UNIVERSITY OF CINCINNATI

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1 UNIVERSITY OF CINCINNATI Date: I,, hereby submit this work as part of the requirements for the degree of: in: It is entitled: This work and its defense approved by: Chair:

2 Enhancing Tissue Engineered Tendon Repair by Optimizing In Vitro Culture Conditions A dissertation submitted to the Division of Research and Advanced Studies of the University of Cincinnati in partial fulfillment of the requirements for the degree of DOCTOR OF PHILOSOPHY (Ph.D.) in the Department of Biomedical Engineering of the College of Engineering 2007 by Victor Sanjit Nirmalanandhan M.S., University of Cincinnati, 2004 B.Tech., Regional Engineering College, Calicut, India, 2001 Committee Chair: David L. Butler, Ph.D.

3 ABSTRACT This dissertation describes experiments designed to test aspects of the governing hypothesis that tissue-engineered constructs containing mesenchymal stem cells can be used in models of tendon injury to improve repair biomechanics and structure. As part of a team, I tested the working hypothesis that optimizing the culture conditions enhances the construct stiffness and the repair biomechanics. The first two studies were designed to identify the appropriate cell-to-collagen ratio and the construct length. I first determined that increasing cell-to-collagen ratio increases the contraction kinetics of the construct such that high ratios result in premature failure of the constructs. Increasing collagen concentration level was also more effective in controlling contraction kinetics than decreasing cell density. The importance of collagen scaffold led to my next study in which I was able to create stiffer and more aligned constructs by increasing the length of the construct in accordance with St. Venant s principle. These results also suggested that the structure of the construct and its biomechanical properties are positively correlated. As one strategy to further improve the biomechanical properties of these constructs, I then studied the effect of mechanical stimulation on the biomechanical properties of cell-gel and cell-sponge constructs. Results from this study indicated that the response of MSCs to mechanical stimulus is scaffold material specific. I next examined the relative importance of scaffold material, construct length and mechanical stimulation in a multi-factorial experiment. As a result of the synergistic effects of these factors, longer, stimulated cell-sponge constructs showed a significantly higher in vitro linear stiffness compared to the other 7 combinations studied. ii

4 I then performed in vitro and in vivo studies to begin to translate my research to preclinical utility. Using a statististical optimization strategy called Response Surface Methodology, I first optimized three components (peak strain, cycle number, and cycle repetition) of the mechanical stimulus while controlling cycle frequency (1 Hz), rise and fall times (25% and 17% of the period, respectively), hours of stimulation/day (8 hours/day) and total time of stimulation (2 weeks). The results from this study indicated that constructs stimulated with 2.4% strain, 3000 cycles/day and one cycle repetition produced the stiffest constructs for later surgery. Finally, I examined whether these in vitro improvements in linear stiffness achieved through optimization indeed translate into increased in vivo repair biomechanics. Limitations in the integrity of the sponges caused me to first stiffen the scaffold material (collagen sponge) before construct creation using dehydrothermally (DHT) crosslinking. Surprisingly, I found that combining DHT crosslinking with mechanical stimulation adversely affected the repair biomechanics, presumably due to the influence of crosslinking. Although this dissertation has demonstrated the potential benefits of increasing construct length and collagen concentration and providing mechanically stimulation, the studies generated new challenges that must be overcome. We must identify ways to stiffen the scaffold and avoid premature failure of constructs due to poor mechanical integrity. Scaffolds must be designed with sufficient mechanical integrity and biologic activity to accomplish our goal of a functionally efficacious tendon repair. Future studies must identify scaffold materials that control/modulate MSC differentiation in vitro and in vivo so that appropriate cellular phenotypes can be produced and extracellular matrices synthesized. iii

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6 ACKNOWLEDGEMENTS 1 Though this dissertation is an individual work, I could never have completed this without the help, support, guidance and efforts of a lot of people. First, I would like to thank my advisor Dr. David Butler for his continuous counsel and supervision throughout the investigation and writing of this dissertation. His enthusiasm and unlimited zeal have been major driving forces through my graduate career at the University of Cincinnati. I will always be grateful. Thanks and gratitude to Dr. Jason Shearn for serving on my dissertation committee and his invaluable contributions to this dissertation and to my professional growth. Special thanks to Jason for teaching me mechanical testing and data collection. I also extend my sincere thanks to Dr. Balakrishna Haridas for being a member of my dissertation committee and for the insightful discussions. Though our personal interaction was limited, I have felt Bala s support. I would like to take this opportunity to thank Drs. Greg Boivin and Marc Galloway for sharing their expertise in rabbit surgery. I would also like to acknowledge and thank Drs. M.B. Rao and Marty Levy for the help they have given with statistics. Drs. Matthew Dressler, at Dordt College in Iowa and Natalia Juncosa-Melvin deserve a very special mention for helping and teaching me all about tendon tissue engineering and data collection. A special word of thanks goes to Natalia for the long and enthusiastic discussions we have had about the projects from this dissertation and her help with PCR analysis. 1 This work was supported by a grant from the National Institutes of Health (AR 46574) given to the University of Cincinnati, by Cincinnati Sports Medicine Research and Education Foundation, and by a Merit Review grant from the Veteran s Administration to Dr. Gregory Boivin. v

7 Special thanks to Cindi Gooch for teaching me cell culture techniques and helping during bone marrow harvests and surgeries. I also like to thank Drs. William Ball, Edward Grood and Carl Huether (Biology) for serving as my Preparing Future Faculty program mentors. Their valuable advice definitely helped me improve my teaching skills. I am also grateful to my fellow graduate students, faculty and staff at the Department of Biomedical Engineering. Special thanks to Drs. Shawn Hunter, Joel Collier, Daria Narmoneva, and Jeff Johnson as well as Kumar Chockalingam, Shun Yoshida, Nisha Sipes, Matt Stonecash, Tatiana Mavridas, Abhishek Jain, Nat Dyment, Saurabh Datta, Denise Smith, Vinod Kaimal, Balaji Kalyanaraman, Heather Powell and Kirsten Kinneberg. I cannot say thanks enough to Linda Moeller, Mike Sanderson, Lori Beth Derenski, and Denis Bailey. Their help and friendship have made this bearable. I would like to thank Dr. Michael Sacks and David Merryman at the University of Pittsburgh for help with the Small Angle Light Scattering measurements. My special thanks to all of my friends, especially those at Church of the Saviour- United Methodist. Their prayers and support have carried me through. I also extend my gratitude to all my Sri Lankan friends, especially to Raj and Malathy for making my stay in Cincinnati an enjoyable one. My very special thanks to the one person whom I owe everything I am today, my cousin, Roshan Gunaratnam. His unwavering faith and confidence in my abilities and in me is what has shaped me to be the person I am today. vi

8 Last, but certainly not the least, I would like to express my special appreciation to my parents, sister and brother for their love, encouragement and support that helped me achieved this educational goal. vii

9 Imagination is more important than knowledge. Knowledge is limited. Imagination encircles the world. Albert Einstein In an Interview by George Sylvester Viereck, for the October 26, 1929 issue of The Saturday Evening Post viii

10 ORGANIZATION OF DISSERTATION This dissertation is organized as a collection of several journal manuscripts, each with the customary sections, including a summary, materials and methods, results and discussion. The presentation of these manuscripts is integrated to produce a coherent dissertation with well-defined objectives and clearly stated conclusions. Chapter 1 begins with an introduction that includes description of the significance, specific aims and hypotheses of the current work. Chapter 2 is a literature review that gives a background of soft tissue injuries and the different approaches for treatment. A brief review of various tissue engineering concepts is also presented in this Chapter. Chapters 3 through 8 present six journal manuscripts. The first manuscript presented in Chapter 3, describes the effect of cell seeding density and collagen concentration on contraction kinetics of MSC-seeded collagen constructs. The next Chapter examine whether more aligned constructs be created by increasing the length of the constructs and hence reducing the end effects acting on the bulk of the constructs (St. Venant s principle) during contraction. The third manuscript in Chapter 5 describes an in vitro study that investigates the influence of two scaffold materials, type 1 collagen gel and sponge on the response of MSCs to mechanical stimulation. Chapter 6 explores the individual and the combined effects of three factors; scaffold material, construct length and mechanical stimulation on the in vitro stiffness of the tissue engineered construct. Chapter 7 optimizes the peak strain, cycle number per day and the cycle repetition of the mechanical stimulus to further improve the in vitro stiffness of the construct. The in vivo study described in Chapter 8 explores whether constructs exposed to the optimized ix

11 stimulus in culture, when placed in a patellar tendon defect site, improve the repair outcome compared to constructs exposed to non optimized stimulus. Finally, Chapters 9 and 10 present discussion of the six studies and recommendations for future research in this field. x

12 Table of Contents Abstract.ii Acknowledgments.v Organization of Dissertation... ix List of Tables.. xii List of Figures....xiii Chapter Chapter 2 - Literature Review Tendon Structure Frequency and Significance of Musculoskeletal Injuries Conventional Clinical Treatments Prosthetic Ligament and Tendon Replacement Tissue Engineering as a Possible Alternative to Conventional Repair Mesenchymal Stem Cells Functional Tissue Engineering Mechanical Stimulation to Control Phenotype In Vitro. 30 Chapter 3 - Effect of Cell Seeding Density and Collagen Concentration on Contraction Kinetics of MSC-Seeded Collagen Constructs Introduction Methods Results Discussion Appendix Chapter 4 - Effect of Length of the Engineered Tendon Construct on its Structure-Function Relationships in Culture Introduction Methods Results Discussion. 63 Chapter 5 - Mechanical Stimulation of Tissue Engineered Tendon Constructs; Effect of Scaffold Materials Introduction Methods Results Discussion. 78 xi

13 Chapter 6 - Effect ff Scaffold Material, Construct Length and Mechanical Stimulation on the In Vitro Stiffness of the Engineered Tendon Construct Introduction Methods Results Discussion. 95 Chapter 7 - Optimizing Mechanical Stimulus to Improve In vitro Biomechanical Properties of Tissue Engineered Tendon Constructs Introduction Experimental Design & Methods Results Discussion Chapter 8 - Combined Effects of Scaffold Stiffening and Mechanical Preconditioning Cycles on Construct Biomechanics and Gene Expression and Tendon Repair Biomechanics Introduction Experimental Design Materials Methods Results Discussion Chapter 9 - Discussion and Conclusions Chapter 10 - Recommendations Cellular Aspects of Tendon Tissue Engineering Scaffold Materials Mechanical Stimulation Chemical Stimulation Other Suggestions Bibliography..161 xii

14 List of Tables Table 3.1. Table 3.2. Table 3.3. Normalized area of HK, LK, HM and LM constructs at 0, 4, 8, 16, 32, 72, 120 and 168 hrs 42 Contrast analysis of treatment means.43 Model parameters for different treatment combinations 46 Table 4.1. Biomechanical properties of short and long constructs. 60 Table 5.1. Biomechanical properties of the cell-gel & cell-sponge constructs.. 70 Table 6.1. Table 6.2. Experimental design to study the combined effects of scaffold material, length, and mechanical stimulation 88 Comparison of long and short constructs...93 Table 6.3. Comparison of cell-gel and cell-sponge constructs Table 6.4. Least significant difference (LSD) multiple comparisons. 95 Table 7.1. Treatments and levels for optimizing the mechanical stimulus. 106 Table 7.2. Linear stiffness of the construct from iteration Table 7.3. Linear stiffness of the construct from iteration 2 & Table 8.1. Table 8.2. Dimensions or the load-related biomechanical properties of DHT crosslinked constructs in culture Gene expression for col 1, col 3, decorin, or fibronectin relative to GAPDH of DHT crosslinked constructs 131 Table 8.3. Dimensions of DHT repairs Table 8.4. Biomechanical properties of DHT repair tissues 133 xiii

15 List of Figures Figure 1.1. Digital and Small Angle Light Scattering images of a construct 4 Figure 2.1. Structural hierarchy of collagen. 12 Figure 2.2. The structural organization of tendon 14 Figure 2.3. ACL reconstruction using central third portion of patellar tendon 18 Figure 3.1. Silicone dish consisting of 4 wells with 2 posts 35 Figure 3.2. Figure 3.3. Digital images of HK, LK, HM and LM constructs at 4, 8, 16, 32, 72, 120 and 168 hours 41 Normalized area of HK, LK, HM & LM constructs as a function of time.. 43 Figure 3.4. Normalized area of HM & L5K constructs as a function of time.. 44 Figure 4.1. Digital and Small Angle Light Scattering images of a construct.. 54 Figure 4.2. Short and long silicone dishes 56 Figure 4.3. Biomechanical properties of short and long constructs. 61 Figure 4.4. SALS images of short and long constructs.62 Figure 4.5. Correlation between orientation index and construct modulus.. 63 Figure 5.1. Silicone dish consisting 4 wells with 2 posts. 72 Figure 5.2. Mechanical stimulation system.. 73 Figure 5.3. Linear modulus of stimulated and non stimulated constructs 77 Figure 5.4. Figure 6.1. Figure 6.2. Figure 6.3. Linear stiffness of stimulated and non stimulated constructs.77 Short and long constructs 88 Effects of length, scaffold material and mechanical stimulation on construct s linear stiffness.. 92 Least significant difference (LSD) multiple comparisons.93 xiv

16 Figure 7.1. A trapezoidal stimulus consists of multiple components Figure 7.2. Figure 7.3. Experimental design showing treatment levels..108 Effects of peak strain, cycles/day and cycle repletion on construct stinffness 109 Figure 7.4. Interaction between peak strain and cycles/day. 110 Figure 7.5. Figure 7.6. The initial response surface 111 The final response surface..112 Figure 8.1. Force elongation curves of DHT repairs and normal tendon. 134 xv

17 Chapter 1 Introduction This dissertation describes experiments designed to test aspects of the global hypothesis that tissue-engineered constructs containing mesenchymal stem cells can be used in models of tendon injury to improve repair biomechanics and structure [1]. Recent advances in cell culture and tissue engineering have allowed the concept of tissue repair and regeneration by cell transplantation to emerge. Many studies have utilized differentiated, specialized cell-transplantation [2-7], which normally requires creating secondary morbidity site(s) in a healthy tissue to harvest the autologous cells. By contrast, mesenchymal stem cell (MSC) transplantation offers the advantage of avoiding the creation of secondary morbidity sites by removing bone marrow [8-10]. Over the past 10 to 12 years, our group has demonstrated that suspending mesenchymal stem cells (MSCs) in type I collagen gels and implanting these constructs in defects in the patellar tendon (PT) produce significant improvements in repair biomechanics over natural repair [11]. The cells in the constructs align when contracted on a suture and when they were then implanted in Achilles tendon (AT) defects, the implants further enhance repair biomechanics and histology [12, 13]. Our group then lowered cell-seeding densities from higher levels (1, 4 and 8 x10 6 cells/ml) to 0.1 x10 6 cells/ml to avoid the ectopic bone formation that was observed in almost one third of the cases [14] and to reduce the premature failure of constructs in culture due to excessive contraction. 1

18 Unfortunately, choosing which cell concentration to use in the MSC-seeded transplants can be somewhat arbitrary. To our knowledge, no studies have justified the MSC concentration used in their transplants [13, 15]. Although Juncosa-Melvin et al [16] investigated the effects of decreasing cell-to-collagen ratio by 20 times from 0.8 to 0.04 cells/mg collagen on the cell-mediated repair process following implantation in vivo, their study did not identify the appropriate cell-to-collagen ratio that produces improved in vivo repair biomechanics without any excessive contraction in culture. Juncosa-Melvin et al [16] reported that constructs containing the highest cell-to-collagen ratios were damaged due to excessive contraction and no significant differences were observed in any structural or material properties or in histological appearance of the repair between the two lowest cell-to-collagen ratio conditions when constructs were implanted in vivo [16]. The fact that constructs with different cell-to-collagen ratios appear to contract differently gave rise to the research question, how does the cell-to-collagen ratio affect the contraction kinetics of these constructs? In order to address this question the following specific aim and hypothesis were developed. Specific Aim 1 (Chapter 3): Examine the effect of cell-to-collagen ratio on the contraction kinetics of MSC- seeded collagen constructs for up to 168 hours in culture. Hypothesis 1: Increasing cell-to-collagen ratio significantly increases contraction kinetics. Preliminary data showed highly significant differences in construct areas among all four ratios after 8 hours of contraction with the exception of the LK (0.08) vs. HM 2

19 (0.4) conditions. This similar pattern raised the second research question as to whether cell density or collagen concentration more influenced these events. To isolate these effects, the following specific aim and hypothesis was developed. Specific Aim 2 (Chapter 3): Compare the contraction kinetics of the HM construct to those of a new construct (L5K) with equivalent cell-to-collagen ratio (0.4) but half the cell density (500K cells/ml) and half the collagen concentration (1.3mg/ml). Hypothesis 2: Increasing cell density increases the contraction kinetics more than decreasing collagen concentration. The cell-assisted repairs exhibit 50% greater maximum force and stiffness at 12 weeks after surgery compared to values for natural repair tissues [16]. However, these constructs often lack the maximum force sufficient to resist the peak in vivo forces acting on the repair site. In fact the strength and stiffness of the cell-gel repairs are still less than 1/3 of normal tendon values [11, 17, 18]. The digital images and the Small Angle Light Scattering (SALS), (a laser-based technique that maps the fibrous structure of planar tissues [19] and measures the spread of the intensity distribution of the projected light pattern (orientation index)) images of these cell-scaffold composites reveal that the orientation of the collagen fibers is not uniform throughout the constructs (Figure 1.1). The color scale method displays the spatial position of data points and the orientation index at each location. Magenta represents higher orders of collagen fiber alignment and blue represents lower orders. 3

20 The darker band observed around the periphery in the digital image and the high fiber orientation observed in the corresponding regions in the SALS image suggest that these cell-scaffold constructs have well-aligned, densely-packed collagen fibers around the periphery and randomly-aligned, loosely-packed collagen fibers in the middle regions. Figure 1.1. Digital and Small Angle Light Scattering (SALS) images of a typical construct showing non-uniform collagen fiber orientation. This gave rise to the third set of research questions; Could marginal fiber alignment in the constructs be the result of the end effects acting on the constructs while they contract around the posts in the silicone dish? Could these changes be minimized by 4

21 increasing the distance between the posts? These questions led to my next two specific aims and two hypotheses. Specific Aim 3 (Chapter 4): Create more aligned constructs by increasing the distance between the posts to reduce the end effects acting on the bulk of the constructs (St. Venant s principle). In culture, compare the biomechanical properties (linear stiffness and linear modulus) of the long constructs with those of the short constructs. Hypothesis 3: Increasing the distance between the posts and thereby lengthening the constructs will significantly improve the fiber orientation and the biomechanical properties of the construct. Specific Aim 4 (Chapter 4): Examine whether fiber alignment in the constructs after 14 days in culture correlates with construct biomechanics. Hypothesis 4: Collagen fiber alignment and construct modulus are linearly related. A previous study in our laboratory evaluated different collagen scaffolds in culture to decide which scaffold was the most suitable to be used in tissue-engineered tendon repair in terms of cell viability, cell penetration and construct integrity to create the stem cell-collagen constructs. Different collagen scaffolds were evaluated: 30 min UV crosslinked collagen fibers (fiber bundles of 25 and 50 fibers), 3 different kinds of collagen films (30 min UV crosslinked, 60 min UV crosslinked, and non-crosslinked 5

