20 days. 7 3 A represents mass loss for 35 % calcium sulfate, 7 3B represents mass loss for 35% DBM, 7 3C represents mass loss for 35% HA and 7 3D represe nts mass loss for PCL scaffolds. Moreover, since the scaffolds were only physically blended, the fillers occupied random spaces in the scaffold making them susceptible to falling off the polymer as mechanical interlocking forces were degraded
111 Figure 7 4 Water absorption of scaffolds immersed in SBF for 20 days. The hydrophobic surface of pure PCL scaffolds made it relatively impervious to water absorption. However, scaffolds composed of DBM and CS had high water absorption rates (91% and 47%, respectiv ely). This is likely due to the swelling properties of the collagen in DBM and the hydrophilic nature of calcium sulfate. Mineralization The ionic composition of SBF is very similar to that of human blood plasma; therefore immersion of scaffolds in SBF is the most widely used model for study of mineralization under in vitro conditions. Favorable conditions that accelerate apatite depositions include the presence of functional groups and a pH in the range of 6.55 6.65. Before immersion, the pH of the SBF was at 7.4. After 20 days of immersion, the pH of the scaffolds were 6.28, 6.6, 6.28, and 6.92 for CS, DBM, HA, and pure PCL scaffolds respectively. The results of the mineralization experiment using Alizarin Red S Staining are presented in Fig 7 5 and Fig 7 6.
112 Figure 7 5. Alizarin Red S staining of scaffolds immersed in SBF for 20 days. 35% DBM (A), Pure PCL (B) 35% Calcium Sulfate (C) and 35% Hydroxyapatite (D) is shown. Notice the lack of staining on pure PCL scaffold and the 35% calcium sulfate scaff old. Using the Alizarin Red S Staining, we were able to qualitatively assess mineralization deposited on the surface of the scaffolds. Calcium deposits were not present on the surface of the PCL scaffolds. Moreover, using acid solubulization, mineralizatio n was not detected on the surface of the PCL scaffolds. Similarly, mineralization was not seen on the surface of the calcium sulfate scaffolds yet, it was detected using the acid solubulization technique. It is possible that the calcium detected in this t echnique was a product of non specific binding, meaning the calcium detected was that of calcium sulfate and not of mineralization as a result of immersion in SBF.
113 Figure 7 6. Absorbance values for stained scaffolds. The absorbance values of the inorga nic fillers (CS and HA) were significantly higher than the organic filler, DBM. Pure PCL scaffolds had a lower absorbance value than the other scaffolds. Obviously, it would reason that a higher OD value would suggest mineralization on the scaffolds. This suggests that mineralization was dependent on the composition of the scaffold. Yet, it is important to note that the higher OD value seen for calcium sulfate, and to a lesser degree, hydroxyapatite, scaffolds could be a result of both specific and non spec ific binding. SEM analysis confirmed the presence of calcium and phosphorous on the HA, CS, and DBM scaffolds but none on the PCL scaffold. One of the drawbacks of EDS analysis is that the scan is at a certain spot on the material, rather than the whole material. Therefore, the results are only based on one particular spot on the scaffold. It does, however, confirm no mineralization on the PCL scaffold.
114 Figure 7 7 SEM w/EDS analysis of scaffolds. 35% Hydroxyapatite (A), 35 Calcium Sulfate (B) 35% DB M (C) and pure PCL (D) is shown. Both calcium and phosphorous peaks were seen on Hydroxyapatite scaffolds. They were, to a lesser degree, also seen on calcium sulfate and demineralized bone matrix. They were not seen on PCL scaffolds. X Ray Diffraction Spe ctroscopy was used to qualitatively evaluate the phase compositions of the mineral deposits before and after immersion in SBF. Unfortunately, the elasticity of the polycaprolactone prevented us from grinding the scaffolds into fine powder and therefore eva luating the various phases deposited on the scaffolds was not possible. Rather, we evaluated mineralization on pure calcium sulfate powder, hydroxyapatite powder, and demineralized bone matrix particles. Analysis of pure HA powder before immersion in SBF s howed a highly crystalline structure that has similar peaks to that of CS
115 Figure 7 8 XRD of fillers before and after immersion in SBF. Pure Calcium Sulfate (A), Pure Hydroxyapatite (B) Pure DBM particles (C) is shown. Notice the highly crystalline /sharp peaks shown for hydroxyapatite. Contrast this with the amorphous signal shown for demineralized bone matrix. In the After picture of DBM, there appears to be re mineralization as depicted by the sharper peaks seen. DBM, however, had very few distin ct peaks, which may be a result of the demineralization process. The majority of the DBM band is an amorphous state that is days, we see similar peaks for CS, HA, and DBM. This suggests that to some degree, it is possible to remineralize DBM. Contact Angle Analysis Analysis of the contact angle of the scaffolds revealed a partial wetting on scaffolds that contained CS, DBM, and HA (Figure 7 9). The pure PCL scaffolds
116 exhibited a hydrophobic scaffold of 77 3.2 Xia et al. reported a 75.5 contact angle for a comparable pure PCL scaffold by . Figure 7 9 Contact angles for 35% Calcium Sulfate (A), 35% DBM scaffold (B) 35% HA scaffold (C) and pure PCL scaffold is shown. The fillers appear to increase the wettability of the hydrophobic polycaprolactone polymer; the contact angle f or the CS scaffold was 57 5, while the DBM and HA containing scaffolds were 64 2.9 and 50 1.1 respectively. This suggests that the various fillers decreased the contact angle, and therefore the hydrophobicity of polycaprolactone. Viability Analysis OD values for the composite scaffolds were found to be statistically different from the pure PCL scaffold. This suggests a lack of proliferation on the PCL scaffolds.
117 However, large deviations were seen on the HA and CS scaffolds. While it is quite clear that scaffolds proliferated far more on the composite scaffolds than the pure PCL scaffolds, the deviations within each subgroup (i.e. scaffold type) are causes for concern. Figure 7 10. The viability of cells on PCL, 35% DBM, 35% HA, 35% CS scaffolds and Tissue Culture Plate (Corning Costar Corp.) at 4 days While the scaffolds were sterilized in a standard series of 2 hour 70% ethanol soaking (2x) followed by a 30 minute PBS rinsing (4X), no confirmation analysis was performed. And while sterilizati on through UV irradiation was discussed, a number of studies have reported on the loss of osteoinductivity in DBM when irradiated  Mechanical Compression Mechanical compressions tests were performed to assess the behavior of the composite scaffolds before and after immersion in SBF
118 Figure 7 11. Before and After Stress Strain Analysis of Composite Scaffolds Immersed in SBF. Figures A and B represent calcium sulfate before and after respectively, Figures C and D represent PCL before and after, Fig ure E and F represent DBM before and after, and Figure G and H represent HA before and after. Y axis (kPa). In particular, we investigated the engineering stress strain behavior of the porous composite scaffolds. The stress strain curve before immersion e xhibits two distinct different regions: a linear elastic region and a plastic region. As the strain increases, the walls of the pores begin to buckle which leads to a stress plateau, exhibited after the linear elastic region. Comparing before and after im mersion behavior, we can clearly
119 see that this plateau region is smoother and extends for a longer period before immersion for all of the composite scaffolds. PCL, however, varied little in behavior before and after immersion. Based on this data it would reason that the scaffolds, initially, behaved in a manner consistent with its composite nature. Basically, 35% calcium sulfate scaffolds took on the behavior of calcium sulfate. Initially, the curve of the 35% calcium sulfate scaffolds exhibits a pattern c onsistent with that of a brittle material. Yet, after immersion in SBF, and more importantly, loss of mass its behavior resembles that of a thermoplastic polymer (i.e. polycaprolactone). Discussion The purpose of this study was to determine the effects of adding Demineralized Bone Matrix, Calcium Sulfate and Hydroxyapatite to the Polycaprolactone polymer. The bulk of this work was conducted on analyzing formation of mineral on the surface of the scaffold. The ability of an implant to bond with native bon e in vivo is paramount to its success as an effective bone substitute. According to Kokubo, we can predict the behavior of a bone substitute in vivo by immersing it in SBF. We assessed mineral formation on the scaffold through Alizarin Red S staining, Scan ning Electron Microscopy and X ray Dispersive Spectroscopy. We also examined the mass loss and water absorption of the various scaffolds. High mass loss and water uptake was mostly seen in the 35% DBM and CS scaffolds. This was not surprising, as DBM has a high collagen content and readily absorbs water. Although calcium sulfate does not have high swelling properties, its surface is prone to cracking which can lead to significant mass loss. While we were ffolds, we performed compression tests,
120 The plateau region of the stress strain curve of the composite scaffolds after immersion were not as smooth and displayed some unevennes s previously not seen. We can deduce that the addition of the fillers to the polycaprolactone polymer was the likely source of this result. Comparing the ultimate compressive strengths (determined from stress strain curve by applying the load until the sca ffold was fractured) of the scaffolds before and after immersion there was a clear but consistent outcome; the addition of the fillers produced a stiffer scaffold, but subsequent to immersion in SBF for 20 days, the composite scaffolds behaved like the so fter polycaprolactone polymer [Figure 7 12]. This result was more pronounced in scaffolds that lost the most mass (35% Calcium Sulfate, 35% Demineralized Bone Matrix). Detection of the mineral deposits on the scaffolds revealed newly mineralized deposits. In Figure 7 5, stained calcium deposits (shown in red) were more prevalent on the 35% DBM and 35% HA scaffolds. The 35% CS scaffold, along with the PCL scaffold, did not stain for any mineral deposits. Positive red staining shows the presence of calcium d eposits. We further analyzed mineral deposits using SEM with EDS analysis. All the composite scaffolds, to some degree, possessed signals for calcium and phosphorous. Yet, it was still not definitive, that mineralization was occurring due to immersion in SBF. In order to definitively determine whether any new mineral deposits were being formed on the scaffolds, we performed diffraction analysis. Unfortunately, due to the elasticity of polycaprolactone, the scaffolds were unable to be analyzed as is as th e protocol required finely ground material. Therefore, we analyzed mineralization of the fillers (calcium sulfate, demineralized bone matrix, and hydroxyapatite) alone. Before
121 immersion in SBF, X ray analysis of calcium sulfate and hydroxyapatite indicated a highly crystalline material with distinct peaks (Figure 7 8). The DBM diffractogram (the image produced by the diffractometer), however, revealed an amorphous material with very few distinct peaks. After immersion in SBF for 20 days, DBM appears to have several distinct peaks a sign of possible demineralization. A study by Figuerido et al.  analyzed the phase and chemical compositions of bone being demineralized with different concentrations of Hydrochloric Acid (HCl). After immersion in 1.2 M HCl for 12 hrs, the obtained diffractogram of the bone showed very few peaks. Rather, an amorphous band was observed. This suggests that the demineralization of bone occurs at 12 hrs in 1.2M HCl. After 48 hours in HCl, the peaks are completely absent. In comp aring our results with this study, we notice many similarities. First, the absence of peaks seen in bone demineralized in 1.2M HCl for 48 hours was also observed in our DBM, pre immersion. Secondly, after immersion in SBF for 20 days, distinct peaks were o bserved on the diffractogram of our DBM similarly to bone before the demineralization process in the Figueiredo et al. study. Figure 7 13 compares the two studies. Figure 7 12. Ultimate Compressive Strength of Scaffolds before (B) and After(A) immersion in SBF.
122 Contact angle is measured in terms of adsorption. The ability of a surface to adsorb molecules onto its surface largely depends on the energy interactions of the substrate (scaffold). Based on our analysis, it is c lear that adding fillers to the polycaprolactone polymer increased its wettability/hydrophilicity. HA Scaffold> CS Scaffold> DBM Scaffold> PCL scaffold This is notable, as mammalian cells are anchorage dependent and tend to favor (in terms of adhesion and proliferation) modestly water wettable or hydrophilic surfaces In vitro viability test demonstrated that the activity (proliferation) of the cells on the PCL scaffold was significantly less than those on the co mposite scaffolds. As a matter of fact, the highest cellular proliferation was found with the HA scaffold. It is interesting to note that the order of the scaffolds in terms of higher cellular proliferation was parallel to the order of hydrophilicity obse rved on the scaffolds. This suggests a relationship between hydrophilicity of the surface of the scaffold and cell adhesion and proliferation. Initially, both a 4 & 7 day AlamarBlue analysis was planned for the viability test. However, the leaking of the C O2 tank prevented a 7 day reading. Future studies should include both days. Composite scaffolds of PCL DBM, PCL CS, PCL HA and pure PCL scaffolds were prepared by a solvent casting/porogen leaching technique. The composites were shown to improve contact a ngle, bioactivity as well as mechanical properties when compared to the pure PCL scaffolds. Mechanical analysis of the composite scaffolds compressive strengths obtained (b oth before and after immersion) are far less than that
123 of cancellous bone (1 100 MPa). Increasing the amount of fillers within the scaffold would likely increase its compressive strength. And while in vitro viability results were stopped earlier than expe cted, initial results indicated that cells proliferated at a higher rate on composite scaffolds when compared to pure PCL scaffolds. Figure 7 13. XRD of bone demineralized for various hours (A) and DBM before (B) and after immersion in SBF (C). Noti ce that at 12 hrs (A) into demineralization, the peaks observed in Hydroxyapatite and bone at 0 hrs, are beginning to disappear. By 48 hrs, those peaks are absent. The same amorphous band is seen in B. Notice, however, after 20 days in immersion in the pea ks reappear. This suggests a remineralization of bone.
124 CHAPTER 8 A NALYSIS OF OSTEOIN DUCTIVE ACTIVITY OF POROUS POLY CAPROLACTONE DEMINERALIZED BONE MATRIX SCAFFOLDS FOR BONE REPAIR: A PRELIMINARY STUDY Introductory Remarks In this study, myogenic C2C12 ce lls which have been shown to express osteoblastic markers in the presence of BMPs were employed . The aim of this study was to investigate and analyze the behavior of C2C12 cells on porous and non porous PCL DBM scaffolds. The behavior of these cells on porous and non porous PCL scaffolds is also reported. In particular, we wanted to evaluate the ability for these scaffolds to support C2C12 cell adhesion, growth and expression. We used Alamar Blue to detect cell viability and growth. ALP expression an d total calcium content were employed to detect osteoblastic expression. Surface characterization of the scaffolds was visualized by Alizarin Red S staining and scanning electron microscopy. Scaffold Preparation PCL (molecular weight 80 kDa) and the solve nt, glacial acetic acid, we re both purchased from Sigma generously donated by Regeneration Technologies (RTIX, Alachua, FL).The size of the DBM and sodi um chloride particles were confirmed using laser diffraction with a Beckman Coulter LS 1320 (Beckman Coulter Inc., Fullerton, CA). PCL/DBM composite scaffolds were formed by using a solvent casting/porogen leaching technique. Initially, PCL pellets were di ssolved in glacial acetic acid at a concentration of 25% (w/v) in a beaker and stirred at 65C. Once the PCL was completely dissolved (1 2 hrs),
125 sodium chloride (NaCl) particles at 90% w/w of PCL were added. The polymer solution was allowed to cool at room temperature for 30 minutes. Once the temperature of the polymer solution reached 37C, the polymer solution was cast into well culture microplates (Corning Life Sciences). Demineralized bone matrix (at 35% wt/wt of the PCL/Glacial Acetic Acid) was then a dded to the PCL DBM resulting scaffolds. The scaffolds were then placed in distilled water to remove any remaining solvent and/or porogen. The scaffolds were freeze dried overnight. Sample Preparation Four types of scaffolds were fabricated for this study Porous and non porous 35% DBM scaffolds and porous and non porous PCL scaffold. The samples were cut into 5 5 mm pieces and sterilized with an ethanol PBS soaking sequence. The samples were then kept in basal media overnight before cells were added. Cell Culture medium (DMEM), supplemented with 10% (v/v) fetal bovine serum and 5% (v/v) 10 U/ml penicillin/streptomycin at 5% CO2 and 37. All cell culture reagents were bought from Sigma Aldrich For experiments, all scaffolds were seeded with 1 10^4 per cm2 of scaffold surface and cultured in basal medium before and after seeding on scaffolds. Media was changed twice every week. While seeding scaffolds with cells, initially an aliquot of cell suspension was deposited on the scaffold. After 10 seconds, media was added to the scaffold to completely immerse the scaffold in liquid. To induce mineral formation by C2C12 cells onto the scaffolds, the basal media was replaced with complete medium conta ining 10 mm b Glycerophosphate (bGP) (Sigma Aldrich, St.
126 Louis, MO) and 50 mg/ml ascorbic acid (Sigma Aldrich). This was done during the seeding of the cells. Cells on the well plate (Corning Costar Corp.) were used for the control. Statistics Three porou s PCL and PCL DBM scaffolds were each tested in two separate experiments (Proliferation and Alkaline Phosphatase). For the quantification of mineralization, three additional scaffolds (of both PCL and PCL DBM) were used. Nonporous scaffolds were used for t he Alizarin Red S Staining as well as for SEM microscopy. All statistical analysis was conducted using the GraphPad Instat program. The data within each scaffold type was examined between time points and analyzed using a one way analysis of variance (ANOVA ) with statistical significance at p<0.05. Further, data comparing the scaffolds was examined using two way analysis of variance (ANOVA) with statistical significance at p<0.05. Post hoc analysis was conducted using the Tukey test. The values were expre ssed as the mean standard error (n=3). Alkaline Phosphatase Activity (ALP) Alkaline Phosphatase activity was measured following incubation of cells on the adapted to acco unt for the scaffold. Briefly, scaffolds were dried on a clean paper towel. The scaffolds were cut in half and then wet weight measurements were taken. Each scaffold was placed in 15 ml conical tube and 2 steel balls and 2 mls cell lysis solution was adde d to the tube. The scaffolds were then disintegrated using a vortexer three times for 10 seconds. In between cycles, the tubes were placed on ice. The contents of the tube were then transferred into a microcentrifuge tube and then
127 centrifuged at 300G for 1 0 min at 4C. The ALP activity was then measured following incubation of the transferred contents and 1 ml of p nitrophenylphosphate substrate solution. The production of p nitrophenol in the presence of ALP was then measured by monitoring light absorbanc e by the solution at 405 nm. Results are expressed in terms of OD. Alizarin Red S Staining Mineralized nodules were stained using Alizarin Red S Staining on the scaffold surfaces. Briefly, the specimens were washed three times with phosphate buffered salin e (PBS), fixed in 3.7 vol % formaldehyde in PBS for 10 min, and stained with 1 mL of 40 mM alizarin red staining, with the pH adjusted to 4.1, for 30 min. Excess and unbound dye was rinsed off the scaffolds using DI water and blotting paper. The scaffold s were then observed using an optical microscope (Olympus BX60 w/SPOT Insight Digital Camera). Calcium Assay Calcium deposition was measured following incubation of cells on the scaffold the scaffold. Briefly, scaffolds were dried on a clean paper towel. The scaffolds were cut in half and then wet weight measurements were taken. Each scaffold was placed in 15 ml conical tube and 2 steel balls and 2 mls of 5% Trichloroacetic acid solution was added to the tube. The scaffolds were then disintegrated using a vortexer three times for 10 seconds. In between cycles, the tubes were placed on ice. The tubes were then incubated at room temperature for 30 minutes. The contents of the tube were then transferred into a microcentrifuge tube and then centrifuged at 300G for 10 min at 4C. 2 mls of 5% Trichloroacetic acid (TCA) was added to the original tube. Once again, the original tube
128 was incubated for 30 minutes and centrifuged at 300G for 10 min. T he two corresponding TCA solutions were then combined and transferred into a clean microcentrifuge tube. The contents were then transferred to a 96 well plate and absorbance was read at 550 nm. SEM Analysis (2% paraformaldehyde, 2.5% glutaraldehyde and .2 M sodium cacodylate buffer) for 2 hours, and then post fixed in 1% osmium tetroxide. The uncoated scaffolds were then examined using a field emission SEM (S HITACHI 3000N). Results Cultured C2C12 cells on 3 D scaffolds were analyzed for viability, proliferation and osteoblastic expression. C2C12 cells on tissue cultured plate (TCP) were used as control. Obviously, the two dimensionality and smooth surface of the culture plate is not an accurate control when compared to the 3D scaffolds. Nevertheless, results are reported. The progression of the C2C12 cells is illustrated in Figure 8 1.
129 Figure 8 1. C2C12 cells cultured on tissue culture plate. A. C2C12 cells at Day 1. B. Day 4. Prolif eration of cells on TCP. C. Day 10 Cells beginning to show osteoblastic morphology (rounded shape). D. Proliferation of C2C12 cells on Day 14. A lkaline Phosphatase Expression Compared to the PCL scaffold, ALP activity in the 35% DBM was significantly highe r. However, ALP activity in the 35% DBM scaffold compared to the cells cultured on TCP was not. This suggests that cells proliferated at a higher rate on the 35% DBM scaffold compared to the PCL scaffold. The results for the scaffolds are reported in Figu re 8 2
130 Figure 8 2. Alkaline Phosphatase activity on scaffolds (n=3). Alizarin Red S Staining Analysis of the cell seeded stained scaffolds revealed a lack of mineralization on the surface of the PCL scaffolds. We reached the same conclusion when we im mersed unseeded PCL scaffolds in simulated body fluid solution for 20 days (Chapter 7). However, 35% DBM scaffolds did stain positive for mineralization (Figure 8 3). It should be noted that the architecture of the 35% DBM scaffold was not able to be prese rved using light microscopy. Therefore, a representative sample of the scaffold was used in this analysis. Cells seeded on pure DBM particles were also stained with Alizarin Red S Stain.
131 Figure 8 3. Cell seeded scaffolds at 14 days. In Figure A, on the left is a stained 35% DBM scaffold, on the right is a stained PCL scaffold. Figure B is a sample of the stained 35% DBM scaffold. Figure D is a sample of the stained PCL scaffold. Figure D is stained cell seeded DBM particle. Notice the dark red staining seen on the 35% DBM scaffold and the DBM particle (arrows) Magnification 50X.
132 Notice in the stained composite scaffold and pure DBM particles, there is a mixture of both intensely stained dark red spots as well as heavily stained black spots. We believe that the red spots seen in both 35% DBM scaffolds and pure DBM are calcium deposits. Whereas the black stains is the result of the dye being adsorbed by the DBM. Again, this is the same c onclusion reached in Chapter 7. Quantification of M ineralization Quan tification of mineralization on the cell seeded scaffolds was performed using a total calcium content assay. Therefore, it is likely that residual mineral released from the DBM contributed to the final result. In order to account for mineralization that co uld have occurred inside the porous scaffold, the scaffold had to be disintegrated which would likely have resulted in an uptick in total calcium content for the 35% DBM scaffold. Results are shown in Figure 8 3. Figure 8 4. Quantification of minerali zation on scaffolds (n=3).
133 In comparing cell seeded 35% DBM scaffold to unseeded 35% DBM scaffold, we can clearly see an increase in mineralization. Furthermore, both seeded and unseeded scaffolds were significantly higher than pure PCL scaffolds. Likewi se, seeded 35 % DBM and unseeded 35% DBM scaffolds had OD values higher than cells on TCP. From this, we can conclude that residual DBM does, in fact, supplies additional mineral Scanning Electron Microscopy Examination of the cell seeded 35% DBM scaffol d by scanning electron microscopy was disappointing. Perhaps, as a result of the fixative, there appeared to be a lot of debris on the scaffold. While elongated structures (similar to morphology of C2C12 cells) appear on the scaffold, we cannot be certain that they were indeed, c ells. SEM images are shown in Figure 8 5 and 8 6. Figure 8 5. Surface of cell seeded 35% DBM scaffold. While the elongated structures resemble the morphology of C2C12 cells, the debris seen on the scaffold precludes us from af firming their presence on the scaffold.
134 Figure 8 6. Surface of cell seeded 35% DBM scaffold. Discussion An illuminating method to understanding the behavior of a biomaterial in vivo is to observe the behavior of cells on the biomaterial in vitro. Obvi ously, the body being a multifaceted system of organs each replete with its own complex network, cannot be wholly replicated. However, in vitro cell culture analysis allows us to closely mimic the challenges found in the human body Bone regeneration is a complex process that involves many types of cells and events that are interdependent. The construction of a bone substitute should take these factors into account. In particular, like natural bone, this bone substitute should degrade in vivo whilst promo ting viable cells to lay new bone.
135 Our approach in this study was to design a composite that combined a biodegradable polymer with an osteoconductive/osteoinductive material. The polymer, polycaprolactone, has been employed for many tissue engineering prod ucts such as stents. The demineralized bone matrix, used in this study, has been shown to induce ectopic bone formation in mouse models. However, its inferior handling properties make it susceptible to migration upon implantation in vivo. The addition of DBM dispersed throughout the PCL results in a scaffold that 1) reinforces and increases the compressive strength of the scaffold 2) provides an additional source of mineral for the remodeling process and 3) provides a cadre of bone morphogenic proteins th at induce pluripotent cells into an osteoblastic lineage. We used a number of analytical assays to monitor and assess the behavior of the cells on the various scaffolds. We used an Alkaline Phosphatase kit to quantify its expression on the scaffolds. And mineralization on the scaffolds was assessed 2 weeks post seeding on the scaffolds using both a stain and assay kit. ALP production is a mid marker expression of osteoblasts, typically detected around 10 days [90 93]. Mineralization, on the other hand, is typically detected on the 14th day. In vitro testing of cellular behavior on TCP is not a complicated task. However, on a 3D scaffold, it can be. In order to account for this, protocols for biochemical assays had to be modified. A schematic of the process used in the biochemical assay for both ALP production and total calcium content is shown in Fig 8 7.
136 Figure 8 7. A composite picture of the modified protocol for both ALP expression and mineralization quantiification. In Fig A, the scaffolds are intiia lly weighed (wet), then cut in half. In Fig B, the cut scaffolds are placed in a conical tube alonvg with steel balls. Using a vortex,t he scaffolds are disintegrtated. This is done a number of times. In between cycles of vortexing, the disintegrated scaff olds are kept on ice Figure C. Our main goal was to retrieve and preserve the cells or its products. Many factors, using this modified protocol, could alter that goal and consequently, the results. Therefore, results reported on both ALP expression and t he mineralization should be taken with that in mind. In our study, C2C12 cells grown onto 35% DBM scaffolds showed an increase in the amount of alkaline phosphatase production when compared to the purely PCL scaffolds. It was not, however, statistically di fferent in terms of production, to the cells
137 cultured on TCP. In quantifying the mineralization of the cells on the scaffold, we detected more mineralization (as noted by OD value) on 35% DBM versus pure PCL scaffolds. It would be inaccurate, however, to ascribe total calcium content to cellular production, rather than a combination of residual mineral left in the DBM and cellular production. To account for this, we also report OD values for unseeded scaffolds. Alizarin Red S Staining showed a mineral rich layer on the surface of 35% DBM scaffolds. Little to no mineral formation was seen on the surface of the PCL scaffold. Alizarin uptake is proportional to mineral formation, so it would reason that the addition of DBM (in particularly, the residual calcium and phosphate ions) aided in the formation of mineral on its surface. The limitations of these methods, and subsequently its findings, are not lost on the authors. A true comparison cannot be made between a 3D scaffold and a flat surface, such as tissue c ulture plate. Secondly, the process of disintegrating the scaffolds can provide inaccurate results. For example, in disintegrating the scaffolds, we found it easier to disintegrate the composite scaffolds than the purely PCL scaffolds. The rubbery, elasti c nature of PCL prevented us from fully disintegrating the scaffold as was done for the composite scaffold. Furthermore, the addition of Glucocorticoids ( Dexamethasone in combination with ascorbic acid and beta Glycerophosphate ) prevents us from solely i dentifying DBM as the singular source behind osteoinduction. Lastly, our results would need to be confirmed using primary human osteoblasts. C2C12 cells were chosen because they were pluripotent, capable of expression osteoblastic markers and readily avail able.
138 To conclude, in this preliminary work, we sought to analyze and observe the behavior of pluripotent C2C12 cells on a composite scaffold of the osteoconductive and osteoinductive material, demineralized bone matrix, and the biodegradable polymer, poly caprolactone. What we observed was that the C2C12 cells produced or expressed osteoblastic markers more than the pure PCL scaffold. The findings of this study give added weight to the theory that DBM can induce pluripotent cells into an osteoblastic linea ge.
139 CHAPTER 9 SUMMARY AND FUTURE D IRECTIONS Introductory Remarks Tissue Engineering combines living cells with a support structure that degrades in vivo, leaving behind a 3D assembly that is functionally, structurally, and mechanically equal to or bette r than the native tissue it replaced Cell selection and material selection are two key components for a successful tissue engineered replacement. Designing, constructing, and optimizing those key components to maximize their respective advantages is a challenge that will continue well beyond this dissertation. The field of tissue engineering has come to serve as both the hope for patients with bone related injuries and the inspiration for researchers to develop the next generation of viable bone graft s ubstitutes. Bone is a dynamic tissue that serves a biological chemical and mechanical purpose in vivo. Thus, the successful bone graft substitute must account for all three factors. Bone substitutes have been made from natural polymers such as chitosan an d coral to synthetic polymers, such as poly (lactide co glycolide) and polycaprolactone. Nevertheless, consistent in these bone substitutes is the combination of a biodegradable material and a biological additive morphogenic proteins, growth factors or ce lls. The complexity of bone renders the development of a complex bone substitute. Plaster of Paris (calcium sulfate hemihydrate, CaSO4) has been used as a bone substitute since 1892. Surprisingly, it is still being used today. Much work has been done on the development of new composite designs that take into account all three functions of bone yet, more work is clearly needed. This is the motivation that drove the work described in this dissertation.
140 The work, thus far, is heavily focused on the synthesi s, characterization, modification and in vitro degradation of the scaffold. A comparative analysis of the effects of demineralized bone matrix, hydroxyapatite, and calcium sulfate on polycaprolactone polymer is also presented. Finally, we discuss prelimin ary results on the behavior of pluripotent C2C12 cells on our composite scaffold. The broad hypothesis behind the construct of this composite scaffold was simple; the combination of an osteoinductive/osteoconductive demineralized bone matrix coupled with a biodegradable and biocompatible polymer would yield a scaffold capable of encouraging and enhancing the formation of new bone in vivo even as the scaffold degraded. Synthesis, Characterization and In Vitro Evaluation of Composite Scaffolds Initially, ou r work focused on the synthesis of the pure polycaprolactone scaffold using a salt porogen technique. Analysis of the scaffolds using scanning electron microscopy revealed an uneven, slightly porous scaffold that retained sodium chloride particles. We sou ght to reduce the amount of sodium chloride through ultrasonication and altering the washing sequence. Using SEM with EDS analysis, scaffolds sonicated for 1 and 3 hours and those washed at 37C and 45C revealed little to no residual salt particles. Thos e sonicated for a longer time or washed at a higher temperature were characterized by higher residual sodium chloride particles, a lightly porous surface and a gummy like appearance. We further explored the physical and chemical characteristics of the sca ffolds by varying DBM content within the scaffold using mass loss, pH measurements, and surface mineralization studies. In our degradation studies, the scaffolds were immersed in SBF solution similar in ionic concentration to that of human plasma.
141 The mas s of all scaffold types declined over the 7 week experiment; however, some declined more than others. The mass of the 50% DBM declined by nearly 43%, while those of the pure PCL scaffolds only declined by 7%. Since the DBM was only physically blended into the scaffold, we expected a decline in scaffold mass. The pH of the SBF solution for the composite scaffolds dipped slightly through the 7 week experiment. The SBF solution that contained pure PCL scaffolds remained stable throughout. It is possible that t he fluctuations in the pH of the composite scaffolds, in particular the 50% DBM scaffold, was a result of residual acidic solvent that possibly remained in the scaffold. Calcium deposits were observed on the composite scaffolds. According to Kokubo, miner alization on the surface of the scaffold can be used to predict in vivo behavior. We observed mineralization using both the Von Kossa Staining [Figure 6 5] and SEM imaging [Figure 6 9]. It was particularly pronounced in the 35% and 50% DBM scaffolds. While the 50% DBM scaffold theoretically contained more nucleation sites for mineralization, no significant difference in mineralization on the 35% and 50% DBM scaffolds was seen. There was, however, a lack of mineralization on both the PCL polymer and the 25% DBM scaffold. This finding was affirmed with EDS analysis. While the signals of calcium and phosphorous were higher for the 50% DBM scaffold, it was not significantly different from that of the 35% DBM scaffold. In Vitro Evaluation of the Effects of A dd ing C ommon B one V oid F illers to a Polycaprolactone S caffold In Chapter 7, we compared the effects of adding common bone void fillers such as hydroxyapatite (HA), calcium sulfate (CS) and demineralized bone matrix (DBM) to a
142 polycaprolactone (PCL) scaffold. Specifically, we examined the in vitro degradation and reprecipitation behavior of the scaffolds after 20 days in simulated body fluid. The hydrophobic surface of pure PCL scaffolds made it relatively impervious to water absorption and mass loss. However, the swelling properties of the collagen in DBM and the hydrophilic nature of the CS produced scaffolds with significant mass loss (34 % and 56% respectively) and high water absorption rates (91 % and 47 % respectively). The ultimate compressive strength o f the calcium sulfate added scaffold post immersion was reduced by nearly a third (28%) and the ultimate compressive strength of the demineralized bone matrix added scaffold was reduced by 10%. Detection of the mineral deposits on the scaffolds revealed ne wly mineralized deposits on the all composite scaffolds. Figure 7 5 shows the stained calcium deposits seen on the 35% DBM and 35% HA scaffolds, and to a lesser extent on the 35% CS scaffolds. They were not seen on the pure PCL scaffolds. Further analysis of mineralization using X ray diffraction before and after immersion in SBF revealed the surface mineralization detected to be a result of calcium and phosphate binding rather than residual mineral within the fillers. The potential for the composite sca ffolds to support cellular activity was examined by seeding C2C12 mouse muscle cells on the scaffolds and analyzing for cellular proliferation. Using an AlamarBlue viability assay, proliferation of the cells on the PCL scaffold was significantly less than those on the composite scaffolds. Furthermore, we found a clear correlation between water contact angle and C2C12 proliferation: the lower the contact angle (hydrophilicity), the higher the proliferation.
143 C2C12 Cell Proliferation and Differentiation on Com posite S caffolds The influence of demineralized bone matrix in composite scaffolds on pluripotent C2C12 cells were analyzed using Alkaline Phosphatase (ALP) expression, mineral te and pure scaffolds were also analyzed for ALP expression and total calcium content in order to account for residual mineral content left in the DBM. Analysis indicated that C2C12 cells produced or expressed osteoblastic markers more on the composite sca ffolds than the pure PCL scaffold. Alkaline Phosphate activity was assessed at Day 10. Compared to the PCL scaffold, ALP activity in the 35% DBM was significantly higher. However, ALP activity in the 35% DBM scaffold compared to the cells cultured on TCP w as not. This suggests that cells proliferated at a higher rate on the 35% DBM scaffold compared to the PCL scaffold. Removal of the cells and/or its products presented us with some challenges. The scaffold had to be completely disintegrated using a fairly complex procedure involving steel balls and centrifugation at high speeds. The elasticity of the scaffolds, in particular the pure PCL scaffold, prevented the complete disintegration of the scaffold. This possibly could have had a dampening effect on ALP analysis Analysis of the Alizarin Red stained scaffolds revealed a lack of new mineralization on the surface of the cell seeded PCL scaffolds. We reached the same conclusion when we immersed unseeded PCL scaffolds in simulated body fluid solution for 2 0 days (Chapter 7). However, 35% DBM scaffolds did stain positive for new mineralization. While the Alizarin Red staining was able to quantify new mineral formation on the scaffold, it also heavily stained residual mineral from the 35% DBM scaffold. To acc ount for the non specific staining, total calcium content was measured using both cell seeded
144 and unseeded scaffolds at Day 14. In comparing cell seeded 35% DBM scaffold to unseeded 35% DBM scaffold, we can clearly see an increase in mineralization. Both seeded and unseeded composite scaffolds were significantly higher than seeded and unseeded PCL scaffolds. Final Conclusions and Future Directions While the preliminary work presented in this dissertation is encouraging, there are a number of features of t his scaffold that can be improved. Although the mechanical properties of this scaffold are enough to handle and hold the very flaky DBM, it is not sufficient for a weight bearing defect site. The addition of DBM could help in this regard. Secondly, the imm ersion of scaffolds in SBF before cell seeding would aid in the formation of bone by osteoblasts by providing additional ions for mineralization. We saw in Chapter 6 that 35% DBM scaffolds immersed in SBF formed a layer mineralization on its surface after 14 days in SBF. Furthermore, the effects of high temperatures on DBM should be examined. Attempts were made to determine if DBM in extreme temperatures (4C and 65C) lost or reduced its osteoinductive potential when compared to DBM at room temperature. U nfortunately, the DBM incubated at high temperatures in media produced bacterial spores and the experiment was stopped. Future work should explore this topic. Moving towards the gold standard of an autograft is the challenge of bone tissue engineering. W hile great strides have been made, an aging population coupled with a great number of soldiers returning home with bone related injuries sustained in battle the need for a viable and effective bone substitute has never been so high.
145 APPENDIX PROTOCOL S Analytical Assay: Alkaline Phosphatase Activity Cell lysis solution: Triton X 100, 0.2% (v/v) + 5 mM magnesium chloride in DI water ALP kit Sodium hydroxide solution, 0.2 M Paper towel Razor blade Eppendorf tubes Microcentrifuge tubes with screw cap, 2 ml Steel balls Parafilm Analytical balance MiniBeadBeater Microcentrifuge Microtitration plate Microplate reader Waterbath 1. Grow C2C12 cells on scaffolds in 12 well plates 2. Dry scaffolds on clean paper towel for 3 min. 3. Measure wet weight of the scaffolds. Cut scaffolds into halves and measure the wet weight of halves. 4. Put each scaffold into a labeled Microcentrifuge tube. 5. Add 2 steel balls and 1 ml cell lysis solution per tube. 6. Disintegrate scaffold by using a MiniBeadBeater 3 times at 25,000 rpm for 10s. P lace on ice between cycles for cooling. 7. Transfer the content of the tube into a clean labeled tube without transferring the steel balls. 8. Centrifuge at 300 g for 10 min at 4 C. 9. Transfer the supernate into a clean, labeled tube. Avoid destruction of the pell et. 10. Run ALP assay immediately.
146 Analytical Assay: Total Calcium TCA 5%: trichloroacetic acid 5% (v/v) in DI water Calcium standard 10 mg/dl Calcium binding reagent Calcium buffer reagent Paper towel Razor blade Microcentrifuge tubes with screw cap, 2 ml Steel balls Parafilm Analytical balance MiniBeadBeater Microcentrifuge Microtiter plate Microplate reader Protocol: 1. Grow C2C12 cells on scaffolds in 12 well plates 2. Dry scaffolds on clean paper towel for 3 min. 3. Measure wet weight of the scaffolds. Cut sca ffolds into halves and measure the wet weight of halves. 4. Transfer each scaffold into a labeled Microcentrifuge tube. 5. Add 2 steel balls and 1 ml TCA 5% per tube. 6. Close tube firmly and wrap with Parafilm. 7. Disintegrate scaffold by using a MiniBeadBeater 3 tim es at 25,000 rpm for 10s. Place on ice between cycles for cooling. 8. Incubate the tubes at room temperature for 20 min. 9. Centrifuge at 3000 g for 10 min at 4 C. 10. Pipette the supernate into a clean, labeled tube without disrupting the pellet. 11. Add 1 ml of TCA 5% into the tube with the steel balls. 12. Incubate the tubes at room temp for 30 min. 13. Combine the 2 corresponding TCA solutions. 14. Centrifuge at 3000g for 10 min. 15. Transfer supernate into clean, labeled tube without disrupting the pellet. 16. Run Calcium Assay as desc ribed.
147 L I S T O F REFERENCES 1. Beller, G. 2010 The cost of our aging population. Journal of Nuclear Cardiology, 17 (3), 345 346. (doi: 10.1007/s12350 010 9234 2) 2. Burg, K. J. L., Porter, S., & Kellam, J. F. 2000 Biomaterial developments for bone tissue enginee ring. Biomaterials, 21 (23), 2347 2359. (doi: 10.1016/s0142 9612(00)00102 2) 3. Laurie, S. W., Kaban, L. B., Mulliken, J. B., & Murray, J. E. 1984 Donor site morbidity after harvesting rib and iliac bone. Plastic and Reconstructive Surgery, 73 (6), 933 938. 4. Pa derni, S., Terzi, S., & Amendola, L. 2009 Major bone defect treatment with an osteoconductive bone substitute. Musculoskeletal Surgery, 93 (2), 89 96. (doi: 10.1007/s12306 009 0028 0) 5. Arrington, E. D., Smith, W. J., Chambers, H. G., Bucknell, A. L., & Davi no, N. A. 1996 Complications of Iliac Crest Bone Graft Harvesting. Clinical Orthopaedics and Related Research, 329 300 309. 6. Sowers, M. R., Clark, M. K., Hollis, B., Wallace, R. B., & Jannausch, M. 1992 Radial bone mineral density in pre and perimenopaus al women: A prospective study of rates and risk factors for loss. Journal of Bone and Mineral Research, 7 (6), 647 657. (doi: 10.1002/jbmr.5650070609) 7. Fatayerji, D., & Eastell, R. 1999 Age Related Changes in Bone Turnover in Men. Journal of Bone and Minera l Research, 14 (7), 1203 1210. (doi: 10.1359/jbmr.1918.104.22.1683) 8. Teitelbaum, S. L. 2000 Bone Resorption by Osteoclasts. Science, 289 ( 5484), 1504 1508. (doi:10.1126/science.289.5484.1504) 9. Ducy, P., Schinke, T., & Karsenty, G. 2000 The Osteoblast: A Sophis ticated Fibroblast under Central Surveillance. Science, 289 (5484), 1501 1504. (doi: 10.1126/science.289.5484.1501) 10. Rodan, G. A. 1992 Introduction to bone biology. Bone, 13, Supplement 1(0), S3 S6. (doi: 10.1016/s8756 3282(09)80003 3) 11. Kearney, J. N., Bojar R., & Holland, K. T. 1993 Ethylene oxide sterilisation of allogenic bone implants. Clinical Materials, 12 (3), 129 135. (doi: 10.1016/0267 6605(93)90063 d) 12. Zhang, Q., Cornu, O., & Delloye, C. 1997 Ethylene oxide does not extinguish the osteoinductive capa city of demineralized bone: A reappraisal in rats. Acta Orthopaedica, 68 ( 2), 104 108 (doi:10.3109/17453679709003989 )
148 13. Haimi, S., Vienonen, A., Hirn, M., Pelto, M., Virtanen, V., & Suuronen, R. 2008 The effect of chemical cleansing procedures combined with peracetic acid ethanol sterilization on biomechanical properties of cortical bone. Biologicals, 36 (2), 99 104.( doi: 10.1016/j.biologicals.2007.06.001) 14. Woodard, J. R., Hilldore, A. J., Lan, S. K., Park, C. J., Morgan, A. W., Eurell, J. A. C., Wagoner John son, A. J. 2007 The mechanical properties and osteoconductivity of hydroxyapatite bone scaffolds with multi scale porosity. Biomaterials, 28 (1), 45 54. (doi: 10.1016/j.biomaterials.2006.08.021) 15. Marie, P. 1997 Growth factors and bone formation in osteoporos is: roles for IGF I and TGF beta. Revue du rhumatisme (English ed.), 64 (1), 44 53. 16. Shea, L. D., Wang, D., Franceschi, R. T., & Mooney, D. J. 2000 Engineered bone development from a pre osteoblast cell line on three dimensional scaffolds. Tissue engineeri ng, 6 (6), 605 617. 17. Yang, X., Tare, R. S., Partridge, K. A., Roach, H. I., Clarke, N. M. P., Howdle, S. M.,Oreffo, R. O. C. 2003 Induction of Human Osteoprogenitor Chemotaxis, Proliferation, Differentiation, and Bone Formation by Osteoblast Stimulating Fac tor 1/Pleiotrophin: Osteoconductive Biomimetic Scaffolds for Tissue Engineering. Journal of Bone and Mineral Research, 18 (1), 47 57. (doi: 10.1359/jbmr.2003.18.1.47) 18. Kenley, R. A., Yim, K., Abrams, J., Ron, E., Turek, T., Marden, L. J., & Hollinger, J. O. 1993 Biotechnology and Bone Graft Substitutes. Pharmaceutical Research, 10 (10), 1393 1401. (doi: 10.1023/a:1018902720816) 19. Yeong, W. Y., Chua, C. K., Leong, K. F., & Chandrasekaran, M. 2004 Rapid prototyping in tissue engineering: challenges and potential. Trends in Biotechnology 22 (12), 643 652. (doi: 10.1016/j.tibtech.2004.10.004) 20. Borenstein, J. T., Terai, H., King, K. R., Weinberg, E. J., Kaazempur Mofrad, M. R., & Vacanti, J. P. 2002 Microfabrication Technology for Vascularized Tissue Engineering. Bio medical Microdevices, 4 (3), 167 175. (doi: 10.1023/a:1016040212127) 21. Friedman, E. A., & Friedman, A. L. 2006 Payment for donor kidneys: Pros and cons. Kidney Int, 69 (6), 960 962. 22. Griffith, L. G., & Naughton, G. 2002 Tissue Engineering -Current Challenges a nd Expanding Opportunities. Science, 295( 5557), 1009 1014. (doi: 10.1126/science.1069210) 23. Lendeckel, S., Jdicke, A., Christophis, P., Heidinger, K., Wolff, J., Fraser, J. K., Howaldt, H. P. 2004 Autologous stem cells (adipose) and fibrin glue used to trea t widespread traumatic calvarial defects: case report. Journal of Cranio Maxillofacial Surgery, 32 (6), 370 373. (doi: 10.1016/j.jcms.2004.06.002)
149 24. Kemp, P. 2005 Cell therapy back on the up curve. Regenerative Medicine, 1 (1), 9 14.( doi: 10.2217/17460751.1 .1.9) 25. Ryan, J., Barry, F., Murphy, J., & Mahon, B. 2005 Mesenchymal stem cells avoid allogeneic rejection. Journal of Inflammation, 2 (1), 1 11. (doi: 10.1186/1476 9255 2 8 ) 26. Hutmacher, D. W. 2000 Scaffolds in tissue engineering bone and cartilage. Biomate rials, 21 (24), 2529 2543. (doi: 10.1016/s0142 9612(00)00121 6) 27. Rai, B., Teoh, S. H., Hutmacher, D. W., Cao, T., & Ho, K. H. 2005 Novel PCL based honeycomb scaffolds as drug delivery systems for rhBMP 2. Biomaterials, 26 (17), 3739 3748. (doi: 10.1016/j.bio materials.2004.09.052) 28. Eppley, B. L., Pietrzak, W. S., & Blanton, M. W. 2005 Allograft and Alloplastic Bone Substitutes: A Review of Science and Technology For the Craniomaxillofacial Surgeon. Journal of Craniofacial Surgery, 16 ( 6), 981 989 (doi: 910.1097 /1001.scs.0000179662.0000138172.dd.) 29. Burchardt, H. 1983 The biology of bone graft repair Clinical Orthopaedics and Related Research 174 28 42. 30. Schwartz, H. E., Matava, M. J., Proch, F. S., Butler, C. A., Ratcliffe, A., Levy, M., & Butler, D. L. 2006 Th e Effect of Gamma Irradiation on Anterior Cruciate Ligament Allograft Biomechanical and Biochemical Properties in the Caprine Model at Time Zero and at 6 Months After Surgery. The American Journal of Sports Medicine, 34 (11), 1747 1755. (doi: 10.1177/036354 6506288851) 31. Kearney, J. N. 2005 Guidelines on processing and clinical use of skin allografts. Clinics in Dermatology, 23 (4), 357 364. (doi: 10.1016/j.clindermatol.2004.07.018) 32. Nguyen, H., Morgan, D. A. F., & Forwood, M. R. 2011 Validation of 11 kGy as a Ra diation Sterilization Dose for Frozen Bone Allografts. The Journal of Arthroplasty, 26 (2), 303 308.(doi: 10.1016/j.arth.2010.03.032) 33. Cook, P. 2011 What Constitutes Adequate Public Consultation? Xenotransplantation Proceeds in Australia. Journal of Bioethi cal Inquiry, 8( 1), 67 70. (doi: 10.1007/s11673 010 9269 8.) 34. Miron, R. J., Hedbom, E., Saulacic, N., Zhang, Y., Sculean, A., Bosshardt, D. D., & Buser, D. 2011 Osteogenic Potential of Autogenous Bone Grafts Harvested with Four Different Surgical Techniques. Journal of Dental Research, 90 (12), 1428 1433. (doi: 10.1177/0022034511422718) 35. Sipe, J. D. 2002 Tissue Engineering and Reparative Medicine. Annals of the New York Academy of Sciences, 961 (1), 1 9. (doi: 10.1111/j.1749 6632.2002.tb03040.x)
150 36. Nair, L. S., & Laurencin, C. T. 2007 Biodegradable polymers as biomaterials. Progress in Polymer Science, 32 (8 9), 762 798. (doi: 10.1016/j.progpolymsci.2007.05.017) 37. Sonia, T. A., & Sharma, C. P. 2012 An overview of natural polymers for oral insulin delivery. Drug Disc overy Today, 17 (13 14), 784 792. (doi: 10.1016/j.drudis.2012.03.019) 38. Wu, K. J., & Odom, R. W. (1998). Peer Reviewed: Characterizing Synthetic Polymers by MALDI MS. Analytical Chemistry, 70 (13), 456A 461A. (doi: 10.1021/ac981910q ) 39. Piskin, E. 1995 Biodegra dable polymers as biomaterials. Journal of Biomaterials Science, Polymer Edition, 6 (9), 775 795. (doi: 10.1163/156856295x00175) 40. Coombes, A. G. A., & Meikle, M. C. 1994 Resorbable synthetic polymers replacements for bone graft. Clinical Materials, 17 (1), 35 67. (doi: 10.1016/0267 6605(94)90046 9) 41. Tayalia, P., & Mooney, D. J. 2009 Controlled Growth Factor Delivery for Tissue Engineering. Advanced Materials, 21 (32 33), 3269 3285. (doi: 10.1002/adma.200900241) 42. Takezawa, T. 2003 A strategy for the development o f tissue engineering scaffolds that regulate cell behavior. Biomaterials, 24 (13), 2267 2275.( doi: 10.1016/s0142 9612(03)00038 3) 43. Oh, S. H., Park, I. K., Kim, J. M., & Lee, J. H. 200 In vitro and in vivo characteristics of PCL scaffolds with pore size gra dient fabricated by a centrifugation method. Biomaterials, 28 (9), 1664 1671. (doi: 10.1016/j.biomaterials.2006.11.024 ) 44. Chen, V. J., & Ma, P. X. 2004 Nano fibrous poly(l lactic acid) scaffolds with interconnected spherical macropores. Biomaterials, 25 (11) 2065 2073. (doi: 10.1016/j.biomaterials.2003.08.058) 45. Karande, T. S., Ong, J. L., & Agrawal, C. M. 2004 Diffusion in Musculoskeletal Tissue Engineering Scaffolds: Design Issues Related to Porosity, Permeability, Architecture, and Nutrient Mixing. Annals o f Biomedical Engineering, 32 (12), 1728 1743. (doi: 10.1007/s10439 004 7825 2) 46. Yang, Y., Bolikal, D., Becker, M. L., Kohn, J., Zeiger, D. N., & Simon, C. G. 2008 Combinatorial Polymer Scaffold Libraries for Screening Cell Biomaterial Interactions in 3D. Ad vanced Materials, 20 (11), 2037 2043. (doi: 10.1002/adma.200702088)
151 47. Pham, D. T., & Gault, R. S. 1998 A comparison of rapid prototyping technologies. International Journal of Machine Tools and Manufacture, 38 (10 11), 1257 1287. (doi: 10.1016/s0890 6955(97)0 0137 5) 48. M K Sah, J Sadanand and K Pramanik. Computational Approaches in Tissue Engineering. International Journal of Computer Applications 27 (4), 13 20. 49. Hutmacher, D. W., Sittinger, M., & Risbud, M. V. 2004 Scaffold based tissue engineering: rationale fo r computer aided design and solid free form fabrication systems. Trends in Biotechnology, 22 (7), 354 362.( doi: 10.1016/j.tibtech.2004.05.005) 50. Chen, Z., Li, D., Lu, B., Tang, Y., Sun, M., & Xu, S. 2005 Fabrication of osteo structure analogous scaffolds via fused deposition modeling. Scripta Materialia, 52 (2), 157 161. (doi: 10.1016/j.scriptamat.2004.08.006) 51. Hou, Q., Grijpma, D. W., & Feijen, J. 2003 Porous polymeric structures for tissue engineering prepared by a coagulation, compression moulding and salt leaching technique. Biomaterials, 24 (11), 1937 1947. (doi: 10.1016/s0142 9612(02)00562 8) 52. Buckwalter, J. A., Glimcher, M. J., Cooper, R. R., & Recker, R. 1996 Bone biology. II: Formation, form, modeling, remodeling, and regulation of cell function. Instru ctional course lectures, 45, 387 399 53. Gimble, J., & Nuttall, M. 2004 Bone and fat. Endocrine, 23 (2), 183 188. (doi: 10.1385/endo:23:2 3:183) 54. Meier, D. E., Orwoll, E. S., & Jones, J. M. 1984 Marked disparity between trabecular and cortical bone loss with a ge in healthy men. Measurement by vertebral computed tomography and radial photon absorptiometry. Annals of internal medicine 5 605 612. 55. Hill, P. A. 1998 Bone remodelling. British journal of orthodontics, 25 (2), 101 107. 56. Ecarot Charrier, B., Glorieux, F H., van der Rest, M., & Pereira, G. 1983 Osteoblasts isolated from mouse calvaria initiate matrix mineralization in culture. The Journal of Cell Biology, 96 (3), 639 643.( doi: 10.1083/jcb.96.3.639) 57. Heinegrd, D., & Oldberg, A. 1989. Structure and biology of cartilage and bone matrix noncollagenous macromolecules. The FASEB Journal, 3 (9), 2042 2051. 58. Osaka, A., Miura, Y., Takeuchi, K., Asada, M., & Takahashi, K. 1991 Calcium apatite prepared from calcium hydroxide and orthophosphoric acid. Journal of Mater ials Science: Materials in Medicine, 2 (1), 51 55. (doi: 10.1007/bf00701687)
152 59. Clarke, B. 2008 Normal Bone Anatomy and Physiology Clinical Journal of the American Society of Nephrology 3 (Supplement 3), S131 S13. (doi: 10.2215/cjn.04151206) 60. Parfitt, A. 1984 The cellular basis of bone remodeling: The quantum concept reexamined in light of recent advances in the cell biology of bone. Calcified Tissue International, 36 (0), S37 S45. (doi: 10.1007/bf02406132.) 61. Kogianni, G., & Noble, B. 2007 The biology of osteocyt es. Current Osteoporosis Reports, 5 (2), 81 86. (doi: 10.1007/s11914 007 0007 z) 62. Aarden, E. M., Nijweide, P. J., & Burger, E. H. 1994 Function of osteocytes in bone. Journal of Cellular Biochemistry, 55 (3), 287 299. (doi: 10.1002/jcb.240550304) 63. Franz Odend aal, T. A., Hall, B. K., & Witten, P. E. 2006 Buried alive: How osteoblasts become osteocytes. Developmental Dynamics, 235 (1), 176 190. (doi: 10.1002/dvdy.20603) 64. Miller, S. C. 1989 Bone lining cells: structure and function. Scanning microscopy, 3 (3), 953 960; D iscussion 960 951. 65. Miller, S., & Jee, W. 1987 The bone lining cell: A distinct phenotype? Calcified Tissue International, 41 (1), 1 5. (doi: 10.1007/bf02555122) 66. Islam, A., Glomski, C., & Henderson, E. S. 1990 Bone lining (endosteal) cells and hemato poiesis: A light microscopic study of normal and pathologic human bone marrow in plastic embedded sections. The Anatomical Record, 227 (3), 300 306. (doi: 10.1002/ar.1092270304) 67. Boyce, B. F., Yao, Z., Zhang, Q., Guo, R., Lu, Y. A. N., Schwarz, E. M., & Xing L. 2007 New Roles for Osteoclasts in Bone. Annals of the New York Academy of Sciences, 1116 (1), 245 254. (doi: 10.1196/annals.1402.084) 68. Lorenzo, J. A. 2011 Do osteoclasts have dual roles: Bone resorption and antigen presentation? IBMS BoneKEy 8 (1), 37 40. (10.1138/20110488) 69. Blair, H. C., Kahn, A. J., Crouch, E. C., Jeffrey, J. J., & Teitelbaum, S. L. 1986 Isolated osteoclasts resorb the organic and inorganic components of bone. The Journal of Cell Biology, 102 (4), 1164 1172.( doi: 10.1083/jcb.102.4.1164 ) 70. Murphy, G. and J. J. Reynolds 2003 Extracellular Matrix Degradation. Connective Tissue and Its Heritable Disorders pp. 343 384. New York, New York: John Wiley & Sons
153 71. Neutzsky Wulff, A., Sorensen, M., Kocijancic, D., Leeming, D., Dziegiel, M., Karsdal, M ., & Henriksen, K. 2010 Alterations in osteoclast function and phenotype induced by different inhibitors of bone resorption implications for osteoclast quality. BMC Musculoskeletal Disorders, 11 (1), 109. 72. Rosenzweig, A. and R. J. Pignolo 2011 Osteobiology of Aging : Fractures in the Elderly pp. 3 42 New York, New York: Humana Press. 73. Long, M., & Rack, H. J. 1998 Titanium alloys in total joint replacement a materials science perspective. Biomaterials, 19 (18), 1621 1639. (doi: 10.1016/s0142 9612(97)00146 4) 74. Galante, J., Rostoker, W., & Ray, R. D. 1970. Physical properties of trabecular bone Calcified Tissue International 5 (1) 236 246. 75. Cummine, J., & Nade, S. 1977 Osteogenesis after bone and bone marrow transplantation: I. Studies with combined myelo osseou s grafts in the guinea pig. Acta Orthopaedica, 48 (1), 15 24. (doi: 10.3109/17453677708985105) 76. Tanaka, Y., Nakayamada, S., & Okada, Y. 2005 Osteoblasts and osteoclasts in bone remodeling and inflammation. Current drug targets. Inflammation and allergy, 4 (3 ), 325 328. 77. Mosekilde, L. 2000 Age related changes in bone mass, structure, and strength effects of loading. Zeitschrift fr Rheumatologie, 59 (7), I1 I9. (doi: 10.1007/s003930070031 ) 78. Barrre, F., van Blitterswijk, C. A., & de Groot, K. 2006 Bone regene ration: molecular and cellular interactions with calcium phosphate ceramics. International Journal of Nanomedicine, 1 (3), 317 332. 79. Manolagas, S. C. 1999 Editorial: Cell Number Versus Cell Vigor What Really Matters to a Regenerating Skeleton? Endocrinolog y, 140 (10), 4377 4381. (doi: 10.1210/en.140.10.4377) 80. Parfitt, A. M. 2002 Targeted and nontargeted bone remodeling: relationship to basic multicellular unit origination and progression Bone 30 (1) 5 7. 81. Jilka, R. L. 2003 Biology of the basic multicellula r unit and the pathophysiology of osteoporosis. Medical and Pediatric Oncology, 41 (3), 182 185. (doi: 10.1002/mpo.10334) 82. Manolagas, S. C., & Jilka, R. L. 1995 Bone Marrow, Cytokines, and Bone Remodeling Emerging Insights into the Pathophysiology of Oste oporosis. New England Journal of Medicine, 332 (5), 305 311. (doi:10.1056/NEJM199502023320506)
154 83. Mundy, G. R. 1995 Bone remodeling and its disorders pp. 21 23 London United Kingdom Taylor and Francis. 84. Raisz, L. G. 1999 Physiology and Pathophysiology of B one Remodeling. Clinical Chemistry, 45( 8), 1353 1358. 85. Vaes, G. 1988 Cellular biology and biochemical mechanism of bone resorption. A review of recent developments on the formation, activation, and mode of action of osteoclasts. Clinical Orthopaedics and R elated Research 231 239 271. 86. Dodds, R. A., Connor, J. R., James, I. E., Lee Rykaczewski, E., Appelbaum, E., Dul, E., & Gowen, M. 1995 Human osteoclasts, not osteoblasts, deposit osteopontin onto resorption surfaces: An in vitro and ex vivo study of remod eling bone. Journal of Bone and Mineral Research, 10 (11), 1666 1680. (doi: 10.1002/jbmr.5650101109) 87. Cowles, E. A., M. E. DeRome, et al. 1998 Mineralization and the Expression of Matrix Proteins During In Vivo Bone Development Calcified Tissue Internationa l 62 (1), 74 82. (doi: 10.1007/s00223990039) 88. Miyazawa, K. 1989 Immunological investigation of intestinal, liver, kidney, bone, placental and serum alkaline phosphatase in cattle. Japanese journal of veterinary science 51 (2) 309 314 89. Seibel, M. J. 2000 Molecular Markers of Bone Turnover: Biochemical, Technical and Analytical Aspects. Osteoporosis International, 11 (18), S18 S29. (doi: 10.1007/s001980070003) 90. Sabokbar, A., Millett, P. J., Myer, B., & Rushton, N. 1994 A rapid, quantitative assay for measurin g alkaline phosphatase activity in osteoblastic cells in vitro. Bone and Mineral, 27 (1), 57 67. (doi: 10.1016/s0169 6009(08)80187 0.) 91. Anderson, H. C., Hsu, H. H., Morris, D. C., Fedde, K. N., & Whyte, M. P. 1997 Matrix vesicles in osteomalacic hypophosphat asia bone contain apatite like mineral crystals. Am J Pathol. 151 (6 1555 1561. 92. Beck, G. R. (2003). Inorganic phosphate as a signaling molecule in osteoblast differentiation. Journal of Cellular Biochemistry, 90 (2), 234 243. (doi: 10.1002/jcb.10622) 93. Gort er de Vries, I., Quartier, E., Boute, P., Wisse, E., & Coomans, D. 1987 Immunocytochemical Localization of Osteocalcin in Developing Rat Teeth. Journal of Dental Research, 66 (3), 784 790. (doi:10.1177/00220345870660031601) 94. Butler, W. T. 1989 The Nature and Significance of Osteopontin. Connective Tissue Research, 23 (2 3), 123 136. (doi:10.3109/03008208909002412)
155 95. Eriksen, E. F., Brixen, K., & Charles, P. 1995 New markers of bone metabolism: clinical use in metabolic bone disease. European Journal of Endo crinology, 132 (3), 251 263. ( doi:0.1530/eje.0.1320251) 96. Butler WT, Ridall AL, McKee MD. 1996 Osteopontin. Principles of Bone Biology pp 167 182. San Diego CA : Academic Press 97. Gerstenfeld, L. C., & Shapiro, F. D. 1996 Expression of bone specific genes by h ypertrophic chondrocytes: Implications of the complex functions of the hypertrophic chondrocyte during endochondral bone development. Journal of Cellular Biochemistry, 62 (1), 1 9. (doi: 10.1002/(sici)1097 4644(199607)62:1<1::aid jcb1>3.0.co;2 x) 98. Fujisawa, K., & Kuboki, Y. 1992 Affinity of bone sialoprotein and several other bone and dentin acidic proteins to collagen fibrils. Calcified Tissue International, 51 ( 6), 438 442.( doi: 10.1007/bf00296677) 99. Maillard, C., Malaval, L., & Delmas, P. D. 1992 Immunologic al screening of SPARC/Osteonectin in nonmineralized tissues. Bone, 13 (3), 257 264. (doi: 10.1016/8756 3282(92)90206 c) 100. 100. Guweidhi, A., Kleeff, J., Adwan, H., Giese, N. A., Wente, M. N., Giese, T., Friess, H. 2005 Osteonectin influences growth and invas ion of pancreatic cancer cells. Annals of Surgery 242 (2), 224 234. 101. Rodrguez, I. R., Moreira, E. F., Bok, D., & Kantorow, M. 2000 Osteonectin/SPARC Secreted by RPE and Localized to the Outer Plexiform Layer of the Monkey Retina Investigative Ophthalmolo gy & Visual Science 41 (9), 2438 2444. 102. Viguet Carrin, S., Garnero, P., & Delmas, P. 2006 The role of collagen in bone strength. Osteoporosis International 17 (3), 319 336. (doi: 10.1007/s00198 005 2035 9 ) 103. Goustin, A. S., Leof, E. B., Shipley, G. D., & Mos es, H. L. 1986 Growth Factors and Cancer. Cancer Research 46 (3), 1015 1029. 104. Mundy, G. R. 1996 Regulation of Bone Formation by Bone Morphogenetic Proteins and Other Growth Factors. Clinical Orthopaedics and Related Research 324 24 27. 105. Riley, E. H., Lan e, J. M., Urist, M. R., Lyons, K. M., & Lieberman, J. R. 1996 Bone Morphogenetic Protein 2: Biology and Applications. Clinical Orthopaedics and Related Research 324 39 46. 106. Bostrom, M. P. G., & Camacho, N. P. 1998 Potential Role of Bone Morphogenetic Pr oteins in Fracture Healing. Clinical Orthopaedics and Related Research 355 S274 S282.
156 107. Heckman, J. D., Boyan, B. D., Aufdemorte, T. B., & Abbott, J. T. 1991 The use of bone morphogenetic protein in the treatment of non union in a canine model. The Journa l of bone and joint surgery. American Volume 73 (5), 750 764. 108. Sandhu, H. S., Khan, S. K., Suh, D. S., & Boden, S. B. 2001 Demineralized bone matrix, bone morphogenetic proteins, and animal models of spine fusion: an overview. European Spine Journal 10 (0) S122 S131. (doi: 10.1007/s005860100303) 109. Reddi, H. 2000 Interplay between bone morphogenetic proteins and cognate binding proteins in bone and cartilage development: noggin, chordin and DAN. Arthritis Research & Therapy 3 (1), 1 5. (doi: 10.1186/ar133) 110. Israel, D. I., Nove, J., Kerns, K. M., Kaufman, R. J., Rosen, V., Cox, K. A., & Wozney, J. M. 1996 Heterodimeric Bone Morphogenetic Proteins Show Enhanced Activity In Vitro and In Vivo. Growth Factors 13 (3 4), 291 300. (doi:10.3109/08977199609003229) 111. Cho K. W. Y., & Blitz, I. L. 1998 BMPs, Smads and metalloproteases: extracellular and intracellular modes of negative regulation. Current Opinion in Genetics & Development, 8 (4), 443 449. (doi: 10.1016/s0959 437x(98)80116 0 ) 112. Li, B. 2008 Bone Morphogenetic Protein Smad Pathway as Drug Targets for Osteoporosis and Cancer Therapy. Endocrine, Metabolic & I mmune D isorders drug targets 8 (3), 208 219. (doi: 10.2174/187153008785700127) 113. Kawabata, M., Im amura, T., & Miyazono, K. 1998 Signal transduction by bone morp hogenetic proteins. Cytokine &; Growth Factor Reviews 9 (1), 49 61. (doi: 10.1016/s1359 6101(97)00036 1) 114. Haydon, R. C., Luu, H. H., & He, T. C. 2007 Osteosarcoma and Osteoblastic Differentiation: A New Perspective on Oncogenesis. Clinical Orthopaedics an d Related Research 454 237 246 (210.1097/BLO.1090b1013e31802b31683c.) 115. Nakashima, M. 1994 Induction of Dentin Formation on Canine Amputated Pulp by Recombinant Human Bone Morphogenetic Proteins (BMP) 2 and 4. Journal of Dental Research 73 (9), 1515 152 2. (doi: 10.1177/00220345940730090601) 116. Ahrens, M., Ankenbauer, T., Schroder, D., Hollnagel, A., Mayer, H., & Gross, G. 1993 Expression of human bone morphogenetic proteins 2 or 4 in murine mesenchymal progenitor C3H10T1/2 cells induces differentiation in to distinct mesenchymal cell lineages. DNA and C ell B iology 12 (10), 871 880 117. Dimitriou, R., Tsiridis, E., & Giannoudis, P. V. 2005 Current concepts of molecular aspects of bone healing. Injury 36 (12), 1392 1404. doi: 10.1016/j.injury.2005.07.019
157 118. Schi ndeler, A., McDonald, M. M., Bokko, P., & Little, D. G. 2008 Bone remodeling during fracture repair: The cellular picture. Seminars in Cell & Developmental Biology 19 (5), 459 466. (doi: 10.1016/j.semcdb.2008.07.004) 119. Stewart, M. G. 20 05. Head, face, and neck trauma: Comprehensive management pp. 89 90. New York, New York: Thieme. 120. LaStayo, P. C., Winters, K. M., & Hardy, M. 2003 Fracture healing: Bone healing, fracture management, and current concepts related to the hand. Journal of Hand Therapy 16 (2), 8 1 93. (doi: 10.1016/s0894 1130(03)80003 0) 121. Lieberman, J. R., & Friedlaender, G. E. 2005 Bone regeneration and repair biology and clinical applications, Retrieved from http://pub lic.eblib.com/EBLPublic/PublicView.do?ptiID=603590 122. Schenk, R. 1978 Histology of Fracture Repair and Non Union Bulletin of the Swiss Association for Study of Internal Fixation, pp. 14. Bern Switzerland : Swiss Association for Study of Internal Fixation 123. R ahn, B. A. 1971 Primary Bone Healing : An Experimental Study in the Rabbit. Journal of bone and joint surgery American volume 53 (4), 783. 124. Draenert, Y. 1980 Gap healing of compact bone. Scanning electron microscopy 4 103 111. 125. Claes, L., Recknagel, S., Ig natius, A. 2012 Fracture healing under healthy and inflammatory conditions Nat Rev Rheumatol 8 (3), 133 143. (doi: 10.1038/nrrheum.2012.1 ) 126. Perren, S. M. 1979 Physical and biological aspects of fracture healing with special reference to internal fixation. Clinical Orthopaedics and Related Research 138 175 196 127. Shapiro, F. 2008 Bone development and its relation to fracture repair. The role of mesenchymal osteoblasts and surface osteoblasts. European cells & materials 15 53 76 128. Weatherholt, A. M., Fuchs R. K., & Warden, S. J. 2012 Specialized Connective Tissue: Bone, the Structural Framework of the Upper Extremity. Journal of Hand Therapy 25 (2), 123 132. (doi: 10.1016/j.jht.2011.08.003) 129. Zioupos, P. 2001Ageing Human Bone: Factors Affecting its Biomecha nical Properties and the Role of Collagen. Journal of Biomaterials Applications 15 (3), 187 229. (doi: 10.1106/5juj tfj3 jvva 3rj0)
158 130. Ozaki, A., Tsunoda, M., Kinoshita, S., & Saura, R. 2000 Role of fracture hematoma and periosteum during fracture healing in rats: interaction of fracture hematoma and the periosteum in the initial step of the healing process. Journal of Orthopaedic Science 5 (1), 64 70. (doi: 10.1007/s007760050010 ) 131. Schmidt Bleek, K., Schell, H., Lienau, J., Schulz, N., Hoff, P., Pfaff, M., Duda, G. 2012 nitial immune reaction and angiogenesis in bone healing. Journal of Tissue Engineering and Regenerative Medicine (doi: 10.1002/term.1505) 132. Kalfas, I. H. 2001 Principles of bone healing. Neurosurgical Focus 10 (4), 1 4. (doi: 10.3171/foc.200 22.214.171.124) 133. Hoppenfeld, S., & Murthy, V. L. 2000 Treatment and rehabilitation of fractures pp. 303. Philadelphia Pa : Lippincott Williams & Wilkins. 134. Marzona, L., & Pavolini, B. 2009 Play and players in bone fracture healing match. Clinical cases in minera l and bone metabolism: the official journal of the Italian Society of Osteoporosis, Mineral Metabolism, and Skeletal Diseases 6 (2), 159 162. 135. McKibbin, B. 1978 The biology of fracture healing in long bones. The Journal of bone and joint surgery British v olume 60 B(2), 150 162. 136. Sen, M. K., & Miclau, T. 2007 Autologous iliac crest bone graft: Should it still be the gold standard for treating nonunions? Injury 38 (1, Supplement), S75 S80. (doi: 10.1016/j.injury.2007.02.012.) 137. Campanacci, M., & Costa, P. 197 9 Total resection of distal femur or proximal tibia for bone tumours. Autogenous bone grafts and arthrodesis in twenty six cases. Journal of Bone & Joint Surgery British Volume 61 B(4), 455 463. 138. Giannoudis, P. V., Dinopoulos, H., & Tsiridis, E. 2005 Bone substitutes: An update. Injury 36 (3, Supplement), S20 S27. (doi: 10.1016/j.injury.2005.07.029) 139. Joshi, A., & Kostakis, G. C. 2004 An investigation of post operative morbidity following iliac crest graft harvesting. Br Dent J 196 (3), 167 171. (10.1038/sj .bdj.4810945) 140. Nguyen, H., Morgan, D., & Forwood, M. 2007 Sterilization of allograft bone: effects of gamma irradiation on allograft biology and biomechanics. Cell and Tissue Banking 8 (2), 93 105. (doi: 10.1007/s10561 006 9020 1) 141. Vaz, K., Verma, K., Protop saltis, T., Schwab, F., Lonner, B., & Errico, T. 2010 Bone grafting options for lumbar spine surgery: a review examining clinical efficacy and complications. SAS Journal 4 (3), 75 86.( doi: 10.1016/j.esas.2010.01.004 )
159 142. Greenwald, A. S., Boden, S. D., Gold berg, V. M., Khan, Y., Laurencin, C. T., Rosier, R. N., & American Academy of Orthopaedic Surgeons. The Committee on Biological, I. 2001 Bone graft substitutes: facts, fictions, and applications. The Journal of bone and joint surgery American volume, 83 A Suppl 2 Pt 2, 98 103. 143. Bostrom, M. P. G., & Seigerman, D. A. 2005 The Clinical Use of Allografts, Demineralized Bone Matrices, Synthetic Bone Graft Substitutes and Osteoinductive Growth Factors: A Survey Study. HSS Journal 1 (1), 9 18. ( doi: 10.1007/s1142 0 005 0111 5 ) 144. Harakas, N. K. 1984 Demineralized Bone Matrix Induced Osteogenesis. Clinical Orthopaedics and Related Research 188 239 251. 145. Wolfinbarger, L., L. M. Eisenlohr, et al. 2008. Demineralized Bone Matrix: Maximizing New Bone Formation for Succe ssful Bone Implantation Musculoskeletal Tissue Regeneration: Biological Materials and Methods pp. 93 117. Totowa, NJ: Humana Press 146. Goebel, J. A., & Jacob, A. 2005 Use of Mimix hydroxyapatite bone cement for difficult ossicular reconstruction. Otolaryngolo gy Head and Neck Surgery 132 (5), 727 734. (doi: 10.1016/j.otohns.2005.01.023 ) 147. Sottosanti, J. S. 1995 Calcium sulfate aided bone regeneration: a case report. Periodontal clinical investigations : O fficial publication of the Northeastern Society of Perio dontists, 17 (2), 10 15. 148. Beuerlein, M. J. S., & McKee, M. D. 2010 Calcium Sulfates: What Is the Evidence? Journal of Orthopaedic Trauma 24 S46 S51 (10.1097/BOT.1090b1013e3181cec1048e.) 149. Glazer, P. A., Spencer, U. M., Alkalay, R. N., & Schwardt, J. 2001 I n vivo evaluation of calcium sulfate as a bone graft substitute for lumbar spinal fusion. The Spine Journal 1 (6), 395 401. ( doi: 10.1016/s1529 9430(01)00108 5 ) 150. Petruskevicius, J., Nielsen, S., Kaalund, S., Knudsen, P. R., & Overgaard, S. 2002 No effect o f Osteoset a bone graft substitute, on bone healing in humans: A prospective randomized double blind study. Acta Orthopaedica 73 (5), 575 578. (doi: doi:10.1080/000164702321022875 ) 151. Barrack, R. L. 2005 Bone Graft Extenders, Substitutes, and Osteogenic Proteins. The Journal of Arthroplasty 20 94 97. (doi: 10.1016/j.arth.2005.03.025) 152. Schoelles K, Snyder D, Kaczmarek J, et al. 2005 The Role of Bone Growth Stimulating Devices and Orthobiologics in Healing Nonunion Fractures Agency for Healthcare Research and Quality. Retrieved from http://www.cms.hhs.gov/determinationprocess/downloads/id29TA.pdf
160 153. Katagiri, T., Yamaguchi, A., Komaki, M., Abe, E., Takahashi, N., Ikeda, T., Suda, T. 1994 Bone morphogenetic protein 2 converts the differentiation pathway of C2C12 myoblasts into the osteoblast lineage. The Journal of Cell Biology 127 (6), 1755 1766. (doi: 10.1083/jcb.127.6.1755 ) 154. Cheng, H., Jiang, W., Phillips, F. M., Haydon, R. C., Peng, Y., Zhou, L., He, T. C. 2003 Osteogenic activity of the fourteen types of human bone morphogenetic proteins (BMPs). The Journal of bone and joint surgery American volume 85 A(8), 1544 1552. 155. Bauer, T. W., & Muschler G. F. 2000 Bone Graft Materials: An Overview of the Basic Science Clinical Orthopaedics and Related Research 371 10 27. 156. Schmitt, J. M., Hwang, K., Winn, S. R., & Hollinger, J. O. 1999 Bone morphogenetic proteins: An update on basic biology and clinic al relevance. Journal of Orthopaedic Research 17 (2), 269 278. (doi: 10.1002/jor.1100170217) 157. Walker DH, Wright NM. 2002 Bone morphogenic proteins and spinal fusion. Neurosurg Focus 1 3 (6 ) ,1 13. 158. Urist, M. R., Silverman, B. F., Dring, K., DubuC, F. L., & Ro senberg, J. M. 1967 24 The Bone Induction Principle. Clinical Orthopaedics and Related Research 53 243 284. 159. Wozney, J., Rosen, V., Celeste, A., Mitsock, L., Whitters, M., Kriz, R., Wang, E. 1988 Novel regulators of bone formation: molecular clones and activities. Science 242 (4885), 1528 1534. (doi: 10.1126/science.3201241) 160. Pagani, F., Francucci, C. M., & Moro, L. 2005 Markers of bone turnover: biochemical and clinical perspectives. [Review]. J Endocrinol Invest 28 (10), 8 13. 161. Lewandrowski, K. U., Tomf ord, W. W., Michaud, N. A., Schomacker, K. T., & Deutsch, T. F. 1997 An Electron Microscopic Study on the Process of Acid Demineralization of Cortical Bone. Calcified Tissue International 61 (4), 294 297. (doi: 10.1007/s002239900338) 162. Pietrzak, W. S., Perns S. V., Keyes, J., Woodell May, J., & McDonald, N. M. 2005 Demineralized Bone Matrix Graft: A Scientific and Clinical Case Study Assessment. The Journal of Foot and Ankle Surgery 44 (5), 345 353. (doi: 10.1053/j.jfas.2005.07.006) 163. Zhang, M., Powers, R. M. Jr., & Wolfinbarger, L., Jr. 1997 Effect(s) of the demineralization process on the osteoinductivity of demineralized bone matrix. [Research Support, Non U S Gov't]. J Periodontol 68 (11), 1085 1092.
161 164. Damien, C. J., Parsons, J. R., Prewett, A. B., Rietvel d, D. C., & Zimmerman, M. C. 1994 Investigation of an organic delivery system for demineralized bone matrix in a delayed healing cranial defect model. Journal of Biomedical Materials Research 28 (5), 553 561. (doi: 10.1002/jbm.820280505) 165. Lye, K. W., Death erage, J. R., & Waite, P. D. 2008 The Use of Demineralized Bone Matrix for Grafting During Le Fort I and Chin Osteotomies: Techniques and Complications. Journal of Oral and Maxillofacial Surgery 66 (8), 1580 1585. (doi: 10.1016/j.joms.2007.12.003.) 166. Han, B. Tang, B., & Nimni, M. E. 2003 Quantitative and sensitive in vitro assay for osteoinductive activity of demineralized bone matrix. Journal of Orthopaedic Research 21 (4), 648 654. (doi: 10.1016/s0736 0266(03)00005 6 ) 167. Liu, L., Wang, Y., Shen, X., & Fang, Y. e. 2005 Preparation of chitosan g polycaprolactone copolymers through ring caprolactone onto phthaloyl protected chitosan. Biopolymers 78 (4), 163 170. (doi: 10.1002/bip.20261) 168. Agarwal, S., & Speyerer, C. 2010 Degradable ble nds of semi crystalline and amorphous branched poly(caprolactone): Effect of microstructure on blend properties. Polymer 51 (5), 1024 1032. (doi: 10.1016/j.polymer.2010.01.020 ) 169. Hutmacher, D. W., Schantz, T., Zein, I., Ng, K. W., Teoh, S. H., & Tan, K. C. 2001 Mechanical properties and cell cultural response of polycaprolactone scaffolds designed and fabricated via fused deposition modeling. Journal of Biomedical Materials Research 55 (2), 203 216. (doi: 10.1002/1097 4636(200105)55:2<203::aid jbm1007>3.0.co ;2 7 ) 170. Priya, S. G., Jungvid, H., & Kumar, A. 2008 Skin tissue engineering for tissue repair and regeneration. Tissue engineering Part B, Reviews 14 (1), 105 118. 171. caprolactone) PCL Scaffolds for Tissue Engineering Applications. Retrieved from http://ethesis.nitrkl.ac.in/1458/1/final_thesis_with_page_nu.pdf 172. Ginde, R. M., & Gupta, R. K. 1987 In vitro chemica l degradation of poly(glycolic acid) pellets and fibers. Journal of Applied Polymer Science 33 (7), 2411 2429.(doi: 10.1002/app.1987.070330712) 173. Goepferich, A., Karydas, D., Langer, R. 1995 Predicting drug release from cylindric polyanhydride matrix discs", European J. of Pharm. and Biopharm 41 (2): 81 87. 174. Bergsma, J. E., de Bruijn, W. C., Rozema, F. R., Bos, R. R. M., & Boering, G. (1995). Late degradation tissue response to poly(l lactide) bone plates and screws. Biomaterials 16 (1), 25 31. (doi: 10.1016/0 142 9612(95)91092 d)
162 175. Vert, M. 2009 Degradable and bioresorbable polymers in surgery and in pharmacology: beliefs and facts. Journal of Materials Science: Materials in Medicine 20 (2), 437 446.( doi: 10.1007/s10856 008 3581 4) 176. Shimao, M. 2001 Biodegradation of plastics. Current Opinion in Biotechnolo gy, 12 (3), 242 247. (doi: 10.1016/s0958 1669(00)00206 8) 177. Tamjid, E., Bagheri, R., Vossoughi, M., & Simchi, A. (2011). Effect of TiO2 morphology on in vitro bioactivity of polycaprolactone/TiO2 nanocomposites. Mat erials Letters 65 (15 16), 2530 2533. (doi: 10.1016/j.matlet.2011.05.037) 178. Commandeur, S., Van Beusekom, H. M. M., & Van Der Giessen, W. J. 2006 Polymers, Drug Release, and Drug Eluting Stents. Journal of Interventional Cardiology 19 (6), 500 506. (doi: 10. 1111/j.1540 8183.2006.00198.x) 179. Fahmy, T. M., Fong, P. M., Goyal, A., & Saltzman, W. M. 2005 Targeted for drug delivery. Materials Today 8 (8), 18 26. (doi: 10.1016/s1369 7021(05)71033 6) 180. Michea, L., Ferguson, D. R., Peters, E. M., Andrews, P. M., Kirby, M. R., & Burg, M. B. 2000 Cell cycle delay and apoptosis are induced by high salt and urea in renal medullary cells. American Journal of Physiology Renal Physiology 278 (2), F209 F218. 181. Han, B., Yang, Z., & Nimni, M. 2005 Effects of moisture and temperat ure on the osteoinductivity of demineralized bone matrix. Journal of Orthopaedic Research 23 (4), 855 861. (doi: 10.1016/j.orthres.2004.11.007) 182. Kokubo, T., & Takadama, H. 2006 How useful is SBF in predicting in vivo bone bioactivity? Biomaterials 27 (15), 2907 2915.(doi: 10.1016/j.biomaterials.2006.01.017) 183. Nagano, M., Nakamura, T., Kokubo, T., Tanahashi, M., & Ogawa, M. 1996 Differences of bone bonding ability and degradation behaviour in vivo between amorphous calcium phosphate and highly crystalline hydr oxyapatite coating. Biomaterials 17 (18), 1771 1777. (doi: 10.1016/0142 9612(95)00357 6) 184. Albrektsson, T. A., & Johansson, C. J. 2001 Osteoinduction, osteoconduction and osseointegration. European Spine Journal 10 (0 ), S96 S101. (doi: 10.1007/s005860100282 ) 185. Johnson, R., Harrison, D., Tucci, M., Tsao, A., Lemos, M., Puckett, A., Benghuzzi, H. 1997 Fibrous capsule formation in response to ultrahigh molecular weight polyethylene treated with peptides that influence adhesion. Biomedical sciences instrumentation 34 47 52. 186. Webster, T. J. 2001 Nanophase ceramics: The future orthopedic and dental implant material Advances in Chemical Engineering 27 125 166
163 187. Gotman, I. 1997 Characteristics of metals used in implants. Journal of endourology / Endourological Societ y 11 (6), 383 389. 188. Nelson, C. L., McLaren, A. C., McLaren, S. G., Johnson, J. W., & Smeltzer, M. S. 2005 Is Aseptic Loosening Truly Aseptic? Clinical Orthopaedics and Related Research, 437 25 30 (10.1097/1001.blo.0000175715.0000168624.0000175713d.) 189. Kim H. M., Himeno, T., Kokubo, T., & Nakamura, T. 2005 Process and kinetics of bonelike apatite formation on sintered hydroxyapatite in a simulated body fluid. Biomaterials 26 (21), 4366 4373.(doi: 10.1016/j.biomaterials.2004.11.022 ) 190. Li, P., Kangasniemi, I. de Groot, K., & Kokubo, T. 1994 Bonelike Hydroxyapatite Induction by a Gel Derived Titania on a Titanium Substrate. Journal of the American Ceramic Society 77 (5), 1307 1312. (doi: 10.1111/j.1151 2916.1994.tb05407.x) 191. Dias, A. G., Gibson, I. R., Santos, J D., & Lopes, M. A. 2007 Physicochemical degradation studies of calcium phosphate glass ceramic in the CaO P2O5 MgO TiO2 system. Acta Biomaterialia 3 (2), 263 269. (doi: 10.1016/j.actbio.2006.09.009) 192. Zyman, Z., Weng, J., Liu, X., Zhang, X., & Ma, Z. 1993 Amorphous phase and morphological structure of hydroxyapatite plasma coatings. Biomaterials 14 (3), 225 228. (doi: 10.1016/0142 9612(93)90027 y) 193. Schweizer, S., & Taubert, A. 2007 Polymer Controlled, Bio Inspired Calcium Phosphate Mineralization from Aque ous Solution. Macromolecular Bioscience 7 (9 10), 1085 1099.(doi: 10.1002/mabi.200600283) 194. Lievremont, M., Potus, J., & Guillou, B. 1982 Use of Alizarin Red S for Histochemical Staining of Ca2 + in the Mouse; Some Parameters of the Chemical Reaction in vitro Cells Tissues Organs 114 (3), 268 280. 195. Li, X., Shi, J., Dong, X., Zhang, L., & Zeng, H. 2008 A mesoporous bioactive glass/polycaprolactone composite scaffold and its bioactivity behavior. Journal of Biomedical Materials Research Part A 84 A(1), 84 91. (d oi: 10.1002/jbm.a.31371) 196. Han, B., Yang, Z., & Nimni, M. 2008 Effects of gamma irradiation on osteoinduction associated with demineralized bone matrix. Journal of Orthopaedic Research 26 (1), 75 82. (doi: 10.1002/jor.20478) 197. Figueiredo, M., Cunha, S., Martins, G., Freitas, J., Judas, F., & Figueiredo, H. 2011 Influence of hydrochloric acid concentration on the demineralization of cortical bone. Chemical Engineering Research and Design 89 (1), 116 124.(doi: 10.1016/j.cherd.2010 .04.013)
164 BIOGRAPHICAL SKETCH Titilayo Moloye was born to Dr. Olugbemi Moloye and Mrs. Yinka Moloye. She has two sisters, Olajompo Moloye and Dolapo Moloye. She was born in Ibadan, Nigeria but grew up in Tallahassee, FL. Upon graduation, she will be movin g to Alexandria, VA where she will begin her career as a patent examiner with the United States Patent and Trademark Office.