Tissue Engineering Strategies to Improve Tendon Healing and Insertion Site Integration

Home , Cellular adaptation

A dissertation submitted to the

Division of Research and Advanced Studies of the University of Cincinnati

in partial fulfillment of the of the requirements for the degree of

DOCTOR OF PHILOSOPHY (Ph.D.)

in the Department of Biomedical Engineering of the College of Engineering and Applied Science

2011

Kirsten Rose Carol Kinneberg

B.S., University of Minnesota, Twin Cities, MN, 2006

Committee Chair: Jason T. Shearn

Abstract

Tendon and ligament tears and ruptures remain common and significant musculoskeletal injuries. Repairing these injuries continues to be a prominent challenge in orthopaedics and sports medicine. Despite advances in surgical techniques and procedures, traditional repair techniques maintain a high incidence of re-rupture. This has led some researchers to consider using tissue engineered constructs (TECs).

Previous studies in our laboratory have demonstrated that TEC stiffness at the time of surgery is positively correlated with repair tissue stiffness 12 weeks post-surgery. This correlation provided the rationale for implanting a soft tissue patellar tendon autograft (PTA) to repair a central-third defect in the rabbit patellar tendon (PT). The PTA was significantly stiffer than previous TECs and matched the stiffness of the normal central-third PT. Accordingly, we expected a significant improvement in repair tissue biomechanics relative to both natural healing

(NH) and TEC repair. At 12 weeks, treatment with PTA improved repair tissue stiffness relative to NH. However, PTA and NH tissues did not differ in maximum force, modulus or maximum stress. Additionally, neither repair group regenerated normal zonal insertion sites.

To enhance integration at the tendon-to-bone insertion site, PTA repairs were 1) given up to 26 weeks to recover and 2) augmented at the patellar and tibial insertions with mesenchymal stem cell (MSC)-collagen gel biologic augmentations (BAs). The role of the native cell population in PTA healing was also tested by de-cellularizing the PTA at surgery. We found that osteotendinous integration improved with recovery time for both de-cellularized PTA (dcPTA) and PTA repairs. However, biomechanical properties were only affected by recovery time for dcPTA repairs. Despite the changes in biomechanical properties demonstrated by dcPTA repairs, biomechanical properties did not vary between dcPTA and PTA repairs at any time point. We

also found that MSC-collagen gel BAs did not enhance osteotendinous integration or repair tissue biomechanical properties relative to PTA repairs at 12 weeks post-surgery.

Overall, the repair tissue biomechanics of our mechanically pre-conditioned MSC- collagen sponge TECs were approximately twice the biomechanical properties for both PTA repairs and NH. This result warranted additional experiments to further improve in vitro TEC stiffness and mRNA expression with the objective to enhance tendon healing. We investigated the effects of pore size, scaffold composition and mechanical pre-conditioning in vitro on MSC- collagen sponge TEC stiffness and mRNA expression levels for genes of interest. Our initial results indicated that for collagen sponge TECs, pore size did not affect linear stiffness and mechanical stimulation only enhanced stiffness when chondroitin-6-sulfate was incorporated into the collagen sponge.

The goals of my research were to 1) understand the effects of the resident cell population and healing time on PTA integration into bone, 2) develop a biologic augmentation that would improve tendon insertion site development, and 3) improve TEC-mediated PT healing. Future studies need to investigate the effects of combining biological and mechanical factors at the insertion site on PTA integration and also validate our in vitro-to-in vivo predictors for MSC- collagen sponge TEC repair using updated sponge materials.

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Acknowledgements1

First and foremost, I would like to thank my advisor, Dr. Jason Shearn. I could not have asked for a better advisor. I would like to specifically thank Dr. Shearn for trusting me to work independently but always being there when I had questions or needed advice on how to move a project forward. Dr. Shearn pushed me to my limits and helped me realize my full potential and capability as a researcher.

I would also like to extend a very special thank you to Dr. David Butler. His methodologies have forever shaped the way I approach research. I feel privileged to have learned research design from Dr.

Butler and also thank him for helping me better communicate my research. With the help of both Drs.

Shearn and Butler I have learned to appreciate the idea of “keep it simple.”

I am also grateful to Dr. Keith Kenter for serving as a member of my dissertation committee. He has brought a unique aspect to my research and helped me to think critically in terms of the clinical problem. His comments and advice will help ensure our continued projects have clinical impact and move the field of tissue engineering in a positive direction.

I would also like to acknowledge the significant contributions of Dr. Marc Galloway (Cincinnati

Sportsmedicine). His research ideas have had a great impact on this dissertation and I thank both him and

Dr. Shearn for the opportunity to work on these projects. Dr. Galloway, along with Dr. Kenter, has helped ensure that our laboratory maintains a focus on clinically relevant problems. I would also like to thank Dr.

Galloway for always creating a positive and fun atmosphere.

Completing the projects described in this dissertation would not have been possible without the significant contributions of Mrs. Cindi Gooch. She has taught me the invaluable skills of cell culture technique and animal care for surgery. Cindi has not only taught me the „tricks of the trade‟ for cell culture but she has also been a good friend and a good source of advice during my time at the University of Cincinnati.

______1 This work was supported by two grants from the National Institutes of Health (NIH AR46574-10 and NIH AR56943-02) and also by a grant from the National Science Foundation (NSF 0333377) awarded to the University of Cincinnati.

A note of gratitude is also due to my colleagues. Specifically, I would like to thank Dr. William

Ball, Dr. Natalia Juncosa-Melvin, Dr. Victor Nirmalanandhan, Dr. Kumar Chokalingam, Nathaniel

Dyment, Andrea Lalley, Jennifer Hurley and Abdul Sheikh for helping to conduct experiments, understand data or discuss ideas for future studies. Their contributions and ideas were invaluable in completing this dissertation.

Sincere thanks also to Lori Beth Derenski, Linda Moeller, Shelly Smith, Kathryn Siefert, and

Michelle Montoya. They have come to my rescue on more than one occasion and I am very grateful for their efforts.

Finally, I would like to thank my family and Patrick Mihalik. They now know more than they ever wanted to know about mesenchymal stem cells, collagen sponges, and patellar tendon autografts. I appreciate their kindness in being there for me and helping me though the difficulties of research. My mom, dad, sister, grandma and Patrick have always supported and encouraged me even when I doubted myself. I would like to also thank Patrick, especially, for his patience while I completed this dissertation.

Life is a succession of lessons which must be done to be understood. - Ralph Waldo Emerson

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Table of Contents

Abstract ii

Acknowledgements v

Table of Contents

List of Tables 2

List of Figures 3

Chapter 1 Literature Review 4

Chapter 2 Research Objectives and Hypotheses 23

Chapter 3 Effects of Implanting a Soft Tissue Autograft in a Central-Third Patellar Tendon Defect: Biomechanical and Histological Comparisons 38

Chapter 4 Effects of Recovery Time and the Role of the Native Cell Population on Insertion Site Formation and Repair Tissue Biomechanics of the Patellar Tendon Autograft 53

Chapter 5 MSC-Collagen Gel Augmentation to Enhance Osteotendinous Integration of a Patellar Tendon Autograft 70

Chapter 6 Chondroitin-6-Sulfate Incorporation and Mechanical Stimulation Increase MSC-Collagen Sponge Construct Stiffness 95

Chapter 7 Discussion 113

Chapter 8 Recommendations for Future Studies 120

Bibliography 124

List of Tables

Table 1 Repair Tissue Dimensions [mean (SEM)] for Whole and Central-Third PT 43

Table 2 Biomechanical Properties (Mean ± SEM) of Natural Healing, Patellar Tendon Autograft, Tissue Engineered Construct and Normal Central-Third PT 45

Table 3 Repair Tissue Dimensions (Mean ± SD) for Whole and Central-Third PT Repair Tissues 59

Table 4 Biomechanical Properties (Mean ± SD) of Patellar Tendon Autograft and De-Cellularized Patellar Tendon Autograft Repairs 61

Table 5 Gene Names*, TaqMan® Assay Identification (ID) Numbers, and Amplicon Length for All Tested Genes 79

Table 6 Biomechanical Properties (Mean ± SD) of Patellar Tendon Autograft and PTA+BA Repairs and Normal Central-Third PT 83

Table 7 Repair Tissue Dimensions (Mean ± SD) for Whole and Central-Third PT 87

Table 8 Experimental Design 99

Table 9 Biomechanics and Gene Expression Data for MSC-Collagen Sponge TECs [Mean (SEM)] Cultured for Two Weeks Statically and with Mechanical Stimulation 106

List of Figures

Figure 1 Schematic of tendon hierarchical structure 6

Figure 2 Experimental design flow chart 25

Figure 3 SEM images of small (A), medium (B), and large (C) pore size scaffold materials 36

Figure 4 Force-displacement curves (mean ± SEM) 44

Figure 5 H&E staining and IHC staining for collagen type III (Col3) of the tendon mid-substance for both NH and PTA after 12 weeks of healing 47

Figure 6 Histological images of patellar and tibial tendon-to-bone insertion sites 48

Figure 7 Maximum force and stiffness plots as percent of normal 62

Figure 8 Tibial insertion sites for de-cellularized PTA, PTA and native struts (NS) at 6, 12 and 26 weeks post-surgery 64

Figure 9 Experimental design overview for MSC-collagen gel BA surgical implantation and in vitro optimization 75

Figure 10 Alkaline phosphatase activity per cell 85

Figure 11 Gene expression of MSC-collagen gel BAs normalized to GAPDH 86

Figure 12 Force-displacement curve for PTA and PTA+BA repairs Tibial insertion of a PTA+BA repairs, PTA repair and native struts (NS) at 12 weeks. 88

Figure 13 Tibial insertion of a PTA+BA repairs, PTA repair and native struts (NS) at 12 weeks 90

Figure 14 Scanning electron microscopy of scaffold materials 103

Figure 15 Linear stiffness normalized by static control (mean ± SEM) 106

Figure 16 mRNA expression of collagen types I and III normalized by GAPDH (mean±SEM) 107

Figure 17 mRNA expression of decorin and fibronectin normalized by GAPDH (mean ± SEM) 107

Chapter 1

Literature Review

Tendons and ligaments are soft connective tissues that serve to transmit load.1 Tendons join muscle to bone and utilize the relatively large forces generated by muscle to facilitate motion at the joints. Ligaments join bone to bone and sustain relatively smaller forces in order to stabilize joints from excessive motion. While the patellar tendon anatomically connects the patella to the tibial tuberosity of the tibia, it is also continuous with the quadriceps tendon and functions to increase the efficiency of the quadriceps muscle in knee extension; these factors justify studying this structure as a tendon rather than a ligament. The patellar tendon is the focus of study in this dissertation.

1. Tendon Structure Tendons are organized in a hierarchical structure (Fig. 1).2-6 The most basic level is tropocollagen (the collagen molecule). Tropocollagen is composed of three polypeptide strands, also known as alpha chains. The primary structure of the alpha chains is the amino acid sequence: glycine (Gly)-X-Y. The X and Y positions are most commonly filled with proline and hydroxyproline, respectively.7 The three alpha chains each form individual left-handed helixes which, in turn, twist around each other to form a right-handed coiled coil that is stabilized by numerous hydrogen bonds.5 Tropocollagen subunits spontaneously self-assemble into collagen fibrils with a stagger distance (D) of 67nm.5 The staggered array of tropocollagen is stabilized within the fibrils by covalent cross-linking catalyzed by the enzyme lysyl oxidase (LOX).5, 8

Collagen fibrils are the smallest structural unit of tendon with diameters ranging from 10 –

150nm.3, 9 Immature animals have uniformly distributed small fibril diameters while mature animals have a bimodal distribution of diameters.8-12

The mechanism of how fibrils advance to their final form in tendon is not well understood. It has been proposed that fibrils are assembled as intermediate structures before being assembled by fibroblasts both longitudinally and laterally into fibril bundles known as fibers.12, 13 Fibers are then packaged into both primary and secondary bundles (fascicles).

Collagen fiber diameters range from 5 – 30µm in rat tail tendons up to 300µm in human tendons.3, 14

The orientation of collagen fibrils and fibers helps impart functional strength to the tendon. Collagen fibrils and fibers (including both primary and secondary bundles) are not only oriented parallel to the axis of tension but also transversely and horizontally with fibrils and fibers crossing each other to form intricate patterns.3 This complex organization acts as a buffer against multi-directional and rotational loads during movement and activity.3 Collagen fibrils and fibers also display a crimped configuration when unloaded.3, 9, 13 The crimped configuration is removed when the tendon is strained; the initial 2% of tendon strain corresponds to the elongation of the crimp.15, 16 The crimping may also serve as a buffer against injury.15, 16

There are three connective tissue sheaths which supply vasculature, lymphatics, and nerves to the tendon: the endotenon, the epitenon, and the paratenon.3, 5, 17, 18 Collagen fibers, primary fiber bundles, and secondary fiber bundles (fascicles) are each surrounded by endotenon.17 Fascicles are grouped into tertiary fiber bundles within the tendon unit and they are surrounded by epitenon.17 In human tendons, fascicle diameters range from 150 – 1000µm and tertiary bundle diameters range from 1000 – 3000µm (the diameters of both fascicles and tertiary bundles are directly related to the macroscopic size of the tendon structure).3 Endotenon is a

reticular network of thin collagen fibrils in a crisscross pattern while epitenon is a relatively dense network of collagen fibrils 8 – 10nm in thickness.3 Tendons are bound by the paratenon

(synovial sheath for some tendons). Paratenon is a loose, fatty, areolar connective tissue that functions as an elastic sleeve.3, 230

Tropocollagen Collagen Collagen Primary Secondary Fiber Tertiary Fiber Bundle Molecule Fibril Fiber Fiber Bundle and Tendon Unit Bundle (Fascicle) Blood Vessels, α1 α2 Nerves & Lymphatics Epitenon & Epitenon Cells: Tenocytes/ Tenoblasts

Paratenon X Synovial Y Gly Cells

Molecule Length = 280nm D = 0.64nm Dark Band = Gap Light Band = Overlap

Epitenon α1 α2 Endotenon & Endotenon Cells: Crimp Tenocytes/ Tenoblasts

Tertiary Fiber Proteoglycans, Proteoglycans, Bundle Glycoproteins Glycoproteins

Figure 1. Schematic of tendon hierarchical structure. This example shows the hierarchy for collagen type I (col1α1). Collagen types II and III (found in cartilage and tendon sheaths, respectively) also form fibrils. However, collagen type II fibrils do not form fibers or bundles and collagen type III fibrils will form fibers but not bundles. Modified from refs 3, 9, 13, 19.

2. Tendon Composition Structurally, tendon is comprised of cells, fibers, and ground substance.24, 230, 231

However, tendon is about 50-75% water, by mass.4, 28 Of the 30% dry mass of tendon, the primary constituents are fibers and ground substance.9 Collectively, fibers and ground substance are known as extracellular matrix (ECM).24 Although cells are not a major component of tendon, their role in maintaining the structural and functional integrity of the tissue should not be discounted.24 Tendons also contain vasculature which provide nutrients and may serve to help maintain the biomechanical integrity of tendons.18, 20, 21

2a. Tendon Composition - Cells The primary cell type in tendon is the fibroblast and it accounts for 90-95% of the total tendon cellular element3. Fibroblasts originate from undifferentiated mesenchymal stem cells

(MSCs).5, 22 The remaining cellular element of tendon is comprised of fibrochondrocytes and chondrocytes in the insertion site and regions exposed to compressive loads (these cells develop from tendon fibroblasts by metaplasia23, 24, most likely in response to increased mechanical/compressive forces23-26), the synovial cells which line the paratenon (or tendon sheath for some tendons)24, 27, and the vascular cells found in the blood supply of the endo- and epitenon3, 24. In pathological conditions, other cell types including inflammatory cells and myofibroblasts may also be observed3 (see Wound Healing).

The fibroblastic tendon cell population can be broken down into two distinct groups: those found in the 1) tendon proper (within fascicles) and 2) paratenon/tendon sheath, epitenon and endotenon (Fig. 1).24, 27 Additionally, the fibroblastic population can exist in two states: active and quiescent.5 The term fibroblast is generally reserved for immature, active cells while

mature, quiescent cells are typically referred to as fibrocytes.5 Tendon fibroblasts can also be classified as tenoblasts and tenocytes for active and quiescent cells, respectively.

Tenoblasts vary in size with lengths ranging from 20 – 60µm and widths of 8 – 20µm.3,

11, 28 Within their cytoplasm, tenoblasts have a well developed rough endoplasmic reticulum

(RER; responsible for protein synthesis5) and Golgi apparatus (prepares proteins from the RER for cellular secretion5) but very few mitochondria (generate chemical energy in the form of adenosine triphosphate, ATP5).5, 11, 28 The numerous cytoplasmic organelles and granular material resembling matrix ground substance surrounding these cells all support the high metabolic activity (intense synthesis of matrix components) of the tenoblast.3 With age, the cell- to-matrix ratio gradually decreases in the tendon and tenoblasts transform into tenocytes.11, 28

Tenocytes are much longer than tenoblasts with lengths of 80 – 300µm.28 The nucleus-to- cytoplasm ratio increases in tenocytes until the cytoplasm is almost completely occupied by the nucleus; the increased nucleus-to-cytoplasm ratio in tenocytes reflects their decreased metabolic activity.3, 28 Tenocytes are still metabolically active, just not at the same level as tenoblasts and not using the same pathways.3 As cells convert from tenoblasts to tenocytes, there is a shift in metabolic pathways from aerobic to more anaerobic energy production.3, 29, 230 The low metabolic rate of tendon cells, along with their well developed capacity for anaerobic energy production, enables the tendon to carry loads and remain under tension for prolonged periods of time without putting the tissue at risk of ischemia, necrosis and subsequent injury.3 However, the low metabolic rate limits recovery and healing after injury.3, 234

Tendon cells are responsible for the synthesis of ECM components such as collagen, elastin, glycosaminoglycans (GAGs), proteoglycans (PGs), and multiadhesive glycoproteins.5, 231

They are also involved in the production of growth factors which influence cellular growth and

differentiation.5, 232, 233 The more specific role(s) of each cell population are well documented in wound healing (discussed in Fibrillogenesis and Wound Healing, below).4, 92, 93 However, the phenotypic expression does not appear well understood outside of that environment (genotype plus environment and random variation lead to phenotypic expression). The same follows for mast cells, macrophages, and plasma cells. Understanding the day-to-day functions of these cell populations as well as their individual and collective roles in tendon maintenance could help tissue engineers optimize repair strategies.

2b. Tendon Composition – Fibers Connective tissue fibers are assemblies of proteins which have polymerized into elongated structures.7 There are three main categories of fibers in tendon: collagenous, reticular and elastic.5, 14 Collagen and reticular fibers both belong to the collagen fiber system while elastic fibers belong to the elastic system.5

The predominant fiber of tendon ECM is collagen; up to 85% of tendon‟s dry mass is collagen.28, 30 Of the total collagen in tendon, about 95% is collagen type I and 5% is collagen type III (reticular fibers) or V.7, 30-33 The primary biomechanical function of collagen type I in tendon is to resist tensile loads.222 Collagen type I is limited to the tendon proper and paratenon but is not found in the endotenon or epitenon.34, 35 Collagen type III reticular fibers are typically composed of loosely packed thin fibrils and arranged in a mesh-like network that is continuous with collagen type I fibers.14, 34, 36 Collagen type III is limited to the epitenon, endotenon and paratenon sheaths where it forms a fine supporting mesh that helps impart tissue elasticity.26, 34, 35

Also, because collagen type III can be synthesized quickly, expression of this protein is upregulated during early stages of tendon wound healing.34, 37 Collagen type III serves to stabilize the structure until tissue remodeling and synthesis of collagen type I begin.37 Collagen

type V can be found intercalated into the core of collagen type I and helps regulate fibril diameter.9, 38, 39 Accordingly, the decrease in tendon strength and stiffness seen with age is associated with a decrease in tendon fibril size and upregulation of collagen type V.40

Elastic fibers account for < 3% of the dry mass of tendons.5, 31, 41 Elastic fibers have a high content of the protein elastin and stretch easily in response to tension.5, 14 Elastic fibers were only found in 10% of healthy human tendons but were found consistently in the Achilles tendons of young and old rabbits and in the tendons of rats.3, 28, 42 The function of elastic fibers is not well understood. They may serve a role in restoring the crimp configuration of fibrils and fibers after muscle contraction or tendon stretch.15

While collagen types I, III and V predominate in tendon, collagen types II, X, XII, and

XIV are also present.12, 25, 43, 44 Collagen types II and X can be found in a region of fibrocartilage where tendon integrates with bone.25, 43, 44 Collagen type II is found in cartilage and functions to resist compressive pressure.25, 43, 44 Collagen type X is found in hypertrophic and mineralizing cartilage but may only be present at certain stages of tendon development.25, 43, 44 The role of collagen types XII and XIV in fibril formation (fibrillogenesis) is not well understood but it has been proposed that they are involved in linear fibril growth and matrix assembly.12

2c. Tendon Composition – Ground Substance Tendon ground substance is composed of glycoproteins, glycosaminoglycans (GAGs), and proteoglycans (PGs).5, 45 Glycoproteins commonly found in tendon include fibronectin, tenascin-C, undulin, and the thrombospondin (TSP) family (including cartilage oligomeric matrix protein, also known as COMP and TSP-5).3, 4, 46-49 While the functions of many glycoproteins are not yet understood, fibronectin belongs to the subclass of adhesive glycoproteins thought to bind either macromolecules or cell surfaces together 46 while tenascin-C

and the TSP family belong to the subclass of matricellular proteins which are thought to function as adaptors and modulators of cell-matrix interactions50. Fibronectin can be found on the surface of collagens (in the endotenon) 34, 51 where it plays a key role in extra cellular matrix (ECM)-cell interactions such as adhesion, migration, growth, and differentiation.34, 52 During the early stages of wound healing, fibronectin can be found wide spread throughout the tissue matrix (granulation tissue) and surrounding connective tissues. Fibronectin may also play a role in forming and/or remodeling granulation tissue during tendon healing by mediating posttranslational collagen fibril modifications and assembly.53 Tenascin-C is found abundantly in musculoskeletal tissues regions of mature animals where high mechanical forces are being transmitted from one tissue to another, such as the tendon-bone junction at the tendon insertion site.46, 50, 54 Accordingly, tensacin-C is regulated by mechanical strain and upregulated in tendinopathy.48 Tenascin-C is involved in cellular adhesion and migration50 and may also play a role in fiber alignment and orientation49. COMP is thought to play a major role in tendon development.19, 55 COMP binds collagen type I and can act to increase the rate of collagen type I fibril formation.19, 55

Proteoglycans are a class of glycoproteins composed of a protein core and one or more covalently bound glycosaminoglycan (GAG) chains.5, 56 Glycosaminoglycans (GAGs) are carbohydrates composed of repeating disaccharides (“two sugar”) forming long un-branched polysaccharides (“many sugars”).5 In mature tendon, PGs are typically arranged orthogonal to the collagen fibrils (Fig. 1).13, 57 In immature tendons, PGs are typically found either orthogonal or parallel to the fibril.13, 57, 58 The PG content of tendon, generally 1-2% of the dry mass31, varies by region and depends on the mechanical loading environment. In the bovine flexor tendon, the

PG content changes from 3.5% of the tendon‟s dry mass in the compression-bearing region to about 0.2 – 0.5% in the tension-bearing region.59-61 Additionally, while compressive loads in

tendon induce production of aggrecan (typically found in cartilage and fibrocartilage), tensile loads stimulate production of decorin (a primary PG of tendon).62

GAGs commonly found in tendon include chondroitin sulfate (CS) and dermatan sulfate

(DS).45, 57, 63, 64 PGs found in tendon include, but are not limited to, decorin, lumican, biglycan, fibromodulin, hyaluronan, epiphycan and keratocan.64-68 The primary PG in tendon is decorin which, along with lumican, biglycan, fibromodulin and others, belongs to the class of small leucine rich proteoglycans (SLRPs).59, 64, 65, 67 Decorin consists of a protein core and one bound

GAG chain; decorin binds the GAG chondroitin sulfate in bone and dermatan sulfate in tendon.45, 69 It is the presence of GAG which gives PGs their unique properties.45

2d. Tendon Composition – Vasculature Supply Tendons typically receive their blood supply from intrinsic systems at the musculotendinous and osteotendinous junctions (MTJ and OTJ, respectively) and extrinsic systems via the endotenon, epitenon, paratenon and/or synovial sheath18, 20, 70. The patellar tendon (PT) also receives a blood supply from the infrapatellar fat pad.20 The intrinsic and extrinsic sources of vasculature appear to complement each other. Intrinsic systems from the

MTJ and OTJ do not fully penetrate the tendon; sources from the MTJ are typically limited to the proximal one-third of the tendon while those from the OTJ are limited to the insertion zone of the tendon.18 Alternatively, extrinsic sources originating within the paratenon form complex arterial branches that penetrate the epitenon and form an intratendinous vasculature network.20, 70

Vasculature may play an important role in maintaining tendon biomechanical properties.

In the Achilles tendon (AT), the vasculature supply is reduced in the mid-substance of the tendon, the area most prone to rupture for the AT.18, 21 Additionally, the PT receives extrinsic vasculature from both the paratenon on its anterior surface and infrapatellar fat pad on its

posterior surface.20 However, the vasculature supply of the PT is relatively lower near the insertions20, the area most prone to rupture for the PT (most prone to rupture at the patellar insertion, specifically). A limited blood supply may contribute to a reduced healing capacity such that when the tendon experiences microtears71, these areas of reduced vasculature are less capable to repair the tissue.72 Accordingly, ischemia (inadequate blood supply) has been implicated in tendinopathy, along with a myriad of other factors.69, 71

3. Biosynthesis of Collagen and Regulators of Fibrillogenesis

The biosynthesis of collagen type I has been well studied.5, 8, 12 Collagen synthesis initiates within the cell but is completed in the extracellular environment.5, 8, 12 Within the cell, polypeptide alpha chains are assembled and modified (signal peptides are cleaved) to form procollagen.5 Once the peptide chain has reached a minimum length, hydroxylation of selected prolyl and lysis residues are catalyzed by specific enzymes.5 This critical step in the process of collagen synthesis depends on the presence of ascorbic acid. Galactosyl and glucosyl are attached to specific hydroxylysyl residues (glycosylation) before the procollagen molecules are assembled into a triple helix.5 The procollagen triple helix is then transported to the extracellular environment where procollagen peptidases cleave the registration peptides and form the insoluble tropocollagen molecule.5 Tropocollagen then goes on to form the hierarchical structure of tendon described earlier in Tendon Structure through a process of fibrillogenesis.5, 8, 12, 19, 38

As mentioned previously, fibrillogenesis is the development and assembly of collagen fibrils (Fig. 1). It has been proposed that fibrillogenesis of type I collagen in tendon is regulated by PG present in the ECM.45, 67, 68 During tendon maturation, a predominance of chondroitin sulfate has been associated with the presence of thin fibrils during early phases of tendon fiber

development.45 At a critical point, there is a sharp decline in the ratio of PG to collagen and the tissue concentration of both chondroitin sulfate and dermatan sulfate is reduced.45, 57 It is at this point that the predominant GAG changes from chondroitin sulfate to dermatan sulfate and a phase of rapid growth is initiated where existing fibrils increase in diameter faster than new fibrils are created.57, 59 It has been proposed that the onset of rapid growth is contingent upon the decrease in CS such that if a sufficient level is sustained, the outcome would be a tissue composed of thin collagen fibers.59

It is important for tissue engineers to understand collagen synthesis and fibrillogenesis because these processes regulate the form and function of the tissue we are trying to recapitulate in vivo. Understanding the spatial and temporal biologic regulators of collagen synthesis and assembly allows us to test for these markers in vitro before our products are implemented in vivo. If a tissue engineered construct is not demonstrating some aspect of collagen synthesis in vitro, we have to ask: Will this TEC produce a functional tendon in vivo?

4. Tendon Biomechanical Properties Tendon biomechanical behavior can be visualized with a stress-strain curve.4 As described earlier, tendons display a crimp pattern when they are at rest.73 The initial 2% of the stress-strain curve (the “toe” region) is concave up and thought to represent a flattening of the crimp.15, 74 However, the physiological and functional relevance of the toe region has been debated with some arguing that it may simply be an artifact of assuming the normal resting length of the tendon occurs at zero load.31 As the tendon is elongated further, it forms a linear region.75 Tendon behaves elastically when loaded up to 4% strain (change in length / initial length).76 However, microscopic failure begins beyond 4% strain and even if the tendon is unloaded, it will not return to its original length. Tendon begins to yield at the end of the linear

region and by 8-10% strain, macroscopic failure occurs.77 Maximum strain of tendon is typically

12-15% and maximum stress (force / cross-sectional area) is usually 50-150MPa.15, 78, 79

However, maximum strain and stress can vary depending on the structure. For the normal central-third PT, the maximum strain is 16.0 ± 1.5% and the maximum stress is 100.7 ± 16.0MPa

(mean ± SD).80, 81

Tendons function at 30-40% of their ultimate strength during normal activity82-85 but higher loads are placed on the tissues during more strenuous activities like jumping and running.

In the human Achilles tendon, forces corresponding to 12.5 times body weight (9kN) were recorded during running.86 This force exceeds the single-load ultimate tensile strength of the

Achilles tendon15 which raises the importance of strain rate. Animal testing has demonstrated that with high strain rates, there was an increased load to failure and a predominance of soft tissue failures relative to slower strain rates where bone avulsion was the primary failure location.78 It is important to note that, in this study, there was a one hundred-fold difference between the fast and slow strain rates which emphasizes that the biomechanical properties of collagen are not highly strain rate dependent.78 Nonetheless, tendons are at highest risk for failure if tension is applied quickly and obliquely.4

5. Prevalence of Tendon Injuries In the United States, an estimated 1 in 4 people report at least one musculoskeletal injury each year87 with roughly 16 million injuries reported for soft connective tissues such as tendon, ligament, and capsular structures.88 The lower extremity accounts for the majority of activity- related injuries, most commonly in the knee. Knee injuries and Achilles tendon ruptures account for roughly 23% and 6% of activity-related injuries, respectively.87

The most commonly injured tendons include the flexor and extensor tendons of the hand

(incidence of 4.83 and 18/100,000 per year, respectively)89, the Achilles tendon (12-18/100,000 per year)89, 90, and the rotator cuff tendons (3.73/100,000 per year)89. One study reported a 34% overall incidence of rotator cuff tears.91

6. Wound Healing Tendon healing occurs in three overlapping phases: inflammation, proliferation and remodeling. Tendon healing is mediated by both intrinsic and extrinsic cellular mechanisms.

Intrinsically, it has been established that collagen is initially synthesized by cells from the epitenon.92 Additionally, while cells from the endotenon synthesize collagen later than those from the epitenon92, endotenon tenocytes also secrete larger and more mature collagen.4 Lastly, when compared to cells from the epitenon and endotenon, cells from the tendon sheath generate less collagen and glycosaminoglycans but also proliferate more rapidly and therefore may play a larger role in healing.92 While the role of each cell population is relatively well understood for wound healing, the phenotypic expression of each cell population has not been extensively studied outside of this environment (genotype plus environment and random variation lead to phenotypic expression). Extrinsically, mast cells and macrophages each play important roles as chemical mediators of inflammation and wound healing. These cells interact with fibroblasts by releasing growth factors, enzymes and cytokines which mediate the inflammatory response, proliferation, synthesis of collagen and other ECM components, ECM re-absorption, and angiogenesis.4, 92, 93 Seemingly most relevant to our objective(s) as tissue engineers is that both mast cells and macrophages are capable of secreting transforming growth factor-β (TGF-β) which has been associated with promoting scar formation in wound healing rather than tissue regeneration.92, 93

It is important to understand how cells interact in vivo because similar interactions may occur with the implanted cells in tissue engineered constructs. By understanding what occurs during natural healing of tendon injuries, we can potentially design and adapt our TECs to avoid the fibrotic scar tissue generated by naturally healing and, instead, produce functional tendon tissue.

7. Current Repair Strategies for Tendon Injury Wound healing in tendon does not restore the biomechanical strength or stability of the native structure. Therefore, treatment is generally required for patients who wish to resume normal activity levels. Without proper management, these injuries may progress and reduce muscle and/or joint function which can lead to increased pain and disability for the patient.

Tendon and ligament injuries are generally treated using one of three techniques: non-operative, surgical repair (also known as direct repair), or graft replacement.

Non-Operative Tendon Repair Non-operative methods generally implement a brace or cast to immobilize the joint.

These methods are associated with a higher rate of re-rupture and may produce a less functional outcome when compared to surgical repair.94 This is true for AT rupture, where both non- operative and direct repair are common.95, 96 Surgical repair of AT rupture, which is generally recommended for young, athletic patients, has a higher risk of complication (approximately 3 –

11% of patients90, 97, 98) but a lower re-rupture rate (approximately 2% to 5%), when compared to non-operative treatments90, 97, 98. Although re-rupture rates are low for traditional surgical repair,

AT ruptures are often misdiagnosed (diagnostic failure rate of 20 – 30%) and delayed repair of

AT ruptures can downgrade surgical outcomes by at least 20%.95

Surgical Tendon Repair Surgical repair involves suturing the damaged ends of tendons/ligaments back together or to the surrounding connective tissue. Surgical repair does not consistently restore function. The failure rate for flexor tendon repair is 4% to 13% with the most common cause of failure being overloaded sutures.99-101 For the Achilles tendon (AT), open operative repairs have a re-rupture rate of roughly 2-4%.97, 102 However, the failure rate of open surgery is approximately double that for percutaneous repair (4.3% and 2.1%, respectively) and post-operative casting followed by functional bracing significantly reduces re-rupture, when compared to bracing alone (2.4% versus 12.2%, respectively).102 Rotator cuff repair outcomes also vary tremendously, with failure rates ranging from 11% to 95% at two years following repair.103-105 Recurrent tears and poor rates of healing in the rotator cuff may be due to the initial size of the tear, the severity of the damage (number of rotator cuff tendons involved in the injury), the chronicity of the tear, the quality of the tissues at the time of surgery, and the age and general health of the patient. 103-105

Similar to the AT, rotator cuff repair can also require an immobilization period. Patients treated for tendon injuries can be immobilized from physiologic loads during healing which makes them more susceptible to joint stiffness and muscle atrophy.106 Overall, surgical repair is limited by the amount of tissue disruption and the intrinsic healing capacity of the tissue.1 If the degree of damage is too great or if the tissue is not capable of repairing itself, a graft may be required.

Graft Replacement Tendon grafts play an integral role in many orthopaedic procedures. Grafts serve to span defects in chronic tendon injuries as well as provide strength and stability to repair sites. Tendon grafts are also utilized extensively for ligament repair and reconstruction with good repeated success.

Tendon grafts can be obtained from the patient (autografts) or a donor (allografts).

Autografts are limited by tissue availability, procurement site morbidity, and secondary pain at the harvest site.94 Allografts are limited by cost, availability, disease transfer, and immuno- rejection. 94 Additionally, the use of allografts raises the issues of relative graft strength, rates of incorporation, and graft survival.94, 107

The success of tendon grafts varies by application. For example, tendon grafts have demonstrated repeated success in repairing the ACL. However, there is a high rate of failure associated with reconstruction of the rotator cuff tendons. Bone-patellar tendon (PT)-bone

(BPTB) allografts and autografts have proven successful for ACL repair with rupture rates of 7 –

13% and 5 – 7%, respectively.108-111 Alternatively, when allograft (fresh-frozen and irradiated patellar, Achilles, or quadriceps tendon) was used to repair massive rotator cuff tears, 100% of patients had radiographic failure at follow-up.112, 113 The success of the BPTB graft is primarily attributed to 1) the initial mechanical integrity of the graft and 2) the transfer of bone ends with the soft tissue.94 Bone-to-bone healing progresses more rapidly than soft tissue-to-bone healing so transferring the bone ends allows for faster incorporation of the graft.94 Unlike the use of

BPTB grafts, current repair techniques for rotator cuff tears require tendon-to-bone fixation without the transfer of bone ends. These techniques generally do not regenerate a tendon-to-bone insertion site and the high failure rates are often attributed to poor graft incorporation.103, 114, 132

Proper formation of a tendon-to-bone insertion site is critical to the success of tendon grafts as the lack of tissue integration is reported to compromise stability and long-term clinical outcome.114-118, 132

8. Tissue Engineering and Functional Tissue Engineering for Tendon Repair Frequent injuries and the inconsistent results of traditional repairs have led researchers to consider novel treatments, like tissue engineering, which offer both advantages and challenges.

Tissue engineered constructs (TECs) have an advantage over autografts because they eliminate problems of procurement site morbidity and pain associated with tissue harvest.94, 107 Also, unlike traditional allografts, TECs avoid any risk of disease transmission, by utilizing autologous cells.94, 107 The challenges associated with TECs include cost, scaling up between animal models, and large scale production.228, 229 There are also stringent FDA regulations for devices which contain a cellular component, as most TECs do.228, 229

TECs are generally composed of scaffold materials including collagen80, 81, 130, 144-146, 157,

164, 186, 189, 191, 223-226, silk119-121, and biodegradable polymers122-125. Type I collagen scaffolds have been investigated for tendon tissue engineering strategies because it is the primary structural component of tendon. However, collagen TECs in the form of gels80, 186, 189, 191, 223, sponges81, 130,

144, 145, 146, 157, 164 and small intestinal submucosa (SIS)224, 225, 226 are limited by inadequate mechanical strength at the time of implantation. Additionally, the cross-linking methods used to increase the initial strength of collagen fibers tend to inhibit cellular migration into the scaffolds, thus limiting tissue ingrowth and remodeling.227 Silk has been widely studied in vitro because of its biocompatibility, slow degradation, and capacity to produce high mechanical properties in twisted and cabled scaffolds.119-121 Biodegradable polymers, such as poly(lactic acid) (PLA), poly(glycolic acid) (PGA), and their copolymer, poly(DL-lactic-co-glycolic acid) (PLGA), have been studied because they are biocompatible, can be processed into unique 3D geometries, possess relatively good mechanical strength, and their degradation rate can be carefully controlled.122-125 However, the behavior of silk is not well characterized in vivo and the hydrophobicity and lack of signaling molecules on polymers limits cell adhesion, proliferation,

and subsequent function.119-121 Additionally, when used as ACL replacements, both silk and biodegradable polymer scaffolds have been shown to rapidly lose mechanical integrity in vivo.120, 122, 124 This may be attributed to the processing of the fibers leaving an internal space too small to allow proper tissue ingrowth and neo-ligament formation.121

There are several challenges to fabricating TECs for orthopaedic applications common to all scaffold materials. The first is to generate mechanical properties adequate to carry the required mechanical loads. The second is to possess the appropriate biochemical cues for sufficient cellular migration from the repair/wound margins to allow for scaffold degradation, tissue ingrowth and subsequent remodeling, and proper formation of tendon/ligament insertion sites. The third is to recapitulate native biomechanics and maintain mechanical integrity over time in vivo. The last challenge raises the question of whether we need to match normal properties in order to consider a tissue engineered (TE) repair successful.

Defining evaluation benchmarks for orthopaedic repairs is a significant challenge for both clinical and tissue engineered repairs. However, functional tissue engineering can help define critical markers of success. Functional tissue engineering (FTE) was founded on the principles of tissue engineering but incorporates functional loading demands into the design and evaluation of TECs.126-129 Our implementation of FTE principles has recently been updated to include biological markers of success in addition to biomechanical ones. It is important for tissue engineers to understand that biology and biomechanics are not mutually exclusive when designing successful TECs.

To help define the biomechanical markers of a successful repair outcome, our group has recorded patterns and peak in vivo forces in numerous tendons in the rabbit82-84 and goat models85. As a result, we have identified the following biomechanical criteria of success. First,

the repair tissue must sustain a maximum force greater than that required for activities of daily living which is 100N.82-84 Second, the repair must match the tangent stiffness of the normal tendon up to 40% of failure force to accommodate for more strenuous activities.85 By including the tangent stiffness we limit the amount of tissue elongation which can occur before reaching the required loads. It is important to consider force and elongation together to understand if our repairs are functioning similar to the native structure.

The biological markers of success are not as well defined as the biomechanical markers.

However, in working with a group of developmental biologists we hope to understand the mechanisms regulating both normal tendon development and healing. By contrasting these methods we hope to elucidate important regulatory factors and enhance our TECs for better repair tissue generation.

In addition to defining evaluation benchmarks for repair outcomes, it is also important to identify benchmarks of success for in vitro results. If correlations can be established between in vitro and in vivo response measures, the in vitro parameters can be used as a screening tool before proceeding in vivo. Studies by our laboratory have demonstrated that in vitro TEC stiffness is significantly and positively correlated with repair tissue stiffness 12 weeks post- surgery.81, 130 This finding has significantly impacted our laboratory and how we choose to conduct in vivo studies. We are currently working to establish in vitro biological markers that correlate with biological and/or biomechanical response measure in vivo. In vitro predictors of in vivo outcomes are an important aspect of tissue engineering research because they can be used to limit in vivo testing and help select only the most promising candidates of success for in vivo implantation. In vitro screening not only helps save time, money and valuable resources but also helps ensure the field of tissue engineering makes significant progress in the right direction.

Chapter 2

Research Objectives and Hypothesis

This dissertation describes tissue engineering strategies to improve both tendon mid-substance healing and tendon integration into bone at the insertion site.

Previous studies performed by this laboratory have demonstrated that TEC stiffness at the time of surgery is positively correlated with repair tissue stiffness 12 weeks post-surgery. 81, 130

Our next step was to test the limits of this correlation by implanting a TEC that matched the stiffness of the normal tendon, a soft tissue PT autograft (PTA).

Aim 1. Determine the biomechanical and histological effects of implanting a soft tissue autograft in a central-third patellar tendon defect relative to natural healing, previous repairs using tissue engineered constructs and normal patellar tendon.

Hypothesis 1: The PTA will produce repair tissue with biomechanical properties superior to natural healing at 12 weeks post-surgery.

Hypothesis 2: The PTA will produce repair tissue with biomechanical properties superior to our best MSC-collagen sponge TEC repairs but inferior to the normal central-third PT.

Treatment with PTA improved repair tissue stiffness relative to natural healing (NH; p =

0.009). However, PTA and NH tissues did not differ in maximum force, modulus and maximum

stress. Therefore, our first hypothesis was rejected except for the case of stiffness. Additionally, while treatment with PTA increased insertion site failures by about 50% relative to NH, neither repair group regenerated a normal zonal insertion between the patellar tendon and bone. The experiments in this dissertation investigate two approaches to improve insertion site development of the patellar tendon autograft. The first approach allows for a longer recovery period (Aim 2) and the second implements biologic augmentations at the patellar and tibial insertion sites (Aims

3 and 4; Fig. 2).

The maximum force, linear stiffness and maximum stress values for both PTA repair and

NH were inferior to corresponding results for TEC repair (p ≤ 0.001) and normal central-third PT

(p < 0.001). The linear moduli of both PTA repair and NH were significantly lower than values for the normal central-third PT (p < 0.001) but these moduli did not differ from TEC repair.

Overall, the repair tissue biomechanics of our mechanically pre-conditioned MSC-collagen sponge TECs were approximately twice the biomechanical properties for both PTA repairs and

NH. Therefore, our second hypothesis was rejected. However, these results warrant additional experiments to further improve in vitro TEC stiffness and mRNA expression with the objective to enhance TEC-mediated healing (Aims 5 and 6; Fig. 2).

Figure 2. Experimental design flow chart. The two primary conclusions of Aim 1 each led to a distinct research pathway with unanswered research questions (RQ). The first pathway (left) focused on improving tendon-to-bone insertion site formation while the second pathway (right) focused on improving tissue engineered construct (TEC)-mediated repair. The primary rationale for exploring each research question is given by briefly stating what is “known”.

The results of Aim 1 are consistent with the findings of other investigators who show that tendon re-attachment to bone does not regenerate a normal zonal insertion site.115, 131-133 In a bone tunnel, evidence of tendon fibers inserting into the surrounding bone have been demonstrated at 12 weeks post-surgery.115 However, even at 26 weeks, a normal insertion site was not completely established.115 In a patellar tendon bone trough, the zonal insertion between the PT and patella was not completely regenerated by 24 weeks but the tidemark separating mineralized and non-mineralized fibrocartilage was present along with fibrochondrocytic cells.133 In both the bone tunnel and bone trough, dramatic tissue maturation occurred at the

osteotendinous junction between 6 and 24-26 weeks post-surgery. Therefore, it is possible that by allowing our rabbits up to 26 weeks for recovery (Aim 2), we could see a relatively mature insertion site with tendon fiber integration into bone.

While the biological processes of intra-articular graft healing are well studied134-137, a critical role for the native cell population has yet to be identified. It is also unclear whether the extra-articular PTA will heal according to the process described for intra-articular graft healing or the process described for tendon healing. A primary difference between the two is that tendon healing allows for contributions from the native cell population while intra-articular graft healing does not (the first phase of graft healing is tissue necrosis32, 135). By removing the native cell population from an autograft material we can study the cellular contributions to healing without the interference of an immune reaction typically triggered by an acellular allograft.

Aim 2. Determine the effect of recovery time (6, 12 and 26 weeks) and the role of the native cell population (cellular vs. de-cellularized) on insertion site formation and repair tissue biomechanics of the patellar tendon autograft.

Hypothesis 3: Osteotendinous integration and biomechanical properties of both de-cellularized

PTA (dcPTA) and PTA repairs will improve with increasing recovery time.

Hypothesis 4: Removing the native cell population will eliminate the capacity for intrinsic healing and thus produce inferior tendon-to-bone integration and repair tissue biomechanical properties for dcPTA repairs.

The objective of this study was to understand the role of the native cell population on graft incorporation. Osteotendinous integration improved with recovery time for both de- cellularized PTA (dcPTA) and PTA repairs. However, biomechanical properties were only affected by recovery time for dcPTA repairs (p < 0.05). Repair tissue modulus increased with time (p < 0.05) while the remaining biomechanical properties increased from 6 to 12 weeks (p <

0.05) but did not show continued improvement up to 26 weeks. Despite the changes in biomechanical properties demonstrated by dcPTA repairs, biomechanical properties did not vary between dcPTA and PTA repairs at any time point. Therefore, in this model of tendon healing, the native cell population offered no histological or biomechanical benefit. It is likely the PTA tissues went through a healing and incorporation processes mediated primarily by an extrinsic cellular component, similar to that described for intra-articular graft healing.

The results of Aim 1 and 2 both demonstrated that the soft tissue PTA did not regenerate a normal tendon-to-bone insertion site when implanted in a central-third PT defect model in the rabbit. While there are several potential explanations as to why the insertion site was not regenerated, Aims 3 and 4 are focused on the possibility that the biological cues necessary to promote bone ingrowth and tendon incorporation may have been absent or present at insufficient levels. We first postulated that placing a biologic augmentation composed of key components of an insertion site, namely collagen type I to span from tendon to bone and MSCs capable of differentiating into both chondrocytes and osteoblasts, would improve osteotendinous integration. MSC-collagen gel biologic augmentations (BAs) with a cellular density of 100K cells/ml were allowed to contract around two restraining posts in a silicone dish for 14 days. The

MSC-collagen gel BAs were then implanted in the patellar and tibial bone troughs posterior to the PTA.

Aim 3. Determine the effect of MSC-collagen gel biologic augmentations on repair tissue biomechanics and insertion site formation of PTA repair in vivo.

Hypothesis 5: MSC-collagen gel biologic augmentations will improve osteotendinous integration relative to un-augmented PTA repairs after 12 weeks of recovery.

Hypothesis 6: MSC-collagen gel biologic augmentations will improve repair tissue biomechanical properties at 12 weeks relative to un-augmented PTA repairs after 12 weeks of recovery.

At 12 weeks, there were no marked differences in insertion site formation between PTA repairs with BAs (PTA+BA) or un-augmented PTA repairs (PTA). Additionally, maximum force, linear stiffness, maximum stress and modulus were not statistically different for PTA or

PTA+BA repairs.

Multiple investigators have reported that the limiting factor in regenerating a normal, zonal insertion site is bone ingrowth into the fibrous scar that typically forms as part of the tendon healing response.115, 131, 132, 138, 139 It became apparent that the MSC-collagen gel BAs would likely need to be preconditioned toward an osteogenic lineage in vitro. The biologic activity of cells is largely determined by both cell-cell and cell-matrix interactions. Cell-cell interactions are regulated by cell density and it has been demonstrated that both cellular density and cellular shape modulate MSC phenotype.80, 81, 130, 140-143 When trying to promote osteogenic and chondrogenic differentiation in a collagen-glycosaminoglycan sponge, rat MSCs were inoculated at roughly 30,000 cells/mm2 of matrix.142 In contrast, when trying to promote tenogenic differentiation in a type I collagen sponge, rabbit MSCs were inoculated at roughly

800 cells/mm2 of matrix.81, 130, 140, 144-146 Additionally, when type I collagen gels created with concentrations of 1, 4 and 8 million cells/ml were implanted in rabbit PT defects, ectopic bone was found in 28% of samples.80 Therefore, we wanted to test both high and low cellular densities to gauge how MSCs would react in the collagen gel matrix. At the low end, we chose 100K cells/ml because this is near the lower threshold of contraction for MSCs inoculated in collagen gel (in our silicone dish culture system). While contraction varies by cell line, MSC-collagen gel

BAs created with 100K cells/ml show minimal contraction but maintain cellular viability. This was also the cellular density of MSC-collagen gel BAs used for the first iteration of in vivo implantation. At the high end, we chose 600K cells/ml because this was roughly the highest cellular density that could be maintained for two weeks in the silicone dish without pre-mature failure (i.e. the MSC-collagen gel contracting so much that it tore off the restraining post in the well of the silicone dish). Another regulator of cellular density is matrix contraction; as cells contract the collagen gel matrix around them, the cellular density inherently increases. To determine if cell-mediated contraction would influence cell-cell interactions, MSC-collagen gel

BAs were fabricated in two different cell culture systems, a silicone dish and around a taut suture in a glass boat, which each allow for different amounts of contraction. MSC-collagen gel BAs were also cultured for 7 and 14 days because time in culture can have a significant impact on cellular behavior.141, 147-149 The alkaline phosphatase activity (ALP activity; early marker for bone) of osteoblast-like cells seeded onto the bone phase of a triphasic scaffold peaked at day 7 and decreased thereafter.147 When MSCs were cultured into high density cellular aggregates, the matrix was composed primarily of collagen type I but, by days 5 and 7, collagen types II and X were present, respectively.141 Lastly, when MSCs-agarose TECs were cultured under cyclic compression, mRNA expression of collagen type II increased from day 7 to day 14 and

decreased thereafter.149 Significant changes in mRNA expression and cellular activity have been documented at both 7 and 14 days of culture.141, 147-149 Therefore, it is important to test both of these time points in order to determine when the MSC-collagen gel BAs might be best suited for insertion site formation in vivo.

We determined the effects of cellular density, TEC contraction, and time in culture on mRNA expression of osteo- and chondrogenic genes in vitro using real time quantitative reverse- transcriptase PCR (qRT-PCR). For the osteogenic pathway, we tested for the protein collagen type I and the transcription factor runx2. For the chondrogenic pathway, we tested for the protein collagen type II and the transcription factor sox9. Gene expression levels were normalized to the endogenous housekeeping gene GAPDH based on the results of a housekeeping panel which demonstrated GAPDH to be the most stable control of the six possibilities tested. We also evaluated ALP activity using a colorimetric kit (AnaSpec, San Jose, CA). ALP activity was normalized to cell number using a MTT (3-(4, 5-dimethylthiazol-2-yl)-2, 5-diphenyltetrazolium bromide) assay. When gene expression levels were normalized to those for MSC-collagen gel

BAs created with 100K cells/ml and cultured in a silicone dish for 14 days (the BA used for the first iteration of in vivo implantation), BAs created with 100K cells/ml and cultured in a glass boat for 14 days had significantly higher expression of collagen types I and II. Additionally, ALP activity was significantly affected by culture system, cellular density, and days in culture (p ≤

0.007). ALP activity per cell was higher for the glass boat culture system versus the silicone dish

(p < 0.001), higher for cellular density of 100K vs. 600K cells/ml and higher at day 7 versus day

14. We then chose the next BA based on each of these outcomes except for one, the days in culture. Although ALP activity was higher at day 7, we chose to implant BAs after 14 days based on the theory that if ALP activity peaked at day 7, perhaps the cells would be further committed

to the osteogenic lineage by day 14. Therefore, for the second iteration of surgery, we implanted

MSC-collagen gel BAs created with a cellular density of 100K cells/ml and cultured around a taut suture in a glass boat for 14 days. Based on the differences in gene expression levels between the two MSC-collagen gel BAs, the BA utilized in the first iteration of surgery with relatively lower gene expression levels will be referred to as low-activity (LA) and the BA with relatively higher expression levels will be referred to as high-activity (HA).

Aim 4. Determine the effect of MSC-collagen gel BAs with enhanced osteogenic capacity on repair tissue biomechanics and insertion site formation for PTA repair in vivo. Contrast the histological and biomechanical results obtained for BAs with relatively high activity (BA-HA) against those for BAs with relatively low activity (BA-LA).

Hypothesis 7: After 12 weeks of recovery, osteotendinous integration of PTA treated with BA-

HA will be superior to the integration observed for PTA alone (un-augmented repairs).

Hypothesis 8: After 12 weeks of recovery, repair tissue biomechanical properties of PTA treated with BA-HA will be superior to those for PTA alone (un-augmented repairs).

Hypothesis 9: After 12 weeks of recovery, PTA+BA-HA repairs will have osteotendinous integration superior to PTA+BA-LA repairs.

Hypothesis 10: After 12 weeks of recovery, PTA+BA-HA repairs will have repair tissue biomechanical properties superior to PTA+BA-LA repairs.

Treatment with MSC-collagen gel BA-HAs did not improve osteotendinous integration relative to PTA repair at 12 weeks or increase PTA repair tissue biomechanical properties at 12 weeks post-surgery. The results of our biomechanical evaluation of PTA+BA and PTA samples at 12 weeks post-surgery agree with our previous finding that MSC-collagen gel BAs fabricated with 100K cells/ml in a silicone dish for 14 days (PTA+BA-LA) did not improve PTA repair tissue biomechanical properties (Aim 3). When treatments were contrasted against each other,

PTA+BA-HA repairs appeared better integrated at the insertion sites histologically versus

PTA+BA-LA repairs. However, this did not translate into superior biomechanical properties.

Based on the biomechanical and histological data of Aims 3 and 4, it is possible that BAs at the tendon-bone interface limit rather than promote osteotendinous integration. This may be due to a variety of factors including: 1) the BA reduces the amount of pressure between the PTA and bone defect, thus inhibiting integration of the tissues, 2) the BA does not possess the correct biological cues to promote integration in the environment of the interface, and 3) the BAs did not remain in place.

As mentioned in Aim 1, after 12 weeks of recovery, the repair tissue biomechanics of our mechanically pre-conditioned MSC-collagen sponge TECs were approximately twice the biomechanical properties for both PTA repair and NH. These results warranted additional experiments to further improve in vitro TEC stiffness and mRNA expression with the objective to enhance TEC-mediated healing (Aims 5 and 6).

Our goal is for tendon repair biomechanics to match normal tendon tangent stiffness up to

40% of failure force to accommodate moderate activities of daily living.82-85 Although MSC- collagen sponge TEC repair is superior to PTA repair and NH healing 12 weeks after surgery, current repairs have only matched normal patellar tendon up to 32% of failure force.81 TEC

stiffness at the time of surgery is positively correlated with repair tissue stiffness 12 weeks post- surgery.81, 130 Therefore, the objective of this research on MSC-collagen sponge TECs was to increase in vitro stiffness and thereby improve repair tissue biomechanical properties.

The biomechanical properties of TECs may be influenced by both cell-matrix interactions and in vitro culture conditions (i.e. mechanical pre-conditioning with uniaxial strain). One strategy to alter cell-matrix interactions and potentially improve MSC-collagen sponge TEC linear stiffness is to incorporate the glycosaminoglycan (GAG) chondroitin-6-sulfate (C6S) into the collagen scaffold.150-152 Although C6S is not the predominant GAG in tensile-load bearing tendons57, it does bind to decorin which is essential for proper collagen fibrillogenesis in tendon45. Scaffold materials composed of collagen and C6S have been used to facilitate regeneration of the dermis153 and sciatic nerve154 but hold the potential to support MSC differentiation down multiple pathways142. Although the role of the GAG component is not completely understood, studies have demonstrated that the presence of GAG delays wound contraction152, 155 and others have suggested the GAG may act as a “linker” molecule between collagen fibrils creating a more homogenous architecture within the sponge.57, 156 Additionally, mechanical stimulation/pre-conditioning with uniaxial strain has not only significantly improved

TEC linear stiffness but it has also up-regulated mRNA expression of collagen types I and III

(decorin and fibronectin levels were not affected by mechanical stimulation).81, 130

Aim 5. Determine the individual and combined effects of chondroitin-6-sulfate incorporation and mechanical stimulation on MSC-collagen sponge TEC biomechanics and mRNA expression of collagen types I and III, decorin and fibronectin.

Hypothesis 11: Under static culture conditions, incorporating C6S will increase MSC-collagen sponge TEC stiffness.

Hypothesis 12: Under static culture conditions, incorporating C6S will increase MSC-collagen sponge TEC mRNA expression for all genes of interest.

Hypothesis 13: Pre-conditioning with uniaxial mechanical strain will increase MSC-collagen sponge TEC stiffness for TECs created using both a collagen sponge (COL) and a collagen-C6S sponge (COL-C6S).

Hypothesis 14: Pre-conditioning with uniaxial mechanical strain will increase MSC-collagen sponge TEC gene expression levels of all genes of interest for TECs created using both a collagen sponge (COL) and a collagen-C6S sponge (COL-C6S).

Hypothesis 15: Mechanical stimulation and C6S incorporation will have a synergistic effect on

MSC-collagen sponge TEC linear stiffness in vitro.

The incorporation of C6S did not improve MSC-collagen sponge TEC biomechanical properties but it did increase expression levels of collagen type I and decorin.157 For COL and

COL-C6S constructs, mechanical stimulation did not increase TEC stiffness above static controls.157 However, mechanical stimulation with 3000 cycles/day did increase collagen type I expression for both COL and COL-C6S constructs.157 Although the COL and COL-C6S constructs did not respond biomechanically to mechanical stimulation, the incorporation of C6S

and mechanical stimulation played a significant role in the gene expression profiles of MSC- collagen sponge TECs.157 We will continue to explore the effect of C6S incorporation and mechanical stimulation on MSC-collagen sponge TEC biomechanics and biochemistry to expand our knowledge of in vitro predictors of in vivo outcomes.

Despite the positive results of this study, the mRNA expression of COL and COL-C6S constructs were orders of magnitude lower than results produced in a previous study using a commercially derived collagen sponge scaffold (CD-COL).140 In contrast, the linear stiffness of non-stimulated CD-COL TECs was roughly half of that for COL and COL-C6S constructs.81 To investigate the discrepancies in gene expression and biomechanics, we evaluated the architecture of each material. Scanning electron microscopy revealed that the average pore diameter within the CD-COL sponge was approximately three times that of the COL and COL-C6S scaffolds.157

These results suggest that in addition to C6S incorporation and mechanical stimulation, mean scaffold pore size may help modulate TEC biomechanical and biochemical properties.

Scaffold architecture is important for cell-matrix interactions because it has been proposed that within the collagen sponge, ligand binding density decreases with increasing pore diameter.158 However, while the small pore creates an attractive surface for cell attachment, pores cannot be made so small that cellular infiltration into the scaffold is inhibited. For mammalian dermal regeneration, optimal scaffold pore diameter was between 20 and 125µm.155

For rat nerve regeneration, collagen-C6S scaffolds were fabricated with a mean pore diameter of

35um.154 Finally, for mouse MC3T3-E1 cells, the fraction of attached cells was highest when mean pore diameter was 95.9um and decreased with increasing pore size.158 These studies highlight the importance of optimizing pore size diameter for specific cell types and specific tissue engineering applications.

Primarily type I collagen sponges, with and without the inclusion of C6S (COL-C6S and

COL, respectively), were fabricated with three distinct average pore diameters: small, medium, and large. These pore sizes were chosen such that each scaffold material was unique with minimal overlap in pore diameter between scaffold materials (Fig. 3).

A B C

Figure 3. SEM images of small (A), medium (B), and large (C) pore size scaffold materials. The average pore diameter for each material was: small (40.4 ± 17.0μm; mean ± SD), medium (106.6 ± 30.4μm) and large (180.1 ± 45.8μm). Scale bar = 200µm.

Aim 6. Determine the individual and combined effects of scaffold pore size, chondroitin-6- sulfate incorporation and mechanical stimulation on MSC-collagen sponge TEC biomechanics and mRNA expression.

Note: The data set for this experiment is not complete and therefore not included in this thesis beyond the statement of the aim and the hypotheses below.

Hypothesis 16: Under static culture conditions, pore size diameter of both COL and COL-C6S scaffold materials will significantly affect MSC-collagen sponge TEC stiffness.

Hypothesis 17: Under static culture conditions, pore size diameter of both COL and COL-C6S scaffold materials will significantly affect mRNA expression levels of all genes of interest for

MSC-collagen sponge TECs.

Hypothesis 18: Pre-conditioning with uniaxial mechanical strain will increase MSC-collagen sponge TEC stiffness for TECs created using small, medium and large pore diameter scaffolds.

Hypothesis 19: Pre-conditioning with uniaxial mechanical strain will increase MSC-collagen sponge TEC gene expression of all genes of interest for TECs created using small, medium and large pore diameter scaffolds.

Hypothesis 20: When pre-conditioned with mechanical stimulation, pore size diameter will have a significant effect on both MSC-collagen sponge TEC linear stiffness and mRNA gene expression levels for all genes of interest.

Chapter 3

Effect of Implanting a Soft Tissue Autograft in a Central-Third Patellar Tendon Defect: Biomechanical and Histological Comparisons* Kirsten R.C. Kinneberg1, Marc T. Galloway2, David L. Butler1, Jason T. Shearn1

1. School of Energy, Environmental, Biological and Medical Engineering, Biomedical Engineering Program, University of Cincinnati, Cincinnati, OH; 2. Cincinnati Sportsmedicine, Cincinnati, OH * Manuscript is in preparation and will be submitted to the Journal of Biomechanical Engineering.

Abstract Previous studies by our laboratory have demonstrated that implanting a stiffer tissue engineered construct at surgery is positively correlated with repair tissue stiffness at 12 weeks.

The objective of this study was to test this correlation by implanting a construct that matches normal tissue biomechanical properties. To do this, we utilized a soft tissue patellar tendon autograft (PTA) to repair central-third patellar tendon (PT) defects. Bilateral defects were established in the central-third PT of New Zealand White rabbits. In one limb, the excised tissue was sutured into the defect site (PTA). In the contralateral limb the defect was left empty (natural healing, NH). We hypothesized that after 12 weeks, PTA repairs would have biomechanical properties superior to NH. We were surprised to find that only stiffness was improved by treatment with PTA relative to NH (p = 0.009). Additionally, neither the PTA nor NH repairs regenerated a normal zonal insertion site between the tendon and bone. Future studies by our laboratory will investigate potential strategies to improve PTA integration into bone using this model.

Introduction Tendon injury represents a significant clinical challenge in orthopaedic and sports medicine. While treatments such as direct repair and auto- or allograft replacement are viable options, the occurrence of re-rupture and post-operative complications is often high.94, 103, 138 As

injury rates rise and the number of primary reconstructions increases, the need for revision surgery is likely to increase as well.159 These issues present a clear need to improve traditional tendon repair strategies.

Currently, multiple strategies are being studied to enhance tendon repair. One strategy is treatment with growth factors.132, 138, 139, 160 Despite the positive result of increasing repair tissue structural properties relative to untreated and/or sham controls132, 139, 160, growth factor treatments appear to promote large amounts of poor quality scar tissue rather than regenerated tendon.138, 139

Another approach is to augment repairs by reinforcing sutures with various biologic scaffolds

(i.e. CuffPatch (Arthrotek, Warsaw, IN), Restore (Depuy, Warsaw, IN), etc).161, 162, 224

Augmentation grafts increase suture fixation strength compared to un-augmented repairs162 and the biochemical composition of the grafts is similar to that of tendon. However, the discrepancy between elastic moduli of grafts and native tendon limits their mechanical role in augmenting tendon repair.161 To offer both a biological and mechanical component to tendon repair strategies, researchers are developing tissue engineered constructs (TECs).

Our laboratory has focused on TECs composed of collagen-based scaffolds and mesenchymal stem cells (MSCs) to repair tendon defects.80, 81, 130, 140, 144, 163, 164 When implanted into a central-third patellar tendon (PT) defect in the rabbit, mechanically stimulated MSC- collagen sponge TECs produced 12-week repairs that matched the tangent stiffness of normal patellar tendon up to 32% of failure force or 50% greater than the loads and displacements required for normal activities of daily living.81 Additionally, paired in vitro-in vivo studies found that in vitro TEC stiffness was significantly and positively correlated with repair tissue stiffness

12 weeks after surgery.81, 130 Our results suggest that repairing a central-third PT defect with a stiffer implant will promote a better repair outcome.

To examine if further increases in implant stiffness would continue to enhance tendon repair, we filled the central-third PT defect with a soft tissue PTA. We hypothesized that, after

12 weeks of recovery, the PTA would produce repair tissue with biomechanics superior to NH.

Additionally, when compared to our previous studies80, 81, we hypothesized that the PTA repair tissue biomechanical properties would be superior to repairs using TECs but inferior to the normal central-third PT.

Materials and Methods All procedures were approved by the University of Cincinnati Institutional Animal Care and Use Committee. At surgery, each rabbit was anesthetized with an anesthesia cocktail of ketamine/acepromazine (40mg/kg and 5mg/kg, respectively) and isoflurane gas (as needed).

Under aseptic conditions, each patellar tendon was exposed through an anteromedial incision. A

3 mm wide, full-length, full-thickness, central-third defect was created in each PT, producing a soft tissue, patellar tendon autograft (PTA). Bone defects were created at the proximal and distal insertion sites using a pneumatic sagittal saw (MicroAire Surgical Instruments, Charlottesville,

VA). In the right limb, the excised PTA was sutured back into the defect at four sites (Prolene

50, Ethicon, Somerville, NJ): two each near the patella and tibia. In the contralateral limb, the defect remained unfilled (natural healing, NH). Each incision was closed with a continuous subcutaneous suture followed by simple interrupted sutures in the skin. After recovering from anesthesia, each animal was returned to its cage and allowed unrestricted cage activity. Twelve weeks post-surgery, each rabbit was anesthetized as described above and then euthanized by intracardiac administration of Euthasol® (1.0cc/4.54kg; Virbac Animal Health, Inc., Fort Worth,

TX) followed by bilateral pneumothorax. Hind limbs were disarticulated at the hip and removed using a scalpel. For biomechanical evaluation, whole limbs were stored at -20°C for

approximately two weeks until testing. For histological evaluation, patella-PT-tibia samples were placed in 10% neutral buffered formalin and stored at room temperature (25°C) until further processing.

Biomechanical Evaluation One day prior to testing, limbs were removed from the freezer and allowed to thaw over night. On the day of testing, all extraneous tissues were removed leaving the patella, patellar tendon (PT) and tibia. The length and width of the whole PT were measured at the medial/central/lateral and proximal/central/distal portions of the tendon, respectively, using

Vernier calipers (accurate to 0.1mm). Tendon length was measured on the posterior surface from tibial insertion to patellar insertion. Due to the ellipsoidal shape of the PT, whole PT width and thickness are reported by region (proximal/central/distal). Thickness of the whole PT was measured at the proximal/central/distal portions of the tendon using a light force (<0.15 N) digital micrometer (accurate to 0.01mm; IDCtype Mitutoyo Digimatic Indicator, MTI Corp.,

Aurora, IL). The native tendon struts were removed to isolate the central-third repair tissue. The tibia/fibula complex was cut roughly 2.5cm distal to the tibial tuberosity to create a bone block.

The length, width and thickness of the patella-central-third repair-tibia samples were then measured as described for the whole PT. Patellar and tibial bone blocks were fixed into custom designed grips using polymethylmethacrylate cement (Dentsply International , York, PA).81, 164

The testing grips for each sample were secured into a Plexiglas tank mounted on a materials testing system (Model 8501, Instron, Inc., Canton, MA). The tank contained phosphate buffered saline (PBS; pH 7.4) heated to 37°C. Repair tissues were preloaded to 14.8 ± 8.8 N (mean ± SD) and then preconditioned for 50 cycles to 3% strain at 1Hz using a sinusoidal strain pattern.

Tissues were then failed in uniaxial tension at a constant strain rate of 20%/sec while force and displacement were continuously recorded.78, 81, 144, 164

Histological Evaluation Tissue processing has been described previously165 but a brief protocol is outlined below.

Fixed patella-PT-tibia samples were sectioned in the transverse plane to isolate a) the patella and proximal PT, b) the tendon mid-substance, and c) the tibia and distal PT. Patellar and tibial bone ends were decalcified in 10% formic acid. Each tissue section was then dehydrated through a gradient of alcohols and xylene before being embedded in a paraffin block. The bony ends

(patella and proximal PT, tibia and distal PT) were processed by cutting eight serial sections

(4µm thick) in the sagittal plane at 1 mm intervals through the sample. These tissues were used to examine tendon integration into bone and insertion site formation. The mid-substance samples were processed by cutting eight serial sections in the coronal plane at 250µm intervals. These tissues were used to examine PTA and NH integration into the tendon native struts. One section from each depth was stained with hematoxylin and eosin (H&E). Select serial sections were then subjected to immunohistochemical (IHC) staining for collagen type II (bone ends only;

Calbiochem, San Diego, CA) and collagen type III (Sigma-Aldrich, St. Louis, MO). Patellar and tibial bone ends were stained with antibodies for collagen type II because it is the primary structural protein of the fibrocartilage which composes the tendon-to-bone insertion site. Mid- substance samples were stained with antibodies for collagen type III because it is an important structural protein during healing. Repair tissue organization, cellularity, neovascularization, repair tissue integration into the native tendon struts, and tendon-bone insertion site formation were evaluated by one of the authors (KRCK).

Statistical Analysis One animal sustained a unilateral rupture of its left limb (natural healing treatment group). In a separate animal, the biomechanical properties were statistical outliers for the right limb (PTA treatment group). These two limbs were excluded from statistical analysis leaving a sample size of n = 6 per treatment group. Data were normal and homoscedastic. Differences in dimensional (n = 6) and biomechanical (n = 6) data were assessed using a one-way ANOVA with repair treatment as a fixed factor. The significance level was set at p < 0.05.

Results

Repair Tissue Dimensions After 12 weeks of healing, both whole PT and central-third repair tissue dimensions were significantly affected by repair treatment (NH vs. PTA; Table 1). Compared to NH, treatment with the PTA significantly increased: 1) whole PT thickness by 35%, 30% and 36% in the proximal, central, and distal regions, respectively (p ≤ 0.044); 2) whole PT cross-sectional area by 46% in the distal region (p = 0.025); and 3) central-third repair tissue width, thickness, and cross-sectional area by 20%, 47%, and 78%, respectively (p = 0.011, 0.004 and 0.001, respectively).

Table 1. Repair Tissue Dimensions [mean (SEM)] for Whole and Central-Third PT Whole PT Central-Third PT Length Width Thickness Length Width Thickness (mm) (mm) (mm) (mm) (mm) (mm) Proximal Central Distal Proximal Central Distal NH 18.7 9.6 10.4 10.4 2.0 2.0 2.2 19.2 2.5 1.5 (n=6) (0.5) (0.2) (0.7) (0.7) (0.2) (0.2) (0.2) (0.4) (0.2) (0.1) PTA 20.7 9.9 11.5 11.8 2.7* 2.6* 3.0* 21.4 3.0* 2.2* (n=6) (0.9) (0.7) (0.5) (0.8) (0.2) (0.2) (0.2) (1.0) (0.1) (0.1) * Significantly greater than NH repair values (p < 0.05)

Repair Tissue Biomechanical Properties and Failure Mechanisms Treatment with PTA significantly enhanced repair tissue stiffness, as compared to NH, at

12 weeks post-surgery (p = 0.006; Fig. 4, Table 2). However, maximum force, maximum stress and modulus were not affected by repair treatment. Of the six PTA repairs tested, five failed at the patellar insertion site and one failed in the mid-substance. By contrast, of the six NH repairs tested, three failed at the patellar insertion site and three failed in the mid-substance. All samples failed by soft tissue pull-out from the insertion site with no detectable bone avulsion for any of the samples.

Figure 4. Force-displacement curves (mean ± SEM). PTA (n=6) and NH (n=6) repairs both exceeded the peak in vivo force required for activities of daily living (100N; in vivo force and displacement, IVF and IVD, respectively)83, 84. However, PTA and NH repairs do not match the normal central-third PT (Normal; n = 880, 81) or TEC (n = 781) repair curves. Portions of this figure re-printed with permission from Juncosa-Melvin, et al