22 film), and 4 different types of collagen sponges (gamma-irradiated glucose-crosslinked, gamma-irradiated non-glucose crosslinked, non-gamma irradiated glucose-crosslinked and non-gamma irradiated non-glucose crosslinked sponge). Type I collagen sponge was identified as a better scaffold material in terms of cell viability, cell penetration and construct integrity to create the stem cell-collagen constructs. This material thus became the scaffold of choice for my remaining studies. Studies from our laboratory indicate that scaffold material may influence how cell-based tissue engineered constructs respond to an in vitro mechanical stimulus The potential benefits of mechanical stimulation have been investigated in our laboratory using a bioreactor for MSC-seeded collagen constructs based on a system used previously by Nabeshima et al [20]. Shearn et al [21] and Juncosa-Melvin et al [22] demonstrated that two weeks of mechanical stimulation of cell-collagen gel-collagen sponge constructs and cell-collagen sponge constructs in culture improve their stiffness and also significantly improve repair biomechanics at 12 weeks post surgery. However, Dressler et al [23] saw no significant improvements when cell-collagen gel constructs were mechanically stimulated using the same bioreactor. These findings, conducted using different cell lines, suggest that the response of MSCs to mechanical stimulus is scaffold material specific. This resulted in the fourth research question; how do scaffold materials, type I collagen gel and type I collagen sponge, influences the response of MSCs to a mechanical stimulus? To address this research question the following specific aim was developed. 6

23 Specific Aim 5 (Chapter 5): Compare the in vitro linear stiffness of mechanically stimulated cell-sponge constructs with the stiffness of mechanically stimulated cell-gel constructs. Hypothesis 5: Mechanically stimulated cell-collagen sponge constructs will exhibit significantly higher linear stiffness than stimulated cell-gel constructs. I then decided to examine the potential interactive effects of scaffold materials, construct length and mechanical stimulation. Although numerous studies have examined the individual effects of scaffold material [11, 13, 16, 24, 25], construct length [26, 27] and mechanical stimulation [21, 22, 25, 28] on the in vitro biomechanical properties of a cell-based construct, no study, to my knowledge, has examined the combined and potentially interactive effects of these three key factors and which one(s) is(are) most important to control. Thus the research question was asked; Do scaffold material, construct length and mechanical stimulation synergistically affect the construct s in vitro stiffness? To answer this question, the following specific aim and corresponding hypothesis was posed. Specific Aim 6 (Chapter 6): Examine how altering scaffold material, the construct length, and the mechanical stimulation affect the biomechanical properties of the resulting constructs after 14 days in culture. Hypothesis 6a: Mechanical stimulation of longer cell-sponge constructs will result in the highest in vitro stiffness among the treatment groups studied. Hypothesis 6b: Interactions will be found among the three treatment factors. 7

24 One strategy to further improve the construct stiffness is by optimizing the mechanical stimulus. Most of the mechanical stimulation studies related to tendon tissue engineering have examined a single stimulus [21, 22, 25, 28] and this stimulus contains multiple components (e.g. peak strain, frequency, duration, etc.) whose effects are unknown. Often times the values chosen are limited by the lack of knowledge of in vivo signals or the capabilities of the bioreactor systems which do not mimic physiological conditions. The amplitude, frequency, and duty cycle that are chosen should be based on actual physiological in vivo stress-strain histories rather than on bioreactor capabilities or restrictions. To my knowledge, no studies have optimized the components of the mechanical stimulus for tendon tissue engineering applications. This gave rise to the research question, would optimizing components of the mechanical stimulus further improve the construct stiffness and repair outcome? This research question led to the final two specific aims and associated hypotheses aimed at determining how optimizing the components of mechanical stimulus might affect in vitro linear stiffness of the construct. Specific Aim 7 (Chapter 7): Examine how altering and optimizing peak strain, cycle number per day and cycle repetition of a mechanical stimulus affect the in vitro linear stiffness using Response Surface Methodology (RSM) [29] Hypothesis 7: Optimizing these three components, peak strain, cycles/day and cycle repetition will further improve in vitro stiffness of the constructs. Finally the research question was asked, Do constructs, exposed to the above optimized stimulus in culture, improve the repair outcome compared to constructs 8

25 exposed to a non optimized stimulus? Based on this research question, the following specific aim and hypothesis were developed. Specific Aim 8 (Chapter 8): Surgically implant optimized and and non-optimized constructs in contralateral rabbit patellar tendon defects and compare repair biomechanics at 12 weeks post surgery. Hypothesis 8: to the repairs resulting from constructs exposed to the optimized stimulus improves repair biomechanics at 12 weeks post surgery compared to the repairs containing constructs exposed to the non-optimized stimulus. In summary, the specific aims proposed in this Chapter are primarily aimed at enhancing tendon repair by examining construct composition (cell-to-collagen ratio, length of the construct and scaffold material) and mechanical stimulation. The first six specific aims (SA) examine the individual (SA 1-5) as well as the combined (SA 6) effects of these factors on the in vitro biomechanical properties of the engineered tendon constructs. The last two specific aims deal with in vitro optimization of mechanical stimulus (SA 7) to improve the in vivo repair outcome (SA 8). Findings from these specific aims are presented in Chapters 3 to 8 of this dissertation. 9

26 Chapter 2 Literature Review 2.1. Tendon Structure Tendons transmit muscle force to bone attachments to effect movement of lower limbs during gait, and motion of the thorax, upper extremities, and head during activities of daily living [18]. Tendon is a biphasic material composed of solid and fluid components. Approximately 65% of the total wet weight of the tissue is water. Type I collagen represents approximately 95% of all collagen present and 65 80% of the dry mass in tendon [30, 31]. The remaining extracellular matrix is elastin, proteoglycans, and other minor collagen types (III, V, etc) [31-33]. Although generally considered inert, the solid constituents of tendon turn over with half-lives of days [34]. Collagen is responsible for the tensile properties of tendon. Although more than 20 genetically distinct types of collagen are known [35], the fibrillar collagens (types I, II, III, V, and XI) are critical components of tendon [30]. The collagen molecule is composed of three left-handed α helices coiled into a righthanded triple helix [36, 37]. Each α chain has approximately 1000 amino acids. These amino acids have the G-X-Y peptide sequence, where G, X, and Y are glycine, any amino acid and proline or hydroxyproline, respectively (Figure 2.1). This primary structure allows the tightly wound conformation of the α helix as the glycines orient themselves into the center of the triple stranded super-helix forcing the prolines and hydroxyprolines to the outside, thereby exposing their rigid five-membered pyrimidine ring [38]. 10

27 Figure 2.1. Known structural hierarchy of collagen. The structural hierarchy of collagen is shown with increasing complexity from bottom to top. At the bottom is the typical protein sequence G-X-Y, where G is glycine, X can be any amino acid, and Y is typically proline or hydroxyproline. The chains form the collagen triple helix. The molecules then form microfibrils that are axially staggered. This arrangement causes striations that exhibit a 67 nm peridocity. From Orgel et al [38]. Collagen undergoes several post-translational modifications to achieve the form that is found in the extracellular matrix (ECM). The premature intracellular 11

28 fibrillogenesis is forbidden by the progression of processing events. The procollagen, collagen precursor, is co-translationally transported into the rough endoplasmic reticulum (RER). The procollagen chains are then modified and assembled into triple helices. These helices consist approximately 150 and 250 additional globular propeptides at the amino- (N) and carboxy- (C) termini, respectively. The inter-chain disulfide bonds of the propeptides assist proper winding of the α helices [36, 37]. The propeptides, after being secreted from the cell are cleaved by extracellular enzymes and fibrillogenesis occurs. Crosslinking occurs between fibrils at the nonhelical amino and carboxyl termini. Covalent aldol crosslinks form between 4 lysine or hydroxylysine residues at the N and C termini (2 at each terminus) [38]. The collagen molecules pack together, forming fibrils in a quarter stagger arrangement (Figure 2.1). The result of this organization is a striation or banding pattern with a nm separation between bands that can be seen under electron microscopy [38]. At the macrostructural level (Figure 2.2), tendon collagen consists of parallel and crimped primary and secondary fiber bundles (fascicles), the periodicity of which dramatically decreases close to the tendon insertion into bone [39, 40]. Each bundle contains fibers visible at the light microscope level. Under scanning and transmission electron microscopy, the fibers are composed of fibrils ( nm in diameter), the fundamental building blocks of tendon, which in turn contain microfibrils (diameter 3.6 nm) [36]. At the molecular level (detected by X-ray diffraction), aligned and staggered 12

29 Figure 2.2. The organization of the tendon structure from collagen fibrils to the entire tendon [30]. tropocollagen molecules (length 280 nm; diameter 1.5 nm) are bound through transient and permanent (pyridinium) cross-links that permit force transmission along the entire tendon structure [35, 36] Frequency and Significance of Musculoskeletal Injuries Soft tissue injuries to tendons, ligaments and capsular structures represent almost 45% of the 32 million musculoskeletal injuries that take place each year in US [41]. Regardless of the anatomic site, tendon and ligament sprains represent a major surgical problem. More than 14.5 million patients schedule visits to see physicians in office-based orthopaedic practices due to ligament and tendon sprains and/or joint dislocations [42]. Furthermore, it is reported that the most common tendon disorders occur in the Achilles tendon (100,000 per year), the rotator cuff in the shoulder (51,000 per year), and patellar 13

30 tendon in the knee (42,000 per year) [43-45]. It is also estimated that 75,000 to 125,000 anterior cruciate ligament (ACL) injuries occur each year [46] in the US alone with an estimated 300,000 ACL tears worldwide. More alarming, perhaps, is the rate at which these injuries have increased. Between 1992 and 1999, total knee replacements, for example, increased over 77.8%, at a compounded annual rate of 15.5% [41]. Other joint injuries are also dramatically frequent, with the hip and shoulder being the other two major problematic joints. The cost of the various musculoskeletal injuries amounted to over $30 billion in 1998 [41]. While 86% of these costs are direct hospital and physician costs, about 12% are costs associated with morbidity due to fractures, dislocations, sprains and strains resulting in reduced or lost productivity [41]. In fact, between 1985 and 1988, soft connective tissue injuries accounted for 33% to 47% of all musculoskeletally related hospitalizations, work and school loss days reported [42] Conventional Treatment Techniques These injuries are being traditionally treated by replacing the damaged tissue with graft replacements (autografts, allografts, xenografts) or directly repairing the damaged tissue using sutures [5, 47-49]. Surgeons have been using a wide variety of autografts, such as iliotibial band [50, 51], fascia lata [52], semitendinosus [53, 54], hamstring tendon [55], gracilis tendon [56-58], central and medial (or lateral) portions of the patellar tendon [59-61], and combined grafts such as the combined semitendenosus and iliotibial band tenodesis [62, 63] and combined semitendenosus and gracilis tendon [64, 65] to replace the damaged tissue. 14

31 A number of animal and clinical studies have demonstrated modest short-term improvements but no significant long-term improvements in graft strength over time, post reconstruction [52, 66]. For example, eight weeks post-operatively, results in the goat model indicate that regardless of the fascia lata graft fixation technique, the anteroposterior translation of the operated knees was significantly increased compared to non-operated control knees [52]. The maximum force of the fascia lata autografts increased from 6% of normal ACL control at time zero to only 15% at 8 weeks, while the stiffness increased slightly from 2% of normal ACL control at time zero to 9% at 8 weeks [52]. Other longer-term studies reported a return of graft strength and modulus in 26 weeks of only 13 % and 27% of normal ACL, respectively, when combining the fascia lata with the lateral portion of the canine patellar tendon [66]. A complication associated with harvesting the autograft is the creation of a secondary morbidity site. When tendon autograft tissues are either too narrow or previously damaged, surgeons often use bone-patellar tendon-bone (BPTB) allografts to reconstruct the ACL. While allograft reconstruction avoids the need to create a secondary morbidity site, associated high risk of transmitting infectious diseases such as hepatitis and HIV makes it less favorable for repair. To minimize the risk of disease transmission, surgeons have used freeze-dried ethylene oxide-sterilized BPTB allografts in various animal models [67, 68] and in human patients [69]. But, Roberts et al [69] cautioned that using ethylene oxide-sterilized allografts in human patients can lead to complete graft dissolution and the formation of large femoral cysts in a significant number of patients who received the allografts [69]. In addition, careful assessment of the effectiveness of ethylene-oxide sterilization against HIV is still lacking [69]. Fideler et al 15

32 [70] reported that a dose of 3 Mrad gamma radiation or more is necessary for the sterilization of a fresh-frozen bone-patellar ligament-bone allograft, so that it can be used for reconstructive procedures without the risk of transmission of the virus to the recipient. Unfortunately, high levels of gamma irradiation also significantly alter allograft strength and viscoelasticity [71-73]. When frozen BPTB grafts were exposed to 3 Mrad of gamma irradiation, the maximum force and strain energy to maximum force of the composite unit were significantly reduced by 27% and 40%, respectively [71]. While a dose of 3 Mrad gamma radiation significantly reduced the maximum stress, maximum strain, and strain energy density to maximum stress, a dose of 2 Mrad didn t alter these properties significantly [71]. In fact, when subjected to higher levels of gamma irradiation, that may be necessary to avoid any risk of disease transmission [70], more significant dose-dependent reductions in initial structural and material properties of BPTB allografts were reported [72]. These reductions in strength were accompanied by significant reductions in hydroxypyridinium crosslink density [72]. Reduction of strength due to gamma radiation made allografts fall out of favor. Jackson et al [74] compared similar-sized patellar tendon autografts and fresh-frozen allografts used for goat anterior cruciate ligament reconstruction. After 6 months, the strength and modulus of the autografts were 44% and 59% of normal ACL controls, respectively. By comparison, the allografts were only 34% and 47% of normal ACL strength and modulus, respectively. The allografts also showed a slower rate of remodeling and a prolonged inflammatory response, that was distinctly different from the more robust remodeling response observed in the autografts [74]. 16

33 Although the incidence of direct patellar tendon injury is relatively low as compared to the frequency of other soft tissue injury, disruption due to its use as an ACL replacement is very common. As a gold standard, surgeons most often reconstruct the ACL using the central third portion of the patellar tendon (Figure 2.3) [75]. PT grafts result in a 90% success rate for ACL reconstruction [76]. However, the use of this graft (as well as any other autograft) to reconstruct the ACL and restore Figure 2.3. ACL reconstruction using central third portion of the patellar tendon. kinematics creates a secondary morbidity site in an otherwise healthy tissue. Patellar tendon rupture after central third harvest [77], transverse patellar fractures [78], and medial or lateral patellar subluxation and dislocation [79] in human patients have been reported. In fact, several animal studies have demonstrated that the secondary morbidity (harvest) site may never return to normal, even after prolonged periods of time [80-82]. Proctor et al [82] reported that the maximum force to failure and the ultimate stress of the repair tissue were significantly decreased by 51% and 65%, respectively, compared to normal PT control, 21 months after central PT removal for ACL reconstruction in the goat knee [82]. Histological sections demonstrated that the harvest site was filled with collagenous scar tissue that 17

34 was less organized, composed primarily of small collagen fibrils, more cellular and more vascular than the normal PT control tissue [82]. The results reported by Proctor et al [82] are similar to previous results reported by Burks et al [83]. The donor site created in the canine model by harvesting the central BPTB grafts was filled with scar tissue, and furthermore, that scar formation extended onto all sides of the remaining two-thirds of the patellar tendon [83]. Moreover, these tissues showed significant decreases in the failure load compared to normal PT controls. Initially, the remaining two-thirds of the patellar tendon were 47% as strong as the normal patellar tendon. Three months later, the strength of the healing patellar tendon increased to 71% of normal patellar tendon, but then declined to 60% of normal patellar tendon strength at 6 months. Changes in PT modulus were rather interesting, as the study reported continued decline from almost 90% of normal PT values initially to 32% and 16% at three and six months, respectively, post graft harvest [83]. Beynnon et al reported similar results in the rabbit model [81]. It is obvious from all these studies that even after long periods of time, the repair tissue that develops after harvest of a central bone patellar tendon bone (BPTB) autograft is inferior to normal tissue not only in terms of its structural and material properties, but also in terms of its histological and ultrastructural appearance. It is for that reason surgeons avoid harvesting the same graft in revision of an ACL reconstruction. Clearly then, the repair of the patellar tendon after harvesting a central BPTB graft represents a clinically significant repair model, thus justifying our decision to utilize it in this study. 18

35 2.4. Prosthetic Ligament and Tendon Replacement Use of synthetic ligaments to replace the damaged ACL offers a number of advantages over ACL reconstruction using central PT third autografts. Most importantly, synthetic ligaments do not require creating secondary morbidity sites. Frequently used synthetic materials in tendon and ligament reconstructions are 1) permanent synthetic substitutes [84]; 2) augmentation devices [85]; and 3) tissue growth-inducing implants [86]. The latter essentially serves as a scaffold onto which fibrous tissue forms and incorporates, and gradually matures in the presence of tensile forces to cause its fibers to align longitudinally as in tendon and ligament [4]. One such prosthesis is carbon fibers. Although they induced tissue growth [86-90], the induced tissue was characterized by chronic foreign-body reaction and a disorganized ECM [91]. In addition fragmentation of carbon fiber implants in vivo resulted in accumulation of carbon particles in regional lymph nodes. Such carbon particles have been found to inhibit fibroblast growth [92]. Although investigators used a wide variety of prosthetic ligament and tendon replacements such as Dacron prosthesis (polyester) [84, 93], Gore-Tex prosthesis (polytetrafluoroethylene) [94-96], and braided polyethylene ligament augmentation devices [85], and Leeds-Keio (polyester) and Aramid ligament prostheses [96], these synthetic ligament replacement systems have fallen out of favor in orthopaedic practice because of high incidence of rupture or loosening in patients [97, 98]. In addition, these prostheses could not fully replace the function of the torn ACL and restore knee stability to prevent osteoarthritis [93]. Further, induced tissue growth is often characterized with a prolonged inflammatory reaction [97]. 19

36 It would, therefore, be desirable to have a functional ligament/tendon replacement that is biologic in origin, elicits no unusual inflammatory reaction, and is incorporated into the host [99]. Reconstituted collagen scaffolds have gained increasing interest as ligament/tendon replacements. Researchers have investigated in vivo repair of Achilles tendon gap defects [48, 100] and ACL reconstruction [101, 102] using collagen scaffolds made from thin, type I collagen fibers, dehydrothermally crosslinked by heating under vacuum, followed by treatment with cyanamide vapor, or by exposure to glutaraldehyde vapor [100]. When rabbit Achilles tendons were reconstructed with these collagen scaffolds, the dehydrothermally-crosslinked collagen scaffolds were resorbed by 10 weeks post-implantation, and neotendon incorporation was evident. By contrast, the glutaraldehyde -crosslinked scaffolds were surrounded by a fibrous capsule, suggesting a prolonged inflammatory response without tissue growth induction [100]. Long term evaluation revealed that the dehydrothermally-crosslinked scaffolds had an initial strength of about 35% of normal Achilles tendon values, dropping rapidly to about 11% at 3 weeks before climbing back to close to 80% at 20 weeks. By one year, however, the strength of the dehydrothermally-crosslinked scaffold was down to about 66% of the strength of the normal Achilles tendon [48]. When compared to the repair generated by simply reattaching the devascularized Achilles tendon autograft, the reconstituted collagen scaffolds provided no significant structural advantage [48, 100]. The ability to induce tissue regeneration and repair using biologic materials, however, is often limited by individual variability. One such variable, which can be attributed to poor healing in older individuals is the reduced number of reparative/regenerative cells with increasing age [103]. Other repair impairment 20

37 mechanisms have also been hypothesized to result from diminished cell proliferative capacity [104]. New concepts in tissue engineering may now offer potential cell-assisted repair approaches to overcome such impairments Tissue Engineering as a Possible Alternative to Conventional Tendon Repair Tissue engineering offers an attractive alternative to these clinical problems and has been studied in a variety of animal models. Tissue engineering has been defined as a field of science that makes use of biological and/or biocompatible synthetic materials in conjunction with cells to create new tissues or biologic substitutes to serve as functional tissue replacements [4]. The source of the cell can be a differentiated, specialized cell or a primitive, multi-potential stem cell. Although the basic requirement is that these cells are autologous, to avoid rejection by the immune system, investigators are trying allogeneic and xenogeneic cells as alternatives [24, 105]. Differentiated, specialized cells harvested from a secondary, non-functional tissue site can then be isolated and culture expanded in vitro, so that they can later be delivered to the site of injury in an appropriate delivery construct. As with autografts, the creation of a secondary morbidity site to harvest the autologous cells is an undesirable compromise. Stem cells, to the contrary, can be obtained from bone marrow and other reservoirs in the body using a simple, minimally-invasive procedure that avoids the creation of a secondary morbidity site. Significant recent advances have added much to our knowledge about mesenchymal stem cells and their lineages, and their potential use to augment repair of mesenchymal tissues such as bone, cartilage, ligament and tendon [8, 10, 103, ]. 21

38 The choice of biologic or biosynthetic scaffolds onto which cells are seeded is governed by several basic requirements. The first and the most important requirement is biocompatibility or non-toxicity of the scaffold material [109, 110]. It is utmost important that these materials don t produce any abnormal responses and/or toxic or carcinogenic effects in local tissues surrounding the implant. Biodegradable materials, especially, should serve their function while releasing products of degradation that are biocompatible and non-toxic [109]. The second requirement of the scaffold is that it avoids infection. Thus, the scaffold material should be sterile before seeding the cells onto the scaffold, and should remain sterile prior to surgical implantation of the cell-seeded prosthesis. Third, the sterilization technique should be carefully chosen so as not to significantly alter the mechanical and physical properties of the scaffold material, particularly if the technique weakens the scaffold and causes premature failure in vivo [109]. Finally, the scaffold material should provide some mechanical resistance initially to replace the function of the lost ligament or tendon before its degradation and replacement with new extracellular matrix. Investigators have been using a vide variety of biologic or biosynthetic materials in cell-seeded soft connective tissue analogs. Vacanti et al [4] employed a non-woven mesh or a parallel array of fibrous polyglycolic acid (PGA) scaffolds onto which fibroblasts were seeded. Polylactic acid (PLA), polyglycolic acid (PGA), and their copolymers, from which sutures are typically fabricated, have also been used as scaffold materials for regenerating both cartilage and bone [111]. Others have utilized amorphous collagen gels [5, 10, 13] as scaffolding materials. Indeed, collagen remains the basic protein of interest in the field of soft connective tissue engineering because it is the most 22

39 abundant protein in these tissues [6]. Nishiyama et al [112] reported that collagen gels seeded with fibroblasts begin to contract after an initial gelation time, resulting in increased collagen density in the contracted gel to a thousand times the initial density. This contraction process proceeds without involvement of chemical reactions, synthesis or degradation of any molecules [113]. The mechanisms of gel contraction remain to be elucidated, however. Takakuda et al [114] proposed that fibroblasts generate tension in collagen fibers by attaching to them. The cells then stretch themselves along the tense fibers and increase tension in that direction. The cell mediated contraction of collagen cell can be facilitated by providing anchoring sites. Huang et al demonstrated [7] the importance of providing anchorage sites (posts) for the contraction of a fibroblast-seeded collagen gel (ligament equivalent). Once dispersed in the collagen gel in culture, the fibroblasts exerted traction forces on the collagen matrix and initiated the gel contraction process. By providing fixed posts in the culture dishes, the contraction was constrained, leading to uniaxial organization of the cells and the extracellular matrix. Yamato et al [113] examined the contraction of a fluorescein 5'-isothiocyanate- (FITC) labeled type-i collagen gels seeded with human diploid fetal lung fibroblasts and reported that the cells changed in shape from spherical initially to an elongated shape at a later stage, with the actin microfilaments running along the long axis of the cell. They also reported that the collagen fibrils began to be attached to the cells by 12 hours of culture. Many collagen fibers appeared to be condensed beneath the cells, implying that the process of contraction and reorganization was mediated by physical traction, and involved cell-surface receptors and cytoskeletal reorganization. 23

40 Cytoskeletal contributions to contraction forces generated by fibroblasts in collagen gels have also been studied in vitro [115, 116]. Brown et al [116] studied the mechanical forces and microtubule contribution to fibroblast contraction using a culture force monitor device and reported that disruption of microfilament abolishes contraction forces. Microtubules appeared to act to maintain cell shape and consequently to store a residual internal tension (RIT) [116]. Using selective disruption of cytoskeletal components, it has been estimated that the microtubular RIT constituted up to 33% of the measurable force exerted on the collagen gel [116]. On the other hand Nishiyama et al [115] reported that actin microfilaments are more intimately involved in fibroblastmediated collagen gel contraction than microtubules. Further, it has been documented that fibroblasts cultured within a stabilized collagen gel generate stress that is transmitted throughout the matrix (stressed gel), while cells that are cultured within a freely floating collagen gel in culture media do not generate stress (relaxed gel) [117]. Fibroblasts in stressed collagen gels developed large bundles of actin microfilaments (stress fibers) and were associated with fibronectin fibril assembly, while fibroblasts in the relaxed gels developed no stress fibers. When fibronectin fibril assembly was inhibited, the cells still formed prominent actin bundles and developed isometric tension, indicating that contraction forces are independent of the presence of fibronectin [117]. The cell-mediated contraction process is obviously more complicated in a threedimensional collagen gel construct as compared to two-dimensional collagen gel construct. The cell-to-collagen ratio and the process of collagen purification seem to control the speed of the collagen gel contraction and the appearance of the collagen fiber bundles after contraction [118]. Investigators have also shown that cell nuclei align more 24

41 rapidly with the suture axis and elongate more fully at higher cell-to-collagen ratios [12]. The density of collagen fibers in pepsinized type I collagen gels appeared to be significantly less than in non-pepsinized collagen gels. Thus, the non-helical end domains of the collagen molecule, that are removed when treated with pepsin, appear to be of crucial importance for identification and organization of the collagen fibers by fibroblasts [118]. Juncosa-Melvin et al [16] investigated the effects of decreasing cell-tocollagen ratio by 20 times from 0.8 to 0.04 cells/mg collagen on the cell-mediated repair process following implantation in vivo. They reported that constructs containing the highest cell-to-collagen ratios were damaged due to excessive contraction and no significant differences were observed in any structural or material properties or in histological appearance between the two lowest cell-to collagen ratios (0.04 and 0.08) when implanted in vivo [16]. In this study, I will investigate the effects of decreasing cellto-collagen ratio by 20 times from 0.8 to 0.04 cells/mg collagen on the contraction kinetics of cell-seeded collagen constructs Mesenchymal Stem Cells Stem cells have gained increasing interest due to their potential applications in medicine. Stem cells possess the ability to self-renew through repeated mitotic divisions, and the ability to generate differentiated specialized progeny [9, 119]. Although several different stem cell systems including hematopoietic stem cells (HSCs) [120], neural stem cells (NSCs) [121, 122] and mesenchymal stem cells (MSCs) [8, 107] have been identified, studies show that these systems have several common characteristics such as the ability to undergo asymmetric cell-divisions in which a stem cell may divide to 25

42 produce one daughter stem cell (identical to the mother cell) and one differentiated (committed progenitor) daughter cell [119]. Stem cells may also exist in mitotic quiescence, and have the ability to regenerate all of the different cell phenotypes that constitute the tissue system to which they belong [119]. Mesenchymal stem cells (MSCs) are adult progenitor cells with the ability to repair soft and hard connective tissues. These multipotential cells have the capacity to form cartilage, bone, tendon, ligament, marrow stroma and other connective tissues [8]. MSCs continue to function throughout life as a continuous supply source of committed cells that are involved in homeostatic remodeling of the skeletal system and repair of injured tissue [106, 107]. By taking biological and mechanical cues from the local environment MSCs differentiate along multiple cell lineages to produce fibroblast (tendon and ligament), chondrocyte (cartilage), osteocyte (bone), and other types of mesenchymal tissue cells [107]. The progression along any of these lineages towards the end phenotype is also dependent on the genomic potential of the cells (autocrine regulation), presence or absence of vascularity and seeding density of MSCs [8, 9, 107]. The populations of MSCs in marrow, however, have been shown to decline significantly with increasing age in humans [8]. Caplan et al [8] have reported that MSCs in human bone marrow diminish in number almost exponentially from birth (1 in 10,000 nucleated cells) to the eighth decade of life (1 in 2,000,000 nucleated cells). It is believed that MSCs from older individuals have the potential to progress along bone and cartilage lineages, albeit this remains to be verified in vivo. Caplan et al [8] have also demonstrated that human MSCs can be cultured through many passages without loss of their osteo- and chondrogenesis potential [8]. It is believed that MSCs, just as they are involved in 26

43 skeletal tissue generation during embryonic phases of life, remain involved in tissue repair and regeneration throughout life [8, 107, 123]. Therefore, the idea of using MSCtransplantation for in vivo tissue repair and regeneration has been investigated in different tissue and animal models using these techniques [9, 10, 13, 15, 105, 106]. Although many investigators have taken advantage of the simple harvest and multipotentiality of the MSCs in bone and cartilage repair [2, 8, 106, ] very few have investigated their use in tendon [11, 13, 15, 24] Functional Tissue Engineering Functional tissue engineering emphasizes the importance of biomechanical considerations in the design and development of cell and matrix-based implants for soft and hard tissue repair [127]. Functional Tissue Engineering ideally seeks to regenerate and substitute damaged tissue, and replicate the mechanical performance of normal tissue. Applying the concept of Functional Tissue Engineering [17, 127] new generations of reparative tissue constructs to regenerate and functionally replace damaged tissue can be developed [18, 128]. Multipotent stem cells, such as mesenchymal stem cells (MSCs) give rise to the progenitors of various structural and connective tissues, including bone, cartilage, fat, tendon, and muscle [129, 130]. Researchers seek to accelerate tendon repair using rabbit mesenchymal stem cells (MSCs) to fabricate the implants [13, 15, 24, 105]. Investigators employed different techniques including simply placing cells and collagen gel in the wound site [11] and by aligning the cells along sutures before implantation [12, 13] to improve tendon repair. Tissue engineering using mesenchymal stem cells is attractive 27

44 [11, 13, 15] but tissue stiffness and strength do not yet meet in vivo loading demands [17, 131, 132]. Recent tissue engineering studies, particularly those using mesenchymal stem cells (MSCs), have examined the effect of different construct constituents on the rate and quality of tendon healing. MSCs are isolated from bone marrow biopsy and then multiplied in culture before being mixed with a carrier or scaffold, typically a type I collagen gel and implanted in a window defect in the rabbit patellar tendon (PT). Modest improvements in repair biomechanics were seen compared to natural repair of unfilled defects [11]. When MSCs are first permitted to contract the gel around a suture, the resulting construct shows twice the repair biomechanical properties at 4, 8 and 12 weeks after surgery compared to inserting only suture in the wound site [12, 13]. Both repair studies produced better results than for natural healing. Awad et al. [15] showed that autologous MSCs, when seeded at three different cell densities (1, 4 and 8 x10 6 cells/ml) in a type I collagen gel, significantly improved repair outcome over natural healing, but that these improvements were not dose-dependent and induced ectopic bone formation in 28% of the repairs. In the rabbit AT, a high MSC density (4 x10 6 cells/ml) construct results in twice the structural and material properties compared to natural healing between 4 and 12 weeks post surgery, suggesting the potential for early and longer-term improvements but not necessarily normal tissue properties [13]. In this study, the modulus and maximum stress for the repair tissue were 34% and 37%, respectively, of normal values at 12 weeks after surgery. Using a PLGA scaffold instead of collagen gel with suture, MSCs were found to produce denser matrix than scaffold alone and an apparent crimp pattern was evident at 4 weeks [24]. At 12 weeks the stiffness of the 28

45 repair tissue treated by PLGA with MSCs reached 87% compared to 56% in repairs using PLGA alone [105] Mechanical Stimulation to Control Phenotype In Vitro One of the central concepts of functional tissue engineering [17] is developing improved therapies using bioreactors that apply physical stimulation to a prospective tissue engineered construct in vitro. Motivated in part by Wolff s Law and studies on the debilitating effect of stress shielding in tendon [113, ], cellular studies have examined the effect of strain and stress on cell function. When subjected to strains created by a deformable substrate in vitro, cells modify their cytoskeletal orientation [ ], proliferation [ ], as well as protein transcription and synthesis [141, ]. Mechanical stimulation studies exclusively using tendon cells have shown similar results, exhibiting increased DNA synthesis [142]and expressing novel genes [ ]. While the studies listed above have clearly shown that cells respond to mechanical stimulus, they also posses three major limitations. First, most studies have been performed in monolayer possibly ignoring important cell-matrix interactions that exist in a 3-D environment. Cells can act markedly different in ECM than they do when cultured on a 2-D dish or flask [149]. This is due, in part, to the types of ECM-cell attachments that are formed, the substrate stiffness at these attachment sites, and the number of attachments. Many 2-D cell culture systems can adequately model the integrin attachments by applying a specific coating to the culture surface, but they cannot reproduce the deformable ECM and, since they are flat, can only involve one half of the 29

46 cellular membrane or less. Consequently 2-D systems neglect the dynamic or adaptive reciprocity that exists between cells and their ECM. A closer approximation of the in vivo system is achieved by utilizing a 3-D matrix, such as a collagen gel. This is extremely valuable when trying to model wound healing and tissue morphogenesis as the cells that participate in the healing cascade are involved with inflammation, histogenesis, and remodeling, all of which are mediated to some extent by cell-matrix interactions. Second, Most of the mechanical stimulation studies related to tendon tissue engineering have examined a single stimulus [21, 22, 25, 28] and this stimulus contains multiple components (e.g. peak strain, frequency, duration, etc.) whose effects are unknown. Often times the values chosen are limited by the capabilities of the bioreactor systems do not mimic physiological conditions. As determined from in vivo measurements [131, 132, 150, 151], the forces in tendon during the gait cycle are complicated and can barely be approximated using the typically cyclic functions (sine, saw-tooth, haversine, etc). The amplitude, frequency, and duty cycle that are chosen should be based on actual physiological in vivo stress-strain histories rather than on bioreactor capabilities or restrictions. Different bioreactors have been developed to impart a mechanical stimulation on cell-seeded constructs for tendon and ligament repair [ ]. The method of stimulation is quite common among the documented systems. In general, the construct is either fabricated in a custom jig [152, 153] and then placed into the bioreactor or simply fabricated directly in the bioreactor [154]. In both cases the construct is then gripped and stretched cyclically. All maintain a physiological environment (37 C, 95% relative humidity, 5% CO2) and allow for media exchange for long term culture. The various 30

47 drive systems for the bioreactors include: linear motor [153], stepper motor with a pulley system [152], and stepper motor with gear system [154]. The tissue engineered tissues can be maintained in culture for several weeks and are then removed for evaluation. Some of the bioreactors have load cells that can monitor the development of force in the tissue in real time[152, 153]. In an attempt to tissue engineer a ligament through directed differentiation via mechanical stimulation, Altman et al developed a unique bioreactor for MSC-seeded collagen gels [141, 154]. Similar in concept to those described above, however, due to the helical orientation of the collagen fibers and the complex loading experienced by ligaments, the bioreactor delivers both a translational and rotational stimulus (10% strain and 90 rotation at Hz (one cycle per minute)). After 21 days of mechanical stimulation in this bioreactor MSCs expressed mrna of collagen types I and III and tenascin-c, which are indicators of a fibroblast phenotype. At the same time the constructs did not exhibit any upregulation of osteo- or chondro-specific markers. Histologically the cells appeared elongated and in parallel with the collagen fibers that were oriented in a helical fashion similar to normal ACL. While several methods for the mechanical stimulation have been reported [124, ], very few have used these methods to mechanically stimulate tendons. To my knowledge, no studies have optimized the stimulus for tendon applications. 31

48 Chapter 3 Effect of Cell Seeding Density and Collagen Concentration on Contraction Kinetics of MSC-Seeded Collagen Constructs 2 Our group has been engineering cell-scaffold constructs to improve tendon repair by contracting mesenchymal stem cells (MSCs) in collagen gels and then evaluating their repair potential in wound sites in rabbits. Since the construct s initial conditions may influence the ultimate repair outcome, this two-part study sought to distinguish which factors most influence contraction kinetics in culture. 1) We optically determined if varying cell-to-collagen ratio significantly affected construct contraction. Temporal changes in construct area were monitored up to 168 hours for four cell-to-collagen ratios (HK= 0.04, LK= 0.08, HM= 0.4, and LM= 0.8 where H, L= 2.6, 1.3mg/ml collagen and K, M= 0.1, 1 million cells/ml). A mathematical model was created with terms that represent the different combinations of cell densities and collagen concentrations in order to predict the contraction kinetics as a function of time. Highly significant differences in construct areas were found among all four ratios after 8 hours of contraction with the exception of the LK (0.08) vs. HM (0.4) conditions. This similar pattern raised the question as to whether cell density or collagen concentration more influenced these events. 2) To isolate these effects, the contraction kinetics of the HM construct were compared to those of a new construct (L5K) with equivalent cell-to-collagen ratio (0.4) but half the cell density (500K MSCs/ml) and half the collagen concentration (1.3mg/ml). 2 This manuscript has been published in Tissue Engineering, Vol. 12(7), pp ,

49 The L5K construct contracted significantly faster and more completely than the HM construct but no differently than the LM construct. These results indicate that above a threshold value of cell density, percentage reductions in collagen concentration influence contraction kinetics more than equivalent percentage increases in cell seeding density. The fact that our model successfully predicted intermediate time points of contraction suggests its utility when examining other cell and collagen densities. Controlling scaffold as well as cellular initial conditions will be critical in achieving our goal of functional tissue engineering (FTE) a successful tendon repair Introduction Repair of tendons, ligaments and capsular structures is common given that these injuries represent almost 45% of the 32 million musculoskeletal cases in the US each year [41] Unfortunately, primary repair and replacement using autografts, allografts, and xenografts is not routinely successful [99, ] due to donor site tissue damage, inadequate graft availability and potential long-term tissue rejection. As a consequence, surgeons and basic scientists have sought to identify new approaches like tissue engineering that might accelerate tissue repair and return the patient to pre-injury activities. Recently, tissue engineers have been creating cell-scaffold constructs in culture using multi-potential mesenchymal stem cells (MSCs) and various biomaterials to repair and reconstruct defects in articular cartilage [10, 158, 159], Achilles [13, 24] and patellar tendons [11, 12, 15, 160], bone [106, 161] and muscle [10, 162]. Although various synthetic and biologically-derived delivery vehicles have been tried, our group has 33

50 seeded MSCs in type I collagen gels to repair defects in the rabbit Achilles and patellar tendons [11, 12, 15, 160]. While introducing large numbers of cells in collagen gel-suture constructs accelerates repair of tendon defects [11, 13, 15], ectopic bone forms in almost a third of all patellar tendon repair sites [15]. To avoid bone formation that might be caused by cell density-dependent contraction of the 3-D construct [12] and the possible stress-shielding effects of the stiff suture [14], we redesigned the constructs to anchor to posts, permitting cells to directly sense cell traction forces. With this new system, preliminary experiments appeared to show that cell-to-collagen ratio did affect contraction of MSC-gel constructs [163], but the dose dependent effect of these ratios and the role of gel concentration were not examined. Thus the current study sought to understand how cell-to-collagen ratio affects contraction kinetics of MSC-collagen constructs as they mature around posts in a silicone dish in culture and which component of this ratio (cell density or collagen concentration) is more important in this process. Temporal changes in normalized area of the construct (ratio of current-to-initial area (A/A 0 )) served as a single measure for kinetics of contraction with initial area being the area of the well excluding the post areas (Figure 3.1). In a two-part experiment, we hypothesized that: 1) increasing cell- Figure 3.1. Silicone dish consisting of 4 wells with posts protruding from the base of each well. Cell-seeded collagen gel constructs contract around the two posts in each well showing the initial well area, A 0, and the area of the construct at a given time, A. 34

51 to-collagen ratio significantly increases contraction kinetics, and that 2) above a certain threshold value of cell seeding density, percentage reductions in collagen concentration influence contraction kinetics more than equivalent percentage increases in cell seeding density. 3.2 Methods Experimental Design Four adult female NZW rabbits were utilized as part of two related experiments with repeated samples (3 replicates for each animal) of MSCs for studying five treatment levels at the same cell passage number. Experiment 1. Four cell-to-collagen ratios (HK= 0.04, LK= 0.08, HM= 0.4, and LM= 0.8 where H, L= 2.6, 1.3mg/ml collagen and K, M= 0.1, 1 million cells/ml) were contrasted with n=12 for each level by seeding MSCs at two concentrations (0.1 and 1 million cells/ml, respectively) in two different collagen concentrations (2.6 and 1.3 mg/ml, respectively). Experiment 2. Two identical cell-tocollagen ratios (0.4 M/mg) were contrasted, one using MSCs at 0.5 million cells/ml in a 1.3 mg/ml concentration of collagen (L5K) and the other using the same HM conditions as in Experiment 1. Digital images of the constructs provided temporal changes in construct areas at 4, 8, 16, 32, 72, 120 and 168 hours in culture. Harvest of MSCs Adapting published methods [11, 12], bone marrow samples were collected from the iliac crest of each rabbit as follows. Each rabbit was anesthetized by an intramuscular injection of a mixture of ketamine hydrochloride (100 mg/ml and 0.25 ml/kg body 35

52 weight), and acepromazine (10 mg/ml and 0.07 ml/kg body weight). The hair was shaved over the lumbar-sacral area, and the shaved area was then sterilely prepped using betadine solution and alcohol. A 12cc syringe was loaded and coated with 1 ml of diluted sodium heparin (3,000 units/ml). The diluted heparin was prepared by sterilely mixing 10 ml of heparin (5,000 units/ml) with 5 ml of Dulbecco s Phosphate Buffered Saline (D-PBS). With the rabbit in sternal recumbency, a volume of 3 to 4cc of bone marrow was withdrawn directly into the heparin from the iliac crest using an Illinois marrow biopsy needle. The syringe was capped and the heparin and bone marrow were mixed thoroughly. Each sample was mixed with growth medium, Dulbecco s modified Eagle s medium (DMEM)-low glucose (Invitrogen, Life Technologies, Inc., Carlsbad, CA) supplemented with 10% fetal bovine serum from selected lots [164] (Atlanta Biologicals, Lawrenceville, GA) and 1% antibiotic/antimycotic (Atlanta Biologicals). The sample was washed with DMEM and centrifuged at 2000 rpm for 6 minutes. The MSCs were separated from the media, plated in 100 mm culture dishes at 0.5 million cells per dish, grown in the incubator for 2 weeks and fed twice a week with growth media prepared as previously described. At the end of first passage, colonies of MSCs that adhered to the plate were detached using 0.25% Trypsin-EDTA (Atlanta Biologicals), counted, their viability determined using the Trypan blue exclusion method [165] and subcultured again into passage two. At the end of the second passage, MSCs were taken from the dishes, centrifuged and re-suspended in growth media for construct preparation. 36

53 Preparation of Constructs All cell suspension aliquots were mixed with neutralized type I purified bovine collagen gel (Cohesion Technologies, Palo Alto, CA). A 1.4 ml aliquot of cell-gel mixture was pipetted into each of four wells within specially-designed silicone dishes to allow contraction around posts (Figure 3.1). In order to control the other factors that might affect the contraction kinetics, such as cell initial conditions, all constructs within a cell line were created simultaneously to allow simultaneous data collection. Contracting constructs were fed high glucose DMEM supplemented with 10% fetal bovine serum and ascorbic acid and incubated at 37 o C, 5%CO 2 for 7 days. Images were analyzed using National Instrument s Vision 7.0 software. Normalized construct areas were plotted against logarithmic time of contraction to create more equally-spaced time intervals for analysis. Statistical Analysis The treatment structure consisted of two factors with fixed effects, termed Treatment and Time, and another factor with random effects, called ANIM (animal) corresponding to the planned replications of the experiment. To test hypothesis 1, four of five treatment combinations (LK, HK, LM, HM) were initially studied in pre-planned experiments. To test hypothesis 2, a supplemental experiment was performed to contrast the effects of L5K vs. HM. Each treatment factor was observed at seven levels of time (4, 8, 16, 32, 72, 120, and 168 hours) with three replications which were then averaged to produce one set of temporal results per animal. Since treatment structure was not entirely factorial (there were 3 levels of cell seeding densities and only 2 levels of collagen 37

54 concentration), it was decided to model the treatment structure as having one factor with five levels corresponding to the combinations of the cell seeding density and collagen concentration, namely, HK, HM, L5K LK, and LM. The levels of the treatment factor were to be observed at each level of the time factor with reduced number of observations per animal when excessive cellular contraction damaged one or more constructs. Despite the unbalanced number of observations per group, one model still produced estimates of all relevant effects and allowed for all relevant hypotheses to be tested (Appendix). ANOVA was performed to analyze multiple contrasts related to both cell-tocollagen ratio (hypothesis 1) and cell density vs. collagen concentration (hypothesis 2). In particular, the effects of increasing cell density at high and low collagen concentrations were individually compared with the effects of decreasing collagen concentration at 100K and 1 million cells/ml, respectively. The effect of increasing cell density from 100K to 1 million cells/ml was calculated at both high and low collagen concentration levels (HK HM; and LK LM). The effects of increasing cell density from 0.1 to 0.5 million cells/ml and from 0.5 to 1 million cells/ml were also calculated at low collagen concentration (LK L5K; and L5K LM). Similarly, the effects of decreasing the collagen concentration from high to low values were calculated for both 0.1 and 1 million cells/ml (HK LK; and HM LM). Adjustments were also made in reported p-values to account for the one-tailed nature of Experiment 2 for multiple comparisons among groups. 38

55 Mathematical Prediction of Contraction Kinetics Based on an examination of plots of the data for all five treatments and on historical precedent in this area of study [12, 118] a nonlinear mixed model was developed that would satisfy the homoscedasticity and normality assumptions by curve fitting permit normalized area and time to be related Results Although some similarities were noted in construct appearance, contraction kinetics varied dramatically among the groups. All constructs became more opaque with increasing time of contraction (Figure 3.2) and, by 16 hours, most developed darker bands around their periphery compared to lighter areas in the central region between the posts (Figure 3.2). At 168 hours of contraction, the construct with the highest cell-tocollagen ratio (LM: 0.8) contracted the most (to 5 % of its initial area) while the construct with the lowest cell-to-collagen ratio (HK: 0.04) contracted the least (to 31 % of initial area). Between 32 and 168 hours, half of the HM, L5K and LM constructs contracted so appreciably that they tore off the posts (Table 3.1). 39

56 Figure 3.2. Digital images of typical HK, LK, HM and LM constructs at 4, 8, 16, 32, 72, 120 and 168 hours. Constructs became darker with increasing time of incubation. Darker bands were frequently visible around the periphery of the construct by 16 hours with lighter regions between the posts. 40

57 Cell/collagen Ratio HK (0.04) LK (0.08) HM (0.4) L5K (0.4) LM (0.8) Time (hrs) ± ± ± ± ± ± ± ± ± ± ± ± ± ± ± ± ± ± ± ± 0.88* ± ± ± 1.44* ± 1.20^ ± 0.85^ ± ± ± 2.02^ ± 0.94^ 6.60 ± 0.78^ ± ± ± 2.03^ ± 1.49^ 5.19 ± 0.61^ * n = 9 replicates from 4 animals ^ n = 6 replicates from 4 animals Table 3.1. Average percentages of initial area (mean ± SEM) of HK, LK, HM, L5K and LM constructs at 0, 4, 8, 16, 32, 72, 120 and 168 hours. Note the decrease in percentage with increasing time and with increasing cell-to-collagen ratio. Sample size, shown for each treatment level and time point, is N = 4 (with 3 replicates for each cell line) unless otherwise noted. The low values of SEM indicate that the measurements do not vary drastically within groups. Hypothesis 1 was accepted (Figure 3.3). Highly significant differences were found in normalized construct areas between HK (0.04) and LK (0.08) (p = ) and between HM (0.4) and LM (0.8) (p = ). The differences between HK and HM (p = ) and between LK and LM (p = ) were also highly significant (Table 3.2). Surprisingly, however, no significant difference was found in the normalized area contraction curves for LK (0.08) and HM (0.4) (p = ). This finding suggested that cell-to collagen ratio might not be the driving factor in the contraction process and hence the individual effects of cell density vs. collagen concentration should be examined. 41

58 Normalized Area Vs Log(time) A/Ao hrs HK (0.04) LK (0.08) HM (0.4) LM (0.8) Log(time) Figure 3.3. Increasing cell-to-collagen ratio increases the rate and magnitude of contraction except for two cases (LK and HM) suggesting that factors other than just cellto-collagen ratio also affect contraction behavior. Normalized Area vs. Log (time) for HK, LK, HM and LM with N = 4 with 3 replicates from each rabbit. Treatment 1 Treatment 2 p-value HK LK L5K HM LM < < LK L5K HM LM L5K HM LM HM LM Table 3.2. Contrast analysis of treatment means. Every treatment is significantly different from any other treatment except LK & HM; and L5K & LM. 42

59 Hypothesis 2 was also accepted. Despite the same cell-to-collagen ratio (0.4), the construct with the lower collagen concentration (L5K: 1.3mg/ml) displayed significantly lower percentages of initial construct area than the construct with the higher cell density (HM: 1M cells/ml) at equivalent time intervals of contraction (p= 0.012; Table 3.1). Surprisingly, significant differences were also found in percentage areas between the LK (0.08) and L5K (0.4) constructs (p = 0.031), but not between the L5K and LM (0.8) constructs (p = ) (Figure 3.4 & Table 3.2). Normalized Area Vs Log(time) hrs A/Ao HM (0.4) LM (0.8) L5K (0.4) Log(time) Figure 3.4. Decreasing collagen concentration (from HM to L5K) causes greater contraction than increasing cell density (from L5K to LM). 43

60 The potential interactive effects of cell density vs. collagen concentration were also examined. Highly significant differences were observed between halving the collagen concentration at either 1 million cells/ml (HM LM) or 0.1 million cells/ml (HK LK) and doubling the cell density at low collagen concentration above 0.5 million cells/ml (L5K LM) (p= and p= , respectively). However, no significant differences were found between halving collagen concentration [either (HM LM) or (HK LK)] and increasing the cell density by five fold ( million cells/ml) at the low collagen concentration (LK L5K) (p= and p= , respectively). Similarly, no significant difference (p = ) was observed between increasing the cell density at low collagen concentration by five-fold (LK L5K) vs. ten-fold (LK LM). These observations validate hypothesis 2 that decreasing collagen concentration affects normalized area more than increasing cell seeding density above a threshold level of 0.5 million cells/ml. The following nonlinear mixed model was developed. A = ( b + Db1 + C1b2 + C2b3 )exp{( r0 + Dr1 + C1r2 + C2r3 )logt} + + ε. ij 0 U i ij Here A ij is normalized area (Appendix), t represents time and C 1, C 2 and D are physical constants related to treatment levels for cell density and collagen concentration (Table 3.3). Estimates of the intercept (b o, b 1, b 2 and b 3 ) and rate (r 0, r 1, r 2, and r 3 ) parameters are given in Table 3.3. U i andε are error terms due to subject and measurement variations, ij respectively. This nonlinear least squares (PROC NLIN) method for estimating exponential fit matched the experimental results very well across the five treatments (R>0.98). The model, which ignored the random effect of the dishes, was particularly effective for interpolation between selected time points. 44

61 Indicators Treatment Parameter D, C1, C2 Density Cell Count Intercept Rate 1, 1, 0 HIGH M b 0 +b 1 +b 2 r 0 +r 1 +r 2 1, 0, 0 HIGH K b 0 +b 1 r 0 +r 1 0, 1, 0 LOW M b 0 +b 2 r 0 +r 2 0, 0, 1 LOW 5K b 0 +b 3 r 0 +r 3 0, 0, 0 LOW K b 0 r 0 A = ( b + Db1 + C1b2 + C2b3 )exp{( r0 + Dr1 + C1r2 + C2r3 )logt} + + ε. ij 0 U i ij Table 3.3. Model parameters, intercepts and rates for different treatment combinations. D = 1 if collagen concentration is at the high level, and 0 otherwise. Also, C 1 = 1 if cell density is one million cells/ml, and 0 otherwise; while C 2 = 1 if cell density is 5K cells/ml, and 0 otherwise. Curve fitting produced intercept and rate values as follows: b 0 = , b 1 = , b 2 = , b 3 = , r 0 = , r 1 = , r 2 = , r 3 = , VAR(U) = and VAR(ε) = Discussion Cell-to-collagen ratio is an important factor that controls contraction kinetics. The fact that the highest cell-to-collagen ratio (LM) caused the constructs to contract more significantly than an intermediate ratio (HM), which in turn contracted more significantly than the lowest ratio (HK), indicates that cell-to-collagen ratio helps to control cellmediated contraction (Figure 3.3). These results are consistent with previous reports [12, 166]. Such cell-based contractions likely compact the collagen fibrils as the constructs draw away from the edges of the dish [166]. As cell-to-collagen ratio increases, this cellmediated contraction likely increases this fibril compaction. At the same time, cell-to-collagen ratio is not the only factor affecting contraction kinetics. The lack of significant differences between the two intermediate cell-tocollagen ratios (LK: 0.08 and HM: 0.4) suggests that cell seeding density and collagen 45

62 concentration be examined separately. Bell et al [166] reported that the rate of contraction depends on both cell seeding density and collagen concentration. However, this study failed to report which of these parameters was more important and thus led to our second hypothesis. While cells contract and align adjacent collagen fibrils (resulting in microstructural rearrangement of the collagenous extracellular matrix (ECM) [113, 167]), collagen fibrils also provide resistance to these cell-driven traction forces. Although contraction is generally viewed as a cell-mediated event where larger numbers of cells can produce more traction forces to facilitate contraction, the significant difference in kinetics between the LK and L5K constructs, with no corresponding differences between the L5K and LM conditions, suggest that increasing cell seeding density above a threshold value (0.5 million cells/ml in this study) does not further enhance the effect. Instead, collagen concentration appears to influence these temporal patterns more than cell seeding density (Figure 3.4) suggesting that availability of protein binding sites may be the more critical factor in cell-mediated contraction. These results are also consistent with observations from two previous reports [12, 166]. Awad et. al detected no significant differences in contraction around a suture when MSC seeding density was increased from 4 to 8 million cells/ml but did find differences between 1 and 4 million cells/ml [12]. Taken together, these results indicate the need to more fully investigate cell-matrix interactions over a broader range of cell-to-collagen ratios and to examine potential mechanisms for these contraction effects by monitoring corresponding changes in formation of integrins and cytoskeletal architecture. Reductions in collagen concentration affect contraction kinetics more than proportional increases in cell density, especially above a seeding density of 0.5 million 46

63 cells/ml. Halving the collagen concentration at either 1 million cells/ml (HM LM) or 0.1 million cells/ml (HK LK) produces a greater reduction in percent area than doubling the cell density at low collagen concentration above 0.5 million cells/ml (L5K LM) [high collagen concentrations were not compared over this cell density range]. In fact, to elicit the same effect as halving collagen concentration [either (HM LM) or (HK LK)], cell density would have to be increased by five fold ( million cells/ml) at the low collagen concentration (LK L5K). The importance of the 0.5 million cell/ml threshold is reinforced by noting the lack of difference in contraction kinetics between a five-fold increase in cell density (which just meets the threshold) and the ten-fold increase (which spans across the threshold) [(LK L5K) and (LK LM)]. Thus, it would seem more efficient and less expensive to regulate collagen concentration levels vs. cell densities in order to better control contraction kinetics. Our comprehensive, nonlinear model, with its physical variables, has helped us to better understand the cell-gel contraction process. Unlike previous models of contraction kinetics [12, 118], the current model allows simultaneous interpolation of contraction kinetics for specific treatments as well as statistical comparisons of rates of contraction using contrast analyses. The model also demonstrates the complex nature of cellmediated contraction, as evidence by the nonlinear relationship between the square root of normalized area and the log (time). Also, the physical variables (C s, D in the model) represent the actual experimental conditions that have been used in this study. Therefore, with the help of this model one can quantify the contraction kinetics at various time points for any particular treatment. 47

64 To better control the contraction process and avoid premature damage to the construct (e.g. LM), more uniform and homogenous structures will need to be created. The anchoring posts at the ends of each well create boundary conditions that likely influence the appearance of the constructs, producing dark bands around the posts and periphery of the construct and light bands in the interior. The inhomogeneous appearance of these dark and light bands is probably associated with local differences in collagen fibril compaction and water expulsion and may also be associated with differences in local stresses and strains. Understanding these local stress-strain fields during contraction (e.g. using finite element models and Small Angle Light Scattering to assess local fiber directions [19]) should be helpful in designing new dishes where high stresses can be controlled and premature construct damage (LM) can be avoided. Controlling MSC differentiation in vitro will be critical to creating desirable phenotypes that express appropriate ECM. Such in vitro approaches should ultimately improve repair outcome after surgery, a goal of functional tissue engineering [17, 127, 168]. Besides cell-to-collagen ratio, cell seeding density and collagen concentration, other factors such as mechanical stimulation and growth factors are also expected to influence contraction kinetics and construct properties. Our combined in vitro and in vivo studies are examining such effects (e.g. mechanical stimulation) on in vivo repair outcome after surgery Appendix The design structure used in the model was a Split Plot design. Treatment was selected as the whole plot factor with its five levels arranged in a one factor, 48

65 completely randomized, design with the available Animals (Anim) serving as specimens. Time was chosen as the split plot factor with 7 levels. The Anim*Treatment interaction served as the basis for the whole plot error used to test the significance of levels of the Treatment factor, while the model residuals were used to test the significance for both Time alone and Time *Treatment interaction. SAS software was used to perform the analyses using PROC MIXED. Prior to finalizing the nonlinear relationship between the response variable, normalized area, and time, diagnostic tests were performed to determine whether the usual statistical assumptions necessary for proper inference were satisfied, namely homoscedasticity and normality of the residuals. The residual plots indicated that the homoscedasticity assumption was suspect and at least one of the normality tests rejected the normality assumption. Various transformations on the contraction scores were then considered, resulting in a model for which A ij was the normalized area of the j th specimen within the i th dish for i = 1,,6, and j varying from 1 up through the number of observations taken within a dish. This model passed the diagnostic tests. The computations were carried out using PROC NLMIXED, a nonlinear mixed model fitting program. The estimation method employed for the exponential fit was maximum likelihood. Note that, the R-squared value was greater than 0.98 when the same exponential model was fit using PROC NLIN, which ignores the random effect of the dishes and uses nonlinear least squares as the method for estimating the exponential fit. 49

66 Chapter 4 Effect of Length of the Engineered Tendon Construct on its Structure-Function Relationships in Culture 3 Constructs containing autogenous mesenchymal stem cells (MSC) seeded in collagen gels have been used by our group to repair rabbit central patellar tendon defect injuries. Although these cell-gel composites exhibit improved repair biomechanics compared to natural healing, they can be difficult to handle at surgery and lack the necessary stiffness to resist peak in vivo forces early thereafter. MSCs are typically suspended in collagen gels around two posts in the base of a well in a specially-designed silicone dish. The distance between posts is approximately the length of the tendon wound site. MSCs contract the gel around the posts prior to removal of the construct for implantation at surgery. We hypothesized that in vitro construct alignment and stiffness might be enhanced in the midregion of the longer construct where the end effects of the posts on the bulk material (St. Venant effects) could be minimized. Rabbit MSC s were seeded in purified bovine collagen gel at 0.04M cells/mg collagen. The cell-gel mixture was pipetted into silicone dishes having two post-to-post lengths (short: 11 mm and long: 51 mm) but equivalent well widths and depths and post diameters. After 14 days of incubation, tensile stiffness and modulus of the constructs were measured using equivalent grip-to-grip lengths. Collagen fiber orientation index or OI (which measures angular dispersion of fibers) was quantified using Small Angle Light Scattering (SALS). 3 This manuscript has been accepted for publication and will appear in Journal of Biomechanics,

67 Long constructs showed significantly lower angular dispersion vs. short constructs (OI of o ± 1.57 o vs o ± 1.27 o,.mean ± SEM, p < 0.001) with significantly higher linear modulus (0.064 ± MPa vs ± MPa, p=0.0022) and linear stiffness (0.031 ± MPa vs ± N/mm, mean ± SEM respectively, p=0.0404). We now plan to use principles of functional tissue engineering to determine if repairs containing central regions of longer MSC-collagen constructs improve defect repair biomechanics after implantation at surgery Introduction The frequency of soft connective tissue injuries and the variable success associated with conventional repair methods have led investigators to develop innovative tendon model systems to study treatment strategies like tissue engineering. Soft tissue injuries to tendons, ligaments and capsular structures represent almost 45% of the 32 million musculoskeletal injuries that take place each year in the US alone [41, 42, 44, 45, 127, 169]. Tears to the rotator cuff and Achilles tendon are particularly common as are surgically-induced defect injuries to create central bone-patellar tendon-bone (BPTB) grafts for cruciate ligament reconstruction. While conventional approaches are successful in repairing some injuries, the repair tissue that develops after BPTB graft harvest shows inferior histological and ultrastructural appearance as well as structural and material biomechanical properties [60, 82, 83, 170]. Such challenges make alternative treatment strategies like tissue engineering more appealing for potential primary and revision surgery applications. Among the most exciting tissue engineering approaches involve the 51

68 possible use of multipotential mesenchymal stem cells (MSCs) seeded in collagen scaffolds. Our group has taken an evolutionary approach using these MSC-collagen constructs to accelerate repair of rabbit patellar tendon defects in a load-controlled setting as part of a functional tissue engineering paradigm [11, 15, 16, 22]. MSCs are isolated following bone marrow biopsy, multiplied in culture, mixed with a type I collagen gel or scaffold, contracted around posts in specially-designed silicone dishes, and introduced into surgically induced central patellar tendon defect sites to try and accelerate repair. Although this approach improves the repair biomechanics of rabbit patellar and Achilles tendon defect wounds by 43-50% compared to natural healing alone [11, 13, 15], the strength and stiffness of these cell-gel based repairs do not match those of normal tendon even after 26 weeks, still achieving only 20-28% of normal tendon unoperated values [15, 17, 18]. Among the many factors that could influence construct stiffness and repair outcome, we questioned whether the initial geometry (i.e. length-to-width or aspect ratio) of the construct might have an important role. Specifically, we wondered if initial geometry might influence initial collagen fiber alignment and resulting biomechanics. In pilot studies, we performed small-angle light scattering (SALS) analysis [19] of these constructs in the lab of one of our co-authors (MSS). This analysis revealed non-uniform collagen fiber alignment in the constructs, with well-oriented fibers around the posts but poorly-aligned fibers between them (Figure 4.1). Digital images of these cell-scaffold composites confirmed the presence of highly-oriented, darker bands around the periphery and randomly aligned lighter regions between the posts. 52

69 These findings raised the question as to whether the spatial variations in fiber alignment between the posts might be due to the end effects of the posts acting on these relatively short constructs. Although this question of how to increase fiber alignment has not specifically been answered in the literature, one study did investigate how aspect ratio influenced spatial variations in cell morphology. Figure 4.1. The darker band observed in the peripheral regions of the digital image and the high fiber orientation observed in the corresponding regions of the SALS image suggest that these cellscaffold constructs have well-aligned, denselypacked collagen fibers around the periphery and more randomly-aligned, loosely-packed collagen fibers in the middle regions. Eastwood et. al [26] reported that mechanical forces generated by cellular activity in fibroblast-populated collagen lattices (FPCL) depended on both the applied strain and the structure s aspect ratio. After providing 16 hours of cyclic loading to low and high aspect ratio structures, cells in the low-ratio FPCL configuration showed no clear orientation whereas cells in the high-ratio FPCLs showed parallel alignment with the axis of the applied load [26]. While these results demonstrated how changes in construct geometry could influence cell morphology 53

70 [26], the study did not address how these changes in aspect ratio might ultimately affect extracellular matrix (ECM) fiber orientation and subsequent construct biomechanics. To address this research question, we sought to uncouple how collagen alignment might be affected by construct dimensions vs. environmental stimuli like mechanical stimulation. Our study specifically examined whether increasing the distance between the posts might improve fiber alignment during maturation in culture by reducing end effects [i.e. by St. Venant s principle [171]]. We hypothesized that increasing post-topost distance and thus construct length would significantly improve collagen fiber alignment which in turn would increase construct stiffness and modulus in vitro. Fiber alignment was quantified using SALS and linear stiffness and linear modulus were respectively measured from force-displacement and stress-strain curves obtained from axial failure tests on constructs with equivalent grip-to-grip lengths Methods Experimental Design Six skeletally mature, female, New Zealand white (NZW) rabbits (Myrtles Rabbitry, Thompson Station, TN) were used in this study. Bone marrow biopsies were performed and mesenchymal stem cells (MSCs) were isolated and expanded as previously described [13]. Cells were seeded at 0.1 x 10 6 cells/ml with purified bovine collagen gel (Cohesion Technologies, Palo Alto, CA) at 0.04M cells/mg collagen. This cell seeding density and collagen concentration were chosen based on the results from two previous studies in our laboratory [16, 172]. The cell-gel mixture was pipetted into specially-designed silicone dishes having two different post-to-post lengths (11 and 51 54

71 mm) but identical dish widths (11mm), depths (6mm) and post diameters (4mm) (Figure 4.2). Twenty-four constructs (12 long and 12 short) were created from each of the six different cell lines and allowed to contract around the posts with half assigned to study collagen fiber alignment and the other half to study construct biomechanical properties. Figure 4.2. Short and long silicone dishes with two different post-to-post lengths (11 and 51 mm) but identical dish widths (11mm), depths (6mm) and post diameters (4mm). Harvest of MSCs Bone marrow samples were collected from the iliac crest by adapting previously published methods [11, 12]. Each sample was mixed with growth medium, Dulbecco s modified Eagle s medium (DMEM)-low glucose (Invitrogen, Life Technologies, Inc., Carlsbad, CA) supplemented with 10% fetal bovine serum from selected lots [164] (Atlanta Biologicals, Lawrenceville, GA ) and 1% antibiotic/antimycotic (Atlanta Biologicals). The sample was washed with DMEM and centrifuged at 2000 rpm for 6 minutes. The MSCs were separated from the media, plated in 100 mm culture dishes at 55

72 0.5 million cells per dish, grown in the incubator for 2 weeks and fed twice a week with growth media prepared as previously described. At the end of first passage, colonies of MSCs that adhered to the plate were detached using 0.25% Trypsin-EDTA (Atlanta Biologicals), counted, their viability determined using the Trypan blue exclusion method [165] and sub-cultured again into passage two. At the end of the second passage, MSCs were taken from the dishes, centrifuged and re-suspended in growth media for construct preparation. Preparation of Constructs All cell suspension aliquots were mixed with neutralized type I purified bovine collagen gel (Cohesion Technologies, Palo Alto, CA). Since the volume of the long well was 2.5 times that of the short well, the volume of the cell-gel mixture pipetted into the long wells was adjusted accordingly (3.5 vs. 1.4 ml, respectively). Contracting constructs were fed high glucose DMEM supplemented with 10% fetal bovine serum and 5% ascorbic acid and incubated at 37 o C, 5% CO 2 for 14 days. After 14 days of incubation the constructs were frozen at -80 C for 2 weeks. Collagen Orientation Collagen orientation was quantified using small angle light scattering (SALS), a method commonly used to show the organization of collagenous tissues [19, ]. Briefly, unpolarized HeNe laser light (wavelength = nm) is transmitted through tissue and is diffracted by the collagen fibers, similar to single-slit light diffraction theory [19]. The diffracted light is projected onto a screen and imaged. To facilitate scattering, 56

73 samples are first dehydrated in a glycerol solution through three aqueous solution concentrations (50%, 75%, 100%). The angular distribution of scattered light was analyzed using previous published methods [19]. Specifically, we used an orientation index (OI) that quantified the degree of fiber alignment in a distribution-independent manner. Note that large OI values represent points with random fiber orientation and small values represent highly-aligned fiber orientation. In our study, the laser scanned the entire construct in 250 µm x 250 µm evenly-spaced sampling points, from which the OI was computed. Average OI was calculated for the entire region between the innermost edges of the posts for the short constructs and an equivalent length of the middle region for long constructs. Biomechanical Evaluation On the day of mechanical testing, specimens were thawed at room temperature after being frozen for two weeks. The width and thickness were measured optically using a digital camera (Coolpix 4500, Nikon, Melville, NY) from which average crosssectional area was calculated. Pieces of gauze with cyanoacrylate were sandwiched onto both ends of the constructs to provide a bearing surface that would minimize slippage or premature failure of the specimen in the grips. Specimens were carefully placed into a custom materials testing system (100R6, TESTRESOURCES, Shakopee, Minnesota) while simultaneously leaving as much tissue as possible (about 5 6 mm on both sides) to minimize gripping effects on specimen properties [171]. The specimens were then failed under displacement control at a rate of 10%/sec in a PBS bath at room temperature [15]. Axial linear stiffness and linear modulus were calculated from the linear regions of 57

74 the force-elongation and stress-strain curves generated by the specimen during failure testing. Nominal stress was determined by dividing force by the initial cross-sectional area and the linear modulus was calculated from the slope of the stress-strain curve in the linear region. Statistical analysis Mean OI, linear stiffness, and linear modulus were each compared between long and short constructs within animals using paired Student t-tests [29]. This paired t-test was justified given that the residuals were found to satisfy normality at a 5% level of significance. Short and long constructs were also compared using multivariate Hotelling s T 2 test with respect to all three of the measurements simultaneously. All conclusions were made at the α = 0.05 experiment-wise level Results Biomechanical Evaluation The final dimensions and biomechanical properties of the short and long constructs are given in Table 4.1. During biomechanical testing, failure modes were inconsistent among both the short and long constructs, occurring either within the midsubstance in four of the constructs or near the grips in another two constructs. Thus, even though highly significant differences were found in failure biomechanical parameters like maximum force and maximum stress between the long and short constructs (Table 4.1), the low aspect ratio of less than 3 to 1 precluded us from analyzing these failure properties across treatment conditions. Thus, the remaining discussion is limited to 58

75 comparisons of two subfailure parameters from tensile testing, modulus in the linear region of the stress-strain curve and stiffness in the linear region of the force-elongation curve. Short Long Dimensions Thickness (mm) 2.0 ± ± 0.0 Width (mm) 4.8 ± ± 0.24 Length (mm)* 21.4 ± ± 0.32 Structural & Mechanical Properties Maximum Force (N) 0.07 ± ± 0.01 Linear Stiffness (N/mm) ± ± Maximum Stress (MPa) ± ± Linear Modulus (MPa) ± ± * gauge length = 10 for both short and long constructs Table 4.1. The final dimensions and the biomechanical properties of the short and long constructs (mean ± SEM) at day 14. Longer constructs demonstrate significantly improved subfailure (linear modulus and linear stiffness) properties. Longer constructs also show higher failure properties (maximum force and maximum stress) compared to shorter constructs although aspect ratio limitations preclude statistical comparisons. Statistical analyses revealed significant differences in subfailure properties due to construct length. Long constructs generated significantly greater linear moduli compared to short constructs (0.064 ± MPa and ± MPa, mean ± SEM respectively, p=0.0022) (Figure 4.3). Long constructs also demonstrated significantly greater linear stiffness compared to short constructs (0.031 ± MPa and ± N/mm, mean ± SEM respectively, p=0.0404) (Figure 4.3). 59

76 Figure 4.3. Long constructs (L; n = 6) demonstrate significantly higher linear stiffness (mean ± SEM; N/mm) and linear modulus (mean ± SEM; MPa) compared to short constructs (S; n = 6) (p = and , respectively). Collagen Orientation The average thicknesses of the short and long constructs that were scanned using SALS were 2.0 ± 0.05 mm and 2.0 ± 0.0 mm, respectively. Long constructs showed more aligned collagen fibers compared to short constructs. SALS images revealed that long constructs displayed a more uniform and frequent magenta color, reflective of greater collagen fiber alignment, over the full width and length of the structure between the posts (Figure 4.4). By contrast, short constructs showed a lower degree of collagen fiber alignment (colors away from red and closer to blue) in peripheral regions of the constructs and less aligned fibers in the regions between the posts (Figure 4.4). 60

77 Figure 4.4. SALS images of long constructs display a more uniform and frequent magenta color that reflects greater collagen fiber alignment over the full width and length of the structure between the posts. Short constructs show a lower degree of collagen fiber alignment (colors away from red and closer to blue) in peripheral regions of the constructs and less aligned fibers in the regions between the posts. 61

78 Statistical differences were noted in the mean OI between the middle portions of the long constructs and between the posts of the short constructs (41.24 o ± 1.57 o vs o ± 1.27 o, mean ± SEM, respectively; p < 0.001). The Hotelling s T 2 test also showed significant differences in mean OI between short and long constructs (p = ). Correlations between Structure and Subfailure Material Properties Linear modulus (LM; in MPa) and mean orientation index (OI; in degrees) of the tissue engineered construct were linearly correlated (R 2 = 0.5) (Figure 4.5). The equation relating these two parameters was as follows: LM (MPa) = *OI ( ) (4.1) Figure 4.5. Linear correlation between orientation index and construct modulus (R 2 = 0.5) suggest that alignment of collagen fibers influences the material properties of the constructs. 62

79 4.4. Discussion A construct s local strain-stress state is mediated by the mechanical anisotropy of the gel, which evolves during contraction. Orientation of collagen fibers induces contact guidance that produces spatial variations in cell density that can influence contraction kinetics [176] and hence mechanical properties of these constructs. It is these properties that serve as initial conditions for cells at the time of implantation. Therefore it is critical to create more homogenous constructs with more aligned fiber orientations to resist tensile forces. Our experimental results show that increasing the distance between the posts can reduce the heterogeneity in the fiber alignment observed in the short constructs. The change in construct dimensions essentially lengthens the region of the construct that experiences more homogenous fiber alignment in the cross section. MSCs are known to contract collagen gels and produce internal strains in the pericellular matrix. Our experimental results are further supported by the FEA predictions reported by Eastwood et. al [26] in that constructs showed no high strain gradients when loaded in low aspectratio configuration. By contrast the strain gradients in constructs that were loaded (identical applied mechanical load) in high aspect ratio were highly condensed. These highly condensed strain gradients in the high aspect-ratio constructs probably contribute to the improved alignment observed in the long constructs. It is interesting to note that collagen fibers in regions just in front of the posts are more poorly aligned for the short vs. long constructs (Figure 4.4). We believe that these differences are due to the fact that fiber alignment in these regions is not only affected by the presence of the adjacent post, but by the end effects produced by the opposite post as 63

80 well. Since collagen fibers in the short constructs experience adverse effects from both ends, the fibers are thus more poorly aligned in these shorter structures. By contrast, the greater distance between the posts for the long constructs means that fibers in any region are only influenced by end effects from one post. This geometry results in improved collagen fiber alignment in long constructs as compared to short constructs. Improved alignment of fibers and minimization of end effects due to contraction around posts likely also contribute to the improved material properties noted for the long constructs. The fact that only marginal improvements were observed in the stiffness of the long vs. short constructs may be caused by the larger cross-sectional area and shorter length for the short vs. long constructs. The long constructs contracted to a greater degree than the short constructs, resulting in a smaller cross-sectional area that would adversely affect stiffness given equivalent material properties. Similarly, shorter structures display greater structural stiffness given equivalent material quality. The significantly greater modulus for the long vs. short constructs attests to these geometric effects on a tissue s structural properties like stiffness. Several factors may have affected our ability to detect length-related differences in construct biomechanics and collagen alignment. Animal-to-animal differences in cell contractility could have reduced significance levels with respect to construct stiffness. Gripping effects may have also influenced the measured material properties and observed failure mechanisms during mechanical testing. We utilized a relatively small aspect ratio for the short constructs in order to contain the holes at both ends of the constructs within the grips. Uniaxial deformation likely induced biaxial strains near the grips that could have influenced loading in the construct midsubstance which in turn might have affected 64

81 the modulus values. It is not surprising, therefore, that 30% of the failures initiated near the grips due to these elevated stress values. We explored methods to eliminate these end effects, such as cutting dog-bone shaped test specimens, but we were concerned that these biological constructs might be damaged in the region between the grips during this specimen preparation. The fact that gauge length was standardized for both short and long constructs suggests that these actual gripping effects would be similar, apart from any St. Venant effects due to aspect ratio. It could be argued that multiple aspect ratios should be examined in order to properly correlate a construct s structure and function. However, this larger study would require many more cells from each cell line, which unfortunately are not available without expansion to later passages. This study would also require fabrication of numerous expensive stainless steel molds to make silicone dishes with different post-topost lengths. We chose not to pursue this approach before establishing whether longer constructs would also demonstrate superior repair biomechanics in the tendon defect sites. Now that these length-related differences have been found in construct stiffness and modulus, an in vivo study will follow. Should the repairs also improve using longer constructs, we will then establish the critical construct length that optimizes construct properties in culture and repair biomechanics thereafter. The final objective of this study was to determine whether construct structure might predict in vitro biomechanical parameters after 14 days in culture. A linear correlation was established between the construct s mean orientation index and its linear modulus (R 2 = 0.5) at this time interval, suggesting potential structure-function relationships for a tissue engineered (TE) material in culture. Developing such 65

82 correlations could serve to speed the process for identifying stiffer constructs without actually performing time-consuming and complicated failure tests. These relationships may also be stronger for material parameters (e.g. linear modulus) that are less affected by tissue dimensions. Other structure-function correlations for multiple construct lengths should and are being sought at this time. In conclusion, this study has demonstrated that long constructs have improved linear stiffness, modulus and fiber alignment compared to short constructs. Juncosa- Melvin et. al [22] have recently demonstrated that in vitro structural and mechanical properties are correlated with corresponding in vivo repair parameters. These correlations suggest that long constructs will improve the repair biomechanics to better resist peak in vivo forces acting on the repair site. However, this study still needs to be conducted. Given these encouraging results, segments of longer constructs that are equivalent in length to short constructs will now be examined as replacement segments at surgery to fill surgically induced central patellar tendon defects. 66

83 Chapter 5 Mechanical Stimulation of Tissue Engineered Tendon Constructs; Effect of Scaffold Materials 4 Our group has shown that numerous factors can influence how tissue engineered tendon constructs respond to in vitro mechanical stimulation. Although one study showed that stimulating mesenchymal stem cell (MSC) collagen sponge constructs significantly increased construct linear stiffness and repair biomechanics, a second study showed no such effect when a collagen gel replaced the sponge. While these results suggest that scaffold material impacts the response of MSCs to mechanical stimulation, a well-designed intra-animal study was needed to directly compare the effects of type I collagen gel vs. type I collagen sponge in regulating MSC response to a mechanical stimulus. Eight constructs from each cell line (n=8 cell lines) were created in specially designed silicone dishes. Four constructs were created by seeding MSCs on a type I bovine collagen sponge and the other four were formed by seeding MSCs in a purified bovine collagen gel. In each dish, 2 cell-sponge and 2 cell-gel constructs from each line were then mechanically stimulated once every five minutes to a peak strain of 2.4%, for 8 hours/day for 2 weeks. The other dish remained in an incubator without stimulation for 2 weeks. After 14 days, all constructs were failed to determine mechanical properties. Mechanical stimulation significantly improved the linear stiffness (0.048 ± vs ± 0.004; mean ± SEM N/mm) and linear modulus (0.016 ± vs ± 0.001; 4 This manuscript has been accepted for publication and will appear in Journal of Biomechanical Engineering,

84 mean ± SEM MPa) of cell-sponge constructs. However, the same stimulus produced no such improvement in cell-gel construct properties. These results confirm that collagen sponge rather than collagen gel facilitates how cells respond to a mechanical stimulus and may be the scaffold of choice in mechanical stimulation studies to produce functional tissue engineered structures Introduction Tissue engineering offers an attractive alternative to direct repair of injuries to tendons, ligaments and capsular structures that represent almost 45% of the 32 million musculoskeletal injuries that occur each year in the United States [41]. Our group is currently using mesenchymal stem cell (MSC)-seeded collagen constructs to repair tendon injuries in the rabbit model [11, 13, 15, 22] Since type I collagen represents approximately 95% of all collagen present and 65 80% of the dry mass in natural tendon [30, 31], these engineered constructs are created by seeding MSCs in type I collagen. Recent studies in our laboratory reveal that mesenchymal stem cell-collagen gel constructs improve the selected biomechanical properties of the repair tissue (e.g. linear stiffness and modulus) of rabbit patellar and Achilles tendon defect wounds by 43-50% compared to natural healing alone [11, 13, 15] However, these cell-collagen-gel based repairs also lack the strength and stiffness of normal tendon after 26 weeks, achieving only 20-28% of normal unoperated values [15, 17, 18]. One strategy to further improve repair outcomes is to mechanically stimulate the tissue engineered construct to direct differentiation and promote extracellular matrix (ECM) development [141, 177, 178]. Numerous investigators have shown 68

85 that application of mechanical signals to cells in two- and three-dimensional scaffolds alter cell alignment [137, 141, 179, 180] and proliferation [ ], and increase secretion of growth factors (TGF-β, bfgf, and PDGF) [143] and collagen [141, 144]. We have demonstrated that stimulating mesenchymal stem cell (MSC) collagen sponge constructs produced significant increases in construct linear stiffness after two weeks [21, 22] in culture. Delivering these stimulated autogenous constructs into rabbit patellar tendon defects also improved repair biomechanics 12 weeks after implantation [21, 22]. However, the effect of delivering mechanical stimulus to MSC collagen gel constructs has not yet been examined in vivo. Therefore, this study was designed to directly compare how the same mechanical stimulation profile would affect the in vitro properties (linear stiffness and modulus) of the MSC-collagen gel vs. MSC-collagen sponge construct using MSCs from the same cell lines Methods Experimental Design MSCs were isolated from the iliac crest of eight one-year-old female New Zealand White rabbits and then culture expanded until passage two. Eight constructs (two replicates for each of the four treatment groups) from each cell line were created in specially designed silicone dishes containing four wells [21, 22, 172]. Four constructs were created by seeding MSCs (0.14 x 10 6 cells/construct) on a type I collagen sponge (Kensey Nash Corporation, Exton, PA) and the other four were formed by seeding MSCs (0.14 x 10 6 cells/construct) on a purified type I bovine collagen gel (2.6 mg/ml; Cohesion Technologies, Palo Alto, CA). For each dish, two cell-sponge (CS) and two cell-gel (CG) 69

86 constructs were then mechanically stimulated. The remaining two cell-sponge (CS) and two cell-gel (CG) constructs from each cell line remained in an incubator without stimulation. After two weeks of incubation, all constructs were failed in tension under displacement control to obtain their structural and material properties. Harvest of MSCs MSCs were isolated from rabbit bone marrow using previously described methods [11, 13, 15] Marrow aspirates were taken from the iliac crest under general anesthesia. Cells were plated at 22 x 10 6 cells per 100 mm dish, placed in an incubator and fed with growth medium formulated from 89% Dulbecco s modified Eagle mediahigh glucose (DMEM-hg, Gibco, Carlsbad, CA), 1% antibiotic/antimycotic (Gibco), and 10% fetal bovine serum (FBS) of selected lots (Atlanta Biologics, Norcross, GA) for days. MSCs proliferated to form colonies between 6 and 8 days in primary culture. Cells were retrieved after reaching confluency, counted and subcultured again to passage two to get enough cells to make 8 constructs from each cell line. Preparation of Cell-Gel Constructs At the end of passage two, MSCs were retrieved and a cell suspension aliquot (0.137 ml) containing 0.16x10 6 MSCs was mixed with ml of neutralized type I collagen gel (Vitrogen, Cohesion, Palo Alto, CA). The final cell-seeding density consisted of 0.1x10 6 MSCs/ml in a solution containing mg collagen/ml, resulting in a cell-to-collagen ratio of 0.038x10 6 cells/mg collagen. A controlled volume of the cell- 70

87 gel mixture (1.4 ml) containing the 0.14x10 6 MSCs was then pipetted into one well of the silicone dish (Figure 5.1). Figure 5.1. Custom silicone dishes were created to deliver mechanical deformation to MSC-collagen constructs. Each silicon dish contains 4 wells with 2 posts. Constructs are created in these dishes and then placed into the mechanical stimulation system. Preparation of Cell-Sponge Constructs Type I collagen sponges (Kensey Nash Corporation, Exton, PA) (90% pore volume; 62 µm mean pore diameter) were cut to fit in the bottom of the well in the silicone dish. Two 4 mm-diameter holes were created, permitting each sponge to be positioned in the base of each well [22]. Sponges were soaked in phosphate buffered saline (PBS, Gibco BRL/Life Technologies Inc., Gaithersburg, MD) for 24 hours. Constructs were created by injecting a cell suspension aliquot (0.4 ml) containing 0.14x10 6 MSCs onto the top surface of the collagen sponge. All mechanically stimulated and non-mechanically stimulated constructs were placed in an incubator (37ºC, 5% CO 2, 95% RH) for 2 weeks and fed three times weekly with high glucose DMEM, 5% ascorbic acid, 1% antibiotic/antimycotic and 10% FBS. 71

88 Mechanical Stimulation After allowing the cell-collagen gel constructs to solidify and the cells to settle in the matrix for 2 days in the incubator, the silicone dishes containing the constructs assigned for stimulation were placed into a pneumatic mechanical stimulation system that had been augmented since it was first described by Nabeshima et al [20]. This computer-controlled system consists of five stations (Figure 5.2) mounted within an incubator (Steri-Cult Model 3033, Figure 5.2. Mechanical stimulation system. The system has five stations in an incubator (37 C, 95% RH, 5% CO2). Strain is induced in the cell-gel constructs by stretching the deformable silicone dish. Forma Scientific, Marietta, OH). The computer controls the displacement profile of the dishes based on user-defined displacement (amplitude) and time (frequency and rest period between cycles) parameters and also collects displacement data from each station using linear variable differential transducers (LVDTs). Each station holds one silicone dish and contains a pneumatic cylinder to stretch the dish and a LVDT to monitor the end-to-end displacement of the dish. The relationship between post-to-post displacement in the well and end-to-end displacement of the dish has been previously established [181]. Dishes were stretched for 8 hours/day to a peak displacement corresponding to a post-to-post strain amplitude of 2.4%. A strain amplitude of 2.4% was chosen since it falls within the peak in vivo strains for rabbit patellar tendon (2% and 4%) that we 72

89 estimated from in vivo force recordings in this tendon for inclined hopping activities, the highest activity of daily living that we studied [132]. After 14 days in the incubator, both mechanically stimulated and non-mechanically stimulated constructs were placed in cryovials and stored in a -80ºC freezer until mechanical testing. Biomechanical Evaluation Constructs were removed from the freezer and thawed to room temperature on the day of testing. The width and thickness were measured optically using a digital camera (Coolpix 4500, Nikon, Melville, NY) from which average cross-sectional area was calculated. Small squares of gauze were sandwiched onto both ends of the constructs to provide a surface that would minimize slippage as well as premature failure of the specimen in the grips. Specimens were carefully placed into a custom materials testing system (100R6, TESTRESOURCES, Shakopee, Minnesota) while simultaneously leaving as much tissue between the grips as possible to minimize Saint Venant s gripping effects on specimen properties [171]. Specimens were then secured in the grips by applying just enough pressure to hold the specimens in the grips without damaging the tissue and failed in tension at a constant strain rate of 10%/s while continuously recording force and grip-to-grip elongation. Both structural (e.g. stiffness and maximum force) and material properties (e.g. modulus and maximum stress) were calculated from the resulting force-elongation and stress-strain curves. 73

90 Statistical Analysis Statistical analysis was performed using two-way analysis of variance with scaffold material and stimulation as fixed factors. Post hoc testing was conducted using Least Significant Difference (LSD) tests [29]. Residuals were tested for normality and homoscedasticity and found to satisfy these criteria at a 5% level of significance. All conclusions regarding the significance of different treatments were made at p < Results Mechanical stimulation had no effect on either construct s final dimensions but showed a positive effect on the biomechanical properties of the cell-sponge construct (Table 5.1). The dimensions of neither construct was influenced by the mechanical stimulus after 14 days in culture (Table 5.1, p > 0.05). Mechanical stimulation significantly improved the mean linear modulus of the cell-sponge constructs (0.016 ± MPa vs ± MPa; mean ± SEM, respectively) (p = ) (Table 5.1) but had no effect on the cell-gel constructs (0.03 ± MPa vs ± MPa; mean ± SEM; stimulated vs. non-stimulated, respectively; p = ) (Figure 5.3). In a similar way, mechanical stimulation significantly improved the linear stiffness of the cellsponge constructs (0.048 ± N/mm vs ± N/mm; mean ± SEM; p = ) (Table 5.1) but had no effect on the linear stiffness of the cell-gel constructs (0.021 ± N/mm vs ± N/mm; mean ± SEM, respectively; p = ) (Figure 5.4). Constructs exhibited inconsistent failure modes. Two-thirds of the constructs failed near the grips while the remaining constructs failed closer to the mid-substance. 74

91 Mechanical stimulation and scaffold type appeared to have no effect on failure mode. The low aspect ratio (length-to-width ratios of 2:1) for the constructs may have influenced failure mode, however. Sponge Stimulated Gel Stimulated Nonstimulated Nonstimulated Dimensions Thickness 2.1 ± ± ± ± 0.1 Width 5.0 ± ± ± ± 0.2 *Length 21.2 ± ± ± ± 0.24 Structural & Mechanical Properties Max Force (N) 0.11 ± ± ± ± 0.01 Linear Stiffnes(N/mm) ± ± ± ± ** ** Max Stress (MPa) ± ± ± ± Linear Modulus (MPa) ± ** ± ** 0.03 ± ± * The initial gage length was 10 mm for all the constructs. ** Statistically significant difference. Table 5.1. The final dimensions and the biomechanical properties of the cell-gel and cellsponge constructs (mean ± SEM) at day 14. Stimulated cell-sponge constructs demonstrated significantly higher linear modulus and linear stiffness than non-stimulated cell-sponge constructs (p = and ; linear modulus and linear stiffness, respectively). However no difference in any property was observed between stimulated cell-gel constructs and non-stimulated cell-gel constructs (p = and ; linear modulus and linear stiffness, respectively). 75

92 p = Linear Modulus (MPa) S p* = NS S 0 Gel NS Sponge Figure 5.3. Stimulated cell-sponge constructs demonstrated significantly higher linear modulus (mean ± SEM) than non-stimulated cell-sponge constructs (p = ; n = 8). However no difference in modulus was observed between stimulated cell-gel constructs and non-stimulated cell-gel constructs (p = ; n = 8) p* = Linear Stiffness (N/mm) p = S 0.01 NS S NS 0 Gel Sponge Figure 5.4. Stimulated cell-sponge constructs demonstrated significantly higher linear stiffness (mean ± SEM) than non-stimulated cell-sponge constructs, stimulated cell-gel constructs and non-stimulated cell-gel constructs (p = , and , respectively; n = 8 for each group). 76

93 5.4. Discussion We sought to determine whether a construct s scaffold material (type I collagen gel vs. type I collagen sponge) would influence how its MSCs responded to a mechanical stimulus in culture. Our results indicate that a controlled mechanical stimulus stiffened only the cell-sponge construct. The three-fold increase that we observed in in vitro linear stiffness for these sponge constructs is consistent with the results from our previous studies [21, 22]. The fact that our intermittent mechanical stimulus did not improve the biomechanical properties of the MSC-seeded cell-gel constructs is contrary to other reports. Other investigators have noted positive effects of stimulation on collagen gelbased constructs using different cell types (tendon fibroblasts [178], smooth muscle cells [177], fibroblasts [182], and MSCs [141]) and methods of stimulation [141, 177, 178, 182]. Possibly these factors (cell types and methods of stimulation) are what contributed to the observed increases in cell proliferation [182], construct stiffness [177, 182], collagen synthesis [178, 182], elastin production [177] and MMP activity [182]. These earlier findings combined with the results from the current study suggest that constructspecific strain profiles may be needed to optimize desired outcome measures. Our strain stimulation protocol may not have allowed the cell-gel construct sufficient time for recovery. It is well known, for example, that viscoelastic cell-seeded collagen constructs undergo cyclic creep and stress relaxation [177, 183, 184]. If such changes did occur in our study, then constructs would have experienced progressively lower forces in subsequent cycles, eventually remaining unloaded during portions of the strain protocol. Under this scenario, the cells might not have derived the benefits of the 77

94 applied strain stimulus. To address this problem would require that we simultaneously measure construct force and deformation to detect time-dependent changes in construct load during relaxation testing. Using an in-series load cell for each construct would permit detection of loss in construct force and would allow real-time readjustment of initial length at first loading (zero strain state) so that constructs would continue to receive a consistent strain pattern during mechanical stimulation. In that regard, we are nearing completion of a new multi-station, electromechanical stimulation system with individual force transducers that permit tracking of real-time construct stiffness. The fact that MSCs seeded in collagen gel did not stiffen due to the mechanical stimulus could also be due to the architectural and volumetric differences of the collagen between the cell-gel and cell-sponge constructs. Although the same number of cells was used at the same passage number to create both constructs, the amount of collagen in these constructs was not the same. Cell-sponge constructs had significantly more collagen than the cell-gel constructs. This difference could have also influenced the response of MSCs to the same mechanical stimulus. An extracellular matrix (ECM) with higher collagen content (sponge) could have facilitated the response of MSCs to the mechanical stimulus. Our future studies will examine the effect of mechanical stimulation of either cell-gel or cell-sponge constructs with different collagen concentrations to test this hypothesis. Mechanical stimulation is known to upregulate collagen types I and III gene expression [141]. Our results would suggest that MSCs in the cell sponge constructs might upregulate these genes more than for the cells in the cell-gel constructs. We are now beginning a study to test this hypothesis by evaluating these construct for expression 78

95 of collagen type I, collagen type III, decorin, fibronectin and glyceraldehydes-3- phosphate dehydrogenase (GAPDH) genes using real time rtpcr. Despite the small aspect ratio of our constructs that could have influenced our subfailure parameters, we noted important differences due to scaffold material. The small aspect (length: width) ratio of the specimens ( 2:1) was chosen so that the ends of the constructs containing the holes from the posts could be included in the grips. It is not surprising, therefore, that two-thirds of these constructs failed near the grips due to these low aspect ratios. These grip failures likely altered failure mechanisms and thus failure biomechanical properties between the grips due to St. Venant s effects [171]. We explored methods to eliminate these end effects, such as cutting a dog-bone shaped test specimen, but we decided to tolerate these complications rather than risk damage to our specimens during preparation. We are also currently modifying our silicone molds for future experiments to dramatically increase the aspect ratio of the constructs to obtain more faithful material properties in the mid region. In light of this, conclusions from the tensile tests were constrained to stiffness and modulus in the linear region of the forceelongation curve. These parameters were chosen since they are less affected by failure mode than failure properties. The sub-failure properties reported in this study are approximately 3 orders of magnitude smaller than those for the repair tissue. However, we have recently demonstrated, using the same rabbit central patellar tendon defect model, that the in vitro structural and mechanical properties of these compliant cell-sponge constructs are in fact positively correlated with corresponding repair parameters, linear stiffness and modulus [22]. While these correlations are certainly in large part due to the protected nature of the 79

96 central PT wound site, the model is still valid for testing the effects of multiple mechanical and chemical stimuli. The fact that we have shown these stimulation-induced effects both in vitro and in vivo makes this model system relevant for these studies. This study has other limitations. We did not specifically examine the cell viability in the constructs, the spatial heterogeneities in cellularity or matrix content in the two scaffolds, and the potential effects of cryo-freezing before mechanical testing. These parameters were not examined because insufficient cells would have been available while also controlling passage number (P2). In conclusion, delivering a controlled mechanical stimulus pattern produced a significant improvement in the sub-failure, biomechanical properties of the cell-sponge construct but no corresponding benefit for the cell-gel construct. These differences in response to a single mechanical signal suggest that mechanical stimuli should be tuned and even optimized to a construct s constituents to enhance its stiffness. Building on this optimization theme, future studies should also investigate whether construct and repair stiffness can be most enhanced by stimulating constructs with patterns that mimic physiologic strains. The current study mimicked only the physiological peak strain (2.4%) previously recorded by our group [132]. Using principles of functional tissue engineering [17], we are now investigating the effects of other in vivo tendon strain components in order to most effectively precondition MSC-based constructs to their future in vivo environment [131, 132, 150, 151] and to determine if faithfully recreating the in vivo strain signal optimizes these response measures. It is worth noting, of course, that other outcome measures will likely also need to be optimized including ECM composition (type I and type III collagen [185, 186] and decorin [37]), rates of collagen 80

97 synthesis [35], degree of collagen crosslinking [185], and cell phenotype and metabolism [187]. Such in vitro systems will ultimately offer the opportunity to also understand structure-function relationships in tissue engineered constructs. 81

98 Chapter 6 Effect of Scaffold Material, Construct Length and Mechanical Stimulation on the In Vitro Stiffness of the Engineered Tendon Construct 5 Introducing mesenchymal stem cell (MSC)-seeded collagen constructs into loadprotected wound sites in the rabbit patellar and Achilles tendons significantly improves their repair outcome compared to natural healing of the unfilled defect. However, these constructs would not be acceptable alternatives for repairing complete tendon ruptures because they lack the initial stiffness at the time of surgery to resist the expected peak in vivo forces thereafter. Since the stiffness of these constructs has also been shown to positively correlate with the stiffness of the subsequent repairs, improving initial stiffness by appropriate selection of in vitro culture conditions would seem crucial. In this study we examined the individual and combined effects of collagen scaffold type, construct length, and mechanical stimulation on in vitro implant stiffness. MSCs were extracted from the iliac crest of six 1 yr-old female NZW rabbits. Eight constructs were created from each of these cell lines to study two levels each of scaffold material (collagen gel vs collagen sponge), construct length (short vs long), and mechanical stimulation (stimulated vs non-stimulated). Constructs assigned for mechanical stimulation were stimulated after two days in culture using a single profile (one pulse every five minutes to a peak strain of 2.4% for 8 hours/day for 12 days). Non-stimulated constructs remained in 5 This manuscript has been submitted to Journal of Biomechanics and is currently being reviewed. 82

99 the incubator for the same period. All 14-day constructs were failed in tension to determine their in vitro linear stiffness. Although all three treatment factors influenced construct linear stiffness, construct length had the greatest effect compared to the other two. Longer, stimulated, cell-sponge constructs showed the highest in vitro linear stiffness and the shorter, non-stimulated, cell-gel constructs showed the lowest stiffness. A significant interaction was also found between length and stimulation. We now plan in vivo studies to determine if high stiffness constructs generate high stiffness repairs 12 weeks after surgery Introduction Mesenchymal stem cell (MSC)-seeded collagen constructs are currently being used by our group to repair patellar tendon defect injuries in the rabbit model. These cellassisted repairs exhibit 200% to 700% of the maximum force and stiffness at 12 weeks after surgery compared to values for natural repair tissues [16, 22]. However, the success obtained using these constructs may also be due to the load-protected nature of the central tendon wound site which would not be present after a complete tendon rupture. In fact, 12 weeks after surgery, the strength and stiffness of the cell-gel repairs are still less than 1/3 of normal tendon values [11, 17, 18] and thus not strong enough to resist the peak in vivo forces that would be carried across a wound site after full tendon rupture. One strategy to improve these repair biomechanical properties would be to enhance the in vitro biomechanical properties of the cell-based construct. Previous studies from our laboratory have demonstrated that in vitro construct stiffness positively correlates with repair stiffness 12 weeks after surgery [22]. Thus, we may be able to 83

100 improve construct stiffness and then repair outcome by optimizing specific construct parameters while still in culture. Three attractive parameters worth investigating would be the individual and combined effects of the type of scaffold material presented to the cells, the length of the cell-scaffold construct during maturation, and the influence of mechanical stimulation before surgery. Although various synthetic and biologically-derived scaffold materials have been used as delivery vehicles to create engineered tissue constructs, our group has selected type I collagen gels and sponges to repair defects in the rabbit patellar tendon model [11, 14-16, 22]. We chose collagen-based scaffolds since type I collagen represents approximately 95% of all collagen present in natural tendon and 65 80% of its dry mass [30, 31]. In an unpublished pilot study, we initially evaluated different collagen scaffolds (fibers, gels, films and sponges) in culture to decide which would be most suitable to use in a tissue-engineered tendon repair. We found that MSCs showed greater penetration and viability in collagen sponges compared to collagen gels as well as better construct integrity and handling. Another approach to improving a construct s in vitro biomechanical properties would be to increase its initial length in culture. This improvement has been demonstrated analytically and experimentally by various investigators [26, 27]. Eastwood et al [26] reported that cells in fibroblast-populated collagen lattices (FPCL) with high aspect ratio showed more parallel alignment with the axis of the applied load compared to the low aspect ratio lattices. Using Small Angle Light Scattering methods to visualize MSC-collagen gel constructs, Nirmalanandhan et al [27] showed that increasing the length of the constructs (to reduce the influence of end effects according to St. Venant s 84

101 principle) improved collagen fiber alignment and in vitro construct stiffness. Whether long constructs formed from collagen sponges might show similar improvements was not studied, however. Mechanically stimulating constructs to precondition them to their in vivo environment before surgery is consistent with functional tissue engineering principles formulated by one of the co-authors (DLB) [17]. Such in vitro construct stimulation in a bioreactor is known to increase cell proliferation [ ] as well as protein transcription and synthesis [141, ]. Mechanical stimulation of individual tendon cells produces similar results by increasing DNA synthesis [142] and expression of novel genes [ ]. Using a pneumatically actuated bioreactor, we also found that compared to non-stimulated controls, mechanically stimulated cell-collagen sponge constructs not only improved construct stiffness before surgery, but also significantly improved repair biomechanical properties 12 weeks after their introduction into rabbit central PT wound sites [22]. Surprisingly, these mechanically-induced benefits in construct stiffness were not observed when cell-gel constructs received identical stimulation profiles [25]. These findings would suggest that scaffold material may influence how cells react to mechanical stimulation in 3-D configurations. Although scaffold material, construct length and mechanical stimulation individually affect the in vitro biomechanical properties of a cell-based construct, no study, to our knowledge, has examined the combined and potentially interactive effects of these three key factors and which one(s) is(are) most important to control. Based on our previous studies, we hypothesized: that 1) mechanical stimulation of longer cell-sponge constructs, with stiffer and more aligned scaffolds, would result in the highest in vitro 85

102 stiffness among the treatment groups studied, and 2) that interactions would be found among the three treatment factors Methods Experimental Design MSCs were extracted from the iliac crest of six skeletally mature, female, New Zealand white (NZW) rabbits (Myrtles Rabbitry, Thompson Station, TN). Eight constructs from each cell line were created to study the effects of two levels of three factors, scaffold material (collagen gel vs. collagen sponge), length (short vs. long), and mechanical stimulation (stimulated vs. non-stimulated) (Table 6.1). The short and long constructs were created in specially-designed silicone dishes having two different post-topost lengths (11 and 51 mm) but identical dish widths (11mm), depths (6mm) and post diameters (4mm) (Figure 6.1). Half of the cell-sponge (one short and one long) constructs and half of the cell-gel (one short and one long) constructs were then mechanically stimulated. The other short and long cell-sponge and cell-gel constructs from each cell line remained in an incubator without stimulation. After two weeks of incubation, all constructs were failed in tension to determine their linear stiffness, the response measure used to contrast different treatments. 86

103 Figure 6.1. The short and long constructs were created in specially-designed silicone dishes having two different post-to-post lengths [A) 11 and B) 51 mm] but identical dish widths (11mm), depths (6mm) and post diameters (4mm). Scaffold Material Mechanical Stimulation Long Length Short Collagen Gel Collagen Sponge Yes n = 6 n = 6 No n = 6 n = 6 Yes n = 6 n = 6 No n = 6 n = 6 Table 6.1. Eight constructs from each of the six (n=6) rabbit cell lines were created to study the effects of two levels of three factors, scaffold material (collagen gel vs. collagen sponge), length (short vs. long), and mechanical stimulation (stimulated vs. nonstimulated). Harvest of MSCs MSCs were isolated from the bone marrow of one-year old New Zealand White rabbits using previously described methods [11, 13, 15]. Following general anesthesia, bone marrow aspirates were taken from the iliac crest of these rabbits. Complete MSC 87

104 growth medium [25 ml of 89% DMEM-lg, (Gibco, Carlsbad, CA), 1% antibiotic/antimycotic (Gibco), and 10% FBS (Atlanta Biologics, Norcross, GA) of selected lot was added to each aspirate. Bone marrow samples were centrifuged and supernatant was discarded. The precipitated pellet was plated at 17x10 6 to 22x10 6 cells/100mm dish. These cells were then cultured in an incubator (Steri-Cult Model 3033, Forma Scientific, Marietta, OH) at 37 C, 95%RH, 5% CO 2 and fed with complete MSC growth media twice weekly until ~95% confluent. MSCs were then harvested, counted, and re-plated at 0.5x10 6 MSCs/dish. After passage one MSCs were cryogenically preserved at 135 C in DMEM with 20% FBS and 10% DMSO [188]. Frozen MSCs were removed from liquid nitrogen storage, thawed in a 37 C bath, quickly diluted with 10 ml of DMEM, rinsed and then subcultured as described above. Preparation of short and long cell-gel constructs A cell suspension at 1.168x10 6 cells/ml was mixed with neutralized type I purified bovine collagen gel (Cohesion Technologies, Palo Alto, CA). The final cellseeding density was 0.1x10 6 MSCs/ml in a solution containing mg collagen/ml. To account for differences in well sizes (Figure 6.1), the volume of the cell-gel mixture for the long constructs (3.5 ml) was 2.5 times that of the mixture for the short constructs (1.4 ml). Long and short constructs were fed high glucose DMEM supplemented with 10% fetal bovine serum and 5% ascorbic acid and incubated at 37 o C, 5% CO 2 for 14 days. 88

105 Preparation of short and long cell-sponge constructs Type I collagen sponges (P1076 collagen pads) were provided by Kensey Nash Corporation (Exton, PA). Sponges were sterilized using a minimum of 25 kgy of gamma irradiation. Each sponge (94% pore volume; 62 µm mean pore diameter) was then cut to fit in the base of each of four wells in a silicone dish containing two restraining posts protruding from the base (Figure 6.1). Two 4 mm-diameter holes were also created, permitting each sponge to be positioned in the base of each well [22]. Sponges were then soaked in phosphate buffered saline (PBS, Gibco BRL/Life Technologies Inc., Gaithersburg, MD) for 24 hours. Short constructs were created by injecting a cell suspension aliquot (0.4 ml) containing 0.14x10 6 MSCs and long constructs by injecting a cell suspension aliquot (1 ml) containing 0.35x10 6 MSCs on the collagen sponge. All constructs were placed in an incubator (37ºC, 5% CO 2, 95% RH) for 2 weeks and fed three times weekly with high glucose DMEM, 5% ascorbic acid and 10% FBS. Mechanical Stimulation After 2 days of initial incubation, the silicone dishes containing the constructs assigned for mechanical stimulation were stretched once every 5 minutes to a peak strain of 2.4% for 8 hours/day for 12 days using previously published methods [21, 22, 25]. After a total of 14 days, both mechanically stimulated and non-mechanically stimulated constructs were placed in cryovials and stored in a -80ºC freezer until mechanical testing. Total storage time in the freezer was approximately two weeks. 89

106 Biomechanical Evaluation Adapting previously published methods [22, 25, 27], constructs were slowly thawed to room temperature over an hour and failed in tension using a custom materials testing system (100R6, TESTRESOURCES, Shakopee, Minnesota). Briefly, specimens were secured in serrated metal grips and then stretched to failure at a rate of 10%/sec based on a 10 mm gage length for both short and long constructs. Linear stiffness was calculated from the linear region of the force-elongation curve generated by the testing system [25, 27]. Statistical Analysis The two levels (gel and sponge; short and long; stimulated and non-stimulated) of the three respective factors (scaffold material, length, stimulation, respectively) constituted a 2 3 factorial design. The fact that the cells from the same animal were used in every treatment combination permitted correlations among the responses. Hence, the data was analyzed using an ANOVA model with effects from the main factors and their interactions with animal treated as random variables. The residuals were tested for normality and homoscedasticity and found to satisfy these criteria at a 5% level of significance. All conclusions regarding the significance of different treatments were made at p < One-way analysis of variance was then conducted to compare the 8 treatment combinations with respect to linear stiffness and least significant difference (LSD) multiple comparisons were conducted when differences were significant. 90

107 6.3. Results Our statistical analysis revealed that length (p = 0.006) has the greatest influence on the linear stiffness compared to scaffold material (p = 0.019) and mechanical stimulation (p = 0.014) (Figure 6.2). In addition we found a significant interaction between length and stimulation (p = 0.039). Surprisingly, no significant interactions were found among any of the other combinations of the three factors studied. Longer constructs showed higher linear stiffness compared to shorter constructs (0.047 ± N/mm vs ± N/mm; mean ± SEM; p = 0.006) (Figure 6.2). For the same type of scaffold material (gel or sponge) and mechanical stimulation condition (stimulated or non-stimulated), long constructs demonstrated significantly higher linear stiffness compared to short constructs (Figure 6.3 & Table 6.2). Figure 6.2. Length (p = 0.006) has the greatest influence on the linear stiffness (mean ± SEM; N/mm; N = 6) compared to scaffold material (p = 0.019) and mechanical stimulation (p = 0.014), at 14 days in culture. 91

108 Group 1 Group 2 p-value Long, Non-stimulated, Gel Short, Non-stimulated, Gel Long, Stimulated, Gel Short, Stimulated, Gel Long, Non-stimulated, Sponge Short, Non-stimulated, Sponge Long, Stimulated, Sponge Short, Stimulated, Sponge Table 6.2. Long constructs (Group 1) demonstrated significantly higher linear stiffness compared to the corresponding short constructs (Group 2) for the same type of scaffold material (gel and sponge) and mechanical stimulation condition (stimulated and nonstimulated) Short; Non-stimulated Short; Stimulated Linear Stiffness (N/mm) Long; Non-stimulated Long; Stimulated * * * * * * * 0 Gel Sponge Figure 6.3. Long, mechanically-stimulated MSC-seeded collagen sponge constructs demonstrated the highest linear stiffness and the short, non-stimulated MSC-seeded collagen gel constructs demonstrated the lowest linear stiffness. No difference was found between long, non-stimulated constructs and short, mechanically-stimulated constructs for both sponge (p = 0.914) and gel (p = 0.878). 92

109 The MSC-seeded collagen sponge constructs showed significantly higher linear stiffness compared to the MSC-seeded collagen gel constructs (0.045 ± N/mm vs ± N/mm; mean ± SEM; p = 0.019) (Figure 6.2). For the same length (short or long) and mechanical stimulation condition (stimulated or non-stimulated) collagen sponge constructs demonstrated significantly higher linear stiffness compared to the corresponding collagen gel constructs (Table 6.3 & Figure 6.3). Similarly, mechanically stimulated constructs exhibited significantly higher linear stiffness compared to non-stimulated constructs (0.045 ± N/mm vs ± N/mm; mean ± SEM; p = 0.014) (Figure 6.2). Mechanical stimulation significantly improved the linear stiffness of cell-sponge constructs (p = and 0.03 for short and long, respectively). However, no improvement in the linear stiffness was observed in the cell-gel constructs due to mechanical stimulation (p = and for short and long, respectively). Group 1 Group 2 p-value Long, Non-stimulated, Sponge Long, Non-stimulated, Gel Long, Stimulated, Sponge Long, Stimulated, Gel Short, Non-stimulated, Sponge Short, Non-stimulated, Gel Short, Stimulated, Sponge Short, Stimulated, Gel Table 6.3. Collagen sponge constructs (Group 1) demonstrated significantly higher linear stiffness than the corresponding collagen gel constructs (Group 2) for the same length (short and long) and mechanical stimulation condition (stimulated and non-stimulated). Long, mechanically-stimulated, MSC-seeded collagen sponge constructs demonstrated the highest linear stiffness (0.066 ± N/mm; mean ± SEM) and the short, non-stimulated MSC-seeded collagen gel constructs demonstrated the lowest linear stiffness (0.011 ± N/mm; mean ± SEM) among all the groups tested (Figure 6.3, 93

110 Table 6.4). Interestingly, no difference was found between long, non-stimulated constructs and short, mechanically-stimulated constructs for both sponge (p = 0.914) and gel (p = 0.878) constructs. Group 1 Group 2 p-value Long, Stimulated, Sponge Long, Non-stimulated, Sponge 0.03 Long, Stimulated, Gel Long, Non-stimulated, Gel < Short, Stimulated, Sponge Short, Non-stimulated, Sponge < Short, Stimulated, Gel < Short, Non-stimulated, Gel < Table 6.4. Longer, mechanically stimulated, cell-sponge constructs showed significantly higher linear stiffness compared to any other group using least significant difference (LSD) multiple comparisons Discussion This study not only determined whether the linear stiffness of tissue engineered constructs could be improved in vitro by altering scaffold material, changing construct length, and introducing mechanical stimulation, but also identified the most dominant of the three factors studied. The fact that changes in all three treatment factors improved the construct s in vitro linear stiffness after 14 days in culture indicates their individual influence on a maturing structure before implantation at surgery. The fact that length and mechanical stimulation showed interactive effects indicates the potential synergistic benefit of introducing multiple factors simultaneously before surgery. For each of the two levels studied for construct length (short vs long), scaffold material (collagen gel vs collagen sponge), and mechanical stimulation (stimulated vs non-stimulated), our statistical analysis reveals that length of the construct has the 94

111 greatest influence on its linear stiffness compared to scaffold material and mechanical stimulation. These findings suggest that the in vitro stiffness of the engineered construct can be improved to a greater extent by simply increasing its length than by using expensive collagen sponges as scaffold materials or providing sophisticated mechanical stimulation. Equally interesting to note is that the stiffness of long, non-stimulated cell sponge constructs match the stiffness of short, mechanically-stimulated cell-sponge constructs (Figure 6.3) that are currently being used in in vivo studies by our group. The significant interaction seen between length and mechanical stimulation suggests that not only the individual effects but also the combined effect of these two factors improves the linear stiffness. The combined effects due to scaffold material, length, and mechanical stimulation likely explain why the long, mechanically-stimulated cell-sponge constructs demonstrated the highest linear stiffness and the short non-stimulated cell-gel constructs exhibited the lowest linear stiffness. Our results support our previous study that increasing the length of a cell-gel construct improves its in vitro biomechanical properties [27]. These biomechanical results also suggest that longer constructs, regardless of the collagenous scaffold material used, exhibit improved fiber alignment as we previously demonstrated after imaging cellgel constructs using a Small Angle Light Scattering (SALS) technique [19, 26, 27]. Our group has already demonstrated that the in vitro biomechanical properties and subsequent repair biomechanics of engineering tendons can be improved by replacing compliant collagen gels with stiffer collagen sponges [22, 25]. Results from the current study are consistent with these previous reports in that short cell-sponge constructs are stiffer than the corresponding cell-gel constructs after 14 days in culture. Our multi- 95

112 factorial experimental design also allowed us to examine how long constructs are affected by the use of compliant gels vs. collagen sponges. The similar trend we observed in the stiffness data for these two long constructs would suggest that regardless of construct length, the use of a collagen sponge improves the in vitro linear stiffness of the constructs more than the use of a collagen gel. This increase in stiffness found after using cellsponge constructs could certainly be attributed to their higher collagen content. Cellsponge constructs had significantly greater collagen content (estimated mg/ml) compared to the cell-gel constructs (2.6 mg/ml), despite the fact that the same number of cells was used at the same passage number to create both cell-gel and cell-sponge constructs. One interesting finding from this study was the differential effect of mechanical stimulation on cell-gel vs. cell-sponge constructs. The fact that a controlled mechanical stimulus stiffened the short cell-sponge constructs but had no effect on the short cell-gel constructs is consistent with the results from three of our previous studies [21, 22, 25]. However, the further differential effect of mechanical stimulation on long cell-sponge vs. cell-gel constructs is new and shows the overarching influence of a scaffold material in controlling the effect of an external stimulus on its constituent cells. Thus, it appears that MSCs in both long and short cell-sponge constructs positively respond to a mechanical stimulus but MSCs in long and short cell-gel constructs do not. The difference in collagen content in these scaffolds could have certainly influenced the response of the MSCs to the same mechanical stimulus. Whether additional increases in scaffold and extracellular matrix (ECM) stiffness might further enhance the effect of the stimulus on the construct s cells is not known, but concerns about stress shielding of a cell-based 96

113 construct (e.g. in inducing a change in cell phenotype and deposition of ectopic bone in soft tissue sites [15]) and poor alignment of collagen fibers (resulting from poor contraction due to increased levels of collagen [172]) would suggest that construct stiffness and amount of collagen present in the construct be altered to optimize stimulus effect in culture. We chose linear stiffness as our response measure for two primary reasons. Monitoring a subfailure parameter like linear stiffness supports our goal of creating tendon constructs that can withstand functional forces associated with activities of daily living rather than failure forces associated with trauma [17, 127]. Our group previously reported that the rabbit s patellar [132], Achilles [131] and flexor digitorum profundus tendons [151] were loaded to 11%, 19% and 28% of normal failure forces during the most vigorous activities studied, respectively. These peak forces load these tendons only into the lower aspects of the linear regions of their force-elongation curves. We also chose linear stiffness because we were concerned that the small aspect (length: width) ratio of the shorter specimens ( 2:1) might have affected the failure properties of the constructs (St. Venant effects). In light of these reasons, our treatment effects are best compared using the stiffness in the linear region of the force-elongation curve. We suspect that the increases we observed in stiffness due to length, scaffold material and mechanical stimulation are primarily due to improved fiber alignment and increased expression of important structural genes like collagen types I and III. We hypothesize that increasing construct length increased linear stiffness by primarily improving geometric fiber alignment as we demonstrated recently using the cell-gel constructs [27]. We also suspect, but must still test, whether applying a mechanical 97

114 stimulus and stiffening the scaffold material improves linear stiffness by upregulating collagen types I and III gene expression [141]. Our future studies will test several hypotheses by not only evaluating these constructs for expression of structural genes like collagen types I and III, but for important fibrillar assembly genes like decorin and fibronectin using real time rtpcr [189] and for changes in collagen fiber alignment by Small Angle Light Scattering (SALS) techniques [19, 189]. Our study is not without limitations. 1) We did not measure the forces acting on the constructs during mechanical stimulation nor did our strain stimulation protocol necessarily allow the cell-gel construct sufficient time to recover. To address these problems, we are designing systems to simultaneously measure construct force and deformation in culture to detect time-dependent changes in construct load during mechanical stimulation. Future mechanical stimulation studies will utilize a new multistation, electromechanical stimulation system with individual force transducers that permit tracking of real-time construct stiffness during stimulation. 2) We did not control the collagen content in the cell-gel and cell-sponge constructs. This was because, unlike the collagen sponge, the collagen gel was not available at higher concentrations. Since this difference in collagen content could have influenced the response of MSCs to the same mechanical stimulus, our future studies will examine the effects of mechanical stimulation separately for different gel concentrations and for different sponge concentrations. These studies will permit us to carefully test the effect of collagen scaffold concentration on the construct stiffness in culture. In conclusion, we have demonstrated that three factors (construct length, mechanical stimulation, and scaffold material) all influence the resulting in vitro linear 98

115 stiffness of a tissue engineering construct in culture. Of the eight groups studied, longer, stimulated cell-sponge constructs show a significantly higher in vitro linear stiffness. Construct length and mechanical stimulation also show significant interactions. As a result of this study, we are currently investigating whether these improved constructs, when placed in a patellar tendon defect site, might also significantly improve repair outcome post surgery and whether the in vitro construct stiffness continues to correlate with in vivo repair parameters like linear stiffness and modulus. 99

116 Chapter 7 Optimizing Mechanical Stimulus to Improve In vitro Biomechanical Properties of Tissue Engineered Tendon Constructs 6 In vitro mechanical stimulation has been reported to induce cell alignment and increase cellular proliferation and collagen synthesis. Our group has previously reported that in vitro mechanical stimulation of tissue-engineered tendon constructs significantly increases both construct stiffness and repair biomechanics after surgery. However, these studies used a single mechanical stimulation profile, but the stimulation is composed of multiple components whose individual and combined effects on construct properties remain unknown. Thus, the purpose of this study was to understand the relative importance of a subset of these components on construct stiffness. To try and optimize the resulting mechanical stimulus, we used an iterative process to vary peak strain, cycle number, and cycle repetition while controlling cycle period (1 sec), rise and fall times (25% and 17% of the period, respectively), hours of stimulation/day (8 hours/day) and total time of stimulation (12 days). Two levels of peak strain (1.2 or 2.4%), cycle number (100 or 3000 cycles/day), and cycle repetition (1 or 20) were first examined. Higher levels of peak strain and cycle number were then examined to optimize the stimulus using Response Surface Methodology. Our results indicate that constructs stimulated with 2.4% strain, 3000 cycles/day and one cycle repetition produced the stiffest constructs. Given 6 This manuscript has been submitted to Tissue Engineering and is currently being reviewed. 100

117 the significant positive correlations we have previously found between construct stiffness and repair biomechanics at 12 weeks post surgery, these in vitro enhancements offer the prospect of further improving repair biomechanics Introduction Soft tissue injuries to tendons, ligaments and capsular structures represent almost 45% of the 32 million musculoskeletal injuries that take place each year in the US [41]. These injuries have been traditionally treated by replacing the damaged tissue with a graft replacement or directly repairing the damaged tissue using sutures [5, 47-49]. While some procedures are successful, many repairs and graft reconstructions fail, requiring revision surgery. The cause of these re-injuries is often due to the large forces sustained during activities of daily living [52, 66]. Tissue engineering offers an attractive alternative to these clinical problems and has been studied in a variety of animal models. In tendon tissue engineering, mesenchymal stem cells (MSC) are multiplied in culture and then often mixed with collagen gels. The cells contract the collagen gel over time in culture to create constructs that can be implanted into the wound sites to repair rabbit patellar and Achilles tendon injuries [11, 13, 24, 105]. The repairs containing MSC-gel composites, introduced into the patellar tendon repair site, result in 33% improvements in the material properties and 50% improvements in maximum force and stiffness at 12 weeks compared to values for natural repair [11]. However, studies also indicate that the repair biomechanics for these tissue engineered constructs do not yet meet in vivo loading demands [17, 131, 132]. 101

118 One strategy to further improve repair outcomes is to mechanically stimulate tissue engineered constructs to direct differentiation and promote extra-cellular matrix (ECM) development [141, 178]. In vitro application of stress and strain to cells in both two- and three-dimensional environments induces cell alignment [141, 190] and increases cell proliferation [141, 142], secretion of growth factors (TGF-β, bfgf, and PDGF) [143] and collagen synthesis [141]. Mechanical stimulation studies using tenocytes show similar results, exhibiting increased DNA synthesis [142] and expression of novel genes not typically seen in tendon cells [145]. Altman et al [141, 154] delivered translational and rotational strains to ligament constructs, and after 21 days of stimulation, MSCs expressed mrna for types I and III collagen and tenascin-c, all of which are indicators of a fibroblastic phenotype. These studies show the ability to positively affect cellular activity through in vitro stimulation. Our group has been using functional tissue engineering principles [17, 127] to create and mechanically stimulate mesenchymal stem cell-collagen gel-collagen sponge constructs in culture to accelerate repair of central-third defects in the rabbit patellar tendon (PT). We initially found that stimulating these constructs increased the linear stiffness and linear modulus of the 12 week repairs to as much as 80% and 40% of normal PT values, respectively [21]. Stimulated repairs also matched the tangent stiffness of normal PT not only up to peak in vivo forces recorded during inclined hopping (100 N) [132], but to 25% beyond this level (125 N). When the scaffold was replaced with a commercial collagen sponge and the gel was removed, the stimulated, 12-week, stem cell-collagen sponge repairs actually matched the tangent stiffness of normal PT up to 50% beyond peak in vivo force levels for inclined hopping (150 N, 32% of PT failure 102

119 force) [22, 132]. This stimulus also significantly increased the stiffness of these stem cellcollagen sponge constructs at 14 days in culture and the linear stiffness of the subsequent rabbit patellar tendon repairs at 12 weeks post surgery [22]. However, these studies examined a single mechanical stimulus which is, in fact, composed of multiple components (e.g. peak strain, frequency, duration, etc.) whose individual and interactive effects are unknown. Studies are now needed to identify and optimize which subset of these components should be applied to more rapidly achieve tangent stiffness up to 40% of normal PT failure force, the highest in vivo forces we have measured in patellar tendon in the goat model [150]. Therefore, the purpose of this study was to understand how three of the components of the mechanical stimulation profile (peak strain, cycle repetition and total number of cycles per day) influence construct stiffness in culture. Our decision to improve and even optimize construct stiffness in culture was based on the positive correlations we previously found between in vitro construct stiffness and repair stiffness 12 weeks post surgery [21, 22]. Using Response Surface Methodology (RSM) [29] to distinguish the effects of altering peak strain, cycle number per day and cycle repetition, we hypothesized that optimizing these components of the stimulus, within physiologic levels, would increase linear construct stiffness Experimental Design & Methods Treatment Factors & Levels for In vitro Mechanical Stimulation Studies (Table 7.1) Previous studies from our laboratory demonstrated that a trapezoidal strain profile significantly improved both construct and repair stiffness [21, 22, 25]. Therefore, in this study, we used this trapezoidal strain profile to optimize the three treatment factors or 103

120 components (Figure 7.1). The ranges for these components are chosen so as to include the physiological levels within them. 1) Four levels of strain amplitude (0.6, 1.2, 2.4, and 4.8% peak strain) were chosen in the range of peak strains estimated from peak in vivo force measurements in the rabbit PT model [131, 132]. 2) Three levels of cycle repetition (1, 20, 50) were also selected to contrast the effects of single, discrete events from the effects of 20 and 50 sequential cycles that might either enhance ECM synthesis on the one hand or desensitize the cell to further stimulation on the other. 3) Four levels of strain cycles per day (10, 100, 3000, and 10,000) were also proposed. We chose the 10 cycle/day level because it falls between the 4 cycles/day that trigger bone cells to maintain bone mass and the 12 to 36 cycles/day that trigger new bone remodeling and periosteal bone formation [191]. We selected the 100 cycle/day level because we previously found it to increase construct stiffness in culture [21, 22, 25]. Finally, we chose the 3000 and 10,000 cycle/day levels because the first is similar to the estimated number of daily walking cycles taken by typical human subjects [192], and the second corresponds to more strenuous activities. Figure 7.1. A trapezoidal stimulus consists of multiple components (e.g. peak strain, frequency, duration, etc.) whose effects are, to date, unknown. 104

121 Response Surface Methodology (RSM) An efficient design strategy was used to test the prohibitively large number of treatment factors and levels (4*3*4 = 48 combinations) (Table 7.1). RSM [29] permits simultaneous evaluation of treatment effects and interactions using a reduced factorial design involving only two levels per factor (2 3 = 8 groups) (Table 7.1), a center point to create the response surface, and then repeated evaluations of individual or targeted levels of the three factors. These targeted levels are determined from a mathematical equation, which defines the surface produced by the reduced factorial design and center point. Analysis of the results for the full response surface model dictates the direction to move for selecting and then optimizing each factor. Treatment Factor Potential Treatment Levels Levels for Full Factorial Design Peak strain (%) 0.6, 1.2, 2.4, , 2.4 Cycle repetition 1, 20, 50 1, 20 Cycle number/day 10, 100, 3K, 10K 100, 3K Table 7.1. Response Surface Methodology permits simultaneous evaluation of treatment interactions using a reduced factorial design involving only two levels per factor (2 3 =8 groups). Iteration 1. First, two levels each of peak strain (1.2 and 2.4%), cycle number (100 and 3000 cycles/day), and cycle repetition (1 and 20) were examined while controlling cycle period (1 sec), rise and fall times (25% and 17% of the period, respectively), hours of stimulation/day (8 hours/day) and total time of stimulation (12 days). One intermediate level of each factor (peak strain 1.8%, 1550 cycles/day and 10 cycle repetitions) was then selected to produce a center point for a response surface to determine the next iterations using RSM. 105

122 MSCs were isolated from the iliac crest of ten one-year-old female New Zealand White rabbits and then culture expanded until passage two using previously described methods [11, 13, 15]. Constructs were created by seeding 0.14 x 10 6 MSCs into a type I collagen sponge (P1076, Kensey Nash Corporation, Exton, PA) cut to fit in the bottom of the wells in the specially designed silicone dish [16, 22, 172]. After allowing the cells to migrate into the collagen sponge for 2 days in the incubator (37ºC, 5% CO 2, 95% RH), the constructs were stimulated using a pneumatically-controlled mechanical stimulation system [21, 22, 25]. All constructs were fed 3 times a week with advanced DMEM, 5% ascorbic acid, 1% antibiotic/antimycotic, 1% Glutamax and 5% FBS. After 2 weeks of incubation, constructs were failed in tension at a strain rate of 10%/sec using an electromechanical testing system (TestResources Inc.; Shakopee, MN). The forceelongation curves were plotted to determine the construct s linear stiffness [25, 27]. A three-way ANOVA was performed with peak strain, cycle number and cycle repetition as fixed factors. For each strain level, a two-way ANOVA was conducted with the other two as fixed factors. Post hoc testing was conducted using Tukey s tests at a significance level of p < Iteration 2. In order to extend the response surface, two targeted groups (the next treatment level combinations that arise from the analysis) were identified based on the response surface from iteration 1 and a mathematical equation derived by curve fitting the data from iteration 1. These two groups included, A) 2.7% peak strain, 4450 cycles/day and 1 cycle repetition and B) 3.15 % peak strain, 5900 cycles/day and 1 cycle repetition. 106

123 Iteration 3. To complete the response surface around the targeted groups from iterations 1 and 2, we analyzed two more groups including: C) 3.15% peak strain, 3000 cycles/day and 1 cycle repetition, and D) 2.4% peak strain, 5900 cycles/day and 1 cycle repetition (Figure 7.2) Iteration 1 Iteration 2 Iteration 3 Cycle number (per day) Response Surface Peak strain (%) Response Surface 2 Figure 7.2. Experimental design showing all three iterations, two response surfaces and the interactions between peak strain and cycle number Results Iteration 1. Increasing peak strain produced the only significant individual change in construct linear stiffness (0.03 ± % to 0.05 ± %; mean ± SEM; p < 0.001; Figure 7.3). A significant interaction was also found between peak strain and cycle number (p = 0.039), but no interactions were found between peak strain and cycle repetition or between cycle number and cycle repetition. While showing 107

124 no effect at 1.2% peak strain, increasing cycle number did increase construct stiffness at 2.4% peak strain from 0.05 ± cycles/day to 0.06 ± cycles/day; p=0.012; Figure 7.4). Stimulating the constructs with intermediate levels of the 3 factors (1.8% peak strain, 1550 cycles/day and 10 cycle repetitions) produced a stiffness value of ± N/mm (Table 7.2). Figure 7.3. Increasing peak strain produced the only significant change in construct linear stiffness (0.03 ± N/mm (1.2%) to 0.05 ± N/mm (2.4%); mean ± SEM; p<0.0001). The following mathematical equation defined the resulting response surface: 2 2 Stiffness = A* strain + B * cycle number + C * strain * cycle number + D * strain + E * cycle number + F where A = 0, B = 0.019, C = 0.006, D = 0.01, E = 0.006, and F =

125 Figure 7.4. While changing cycle number had no effect at 1.2% peak strain, construct stiffness increased at 2.4% strain from 0.05 ± N/mm (100 cycles/day) to 0.06 ± N/mm (3000 cycles/day; p = 0.012). Group Peak strain (%) Cycle number/day Cycle repetition Linear stiffness ± SEM (N/mm) ± ± ± ± ± ± ± ± ± Table 7.2. The conditions of each group tested in iteration 1 and their corresponding stiffness. 109

126 This final response surface showed that imposing 2.4% strain, 3000 cycles/day and 1 cycle repetition most stiffened the construct (0.068 ± 0.009; N/mm; mean ± SEM; Figure 7.5). Analysis of the results for the first response surface model also revealed that linear stiffness could be further improved by increasing the peak strain and cycle number. Figure 7.5. The initial response surface for peak strain and cycle number per day suggests that increasing both components could further improve linear stiffness of the construct. Iterations 2 and 3. The first two targeted groups (groups A and B) exhibited lower linear stiffness (0.03 ± and ± 0.004; N/mm; mean ± SEM) compared to the 3000 cycle per day, 2.4% peak strain group that produced the highest stiffness in iteration 1 (Table 7.3). Groups C and D also showed lower linear stiffness (0.036 ± and ± 0.004; N/mm; mean ± SEM) compared to this 3000 cycle per day, 2.4% peak 110

127 strain group (Table 7.3). This final response surface also showed 2.4% strain, 3000 cycles/day and 1 cycle repetition most stiffened the constructs (Figure 7.6). Group Peak strain (%) Cycle number/day Cycle repetition Linear stiffness ± SEM (N/mm) A ± B ± C ± D ± Table 7.3. Higher peak strains (2.7% and 3.15%) and cycle numbers/day (4450 and 5900) did not improve the stiffness. Figure 7.6. The broader response surface indicates that peak linear stiffness is produced by imposing 2.4% peak strain and 3000 cycles/day in culture to the construct for 12 days. 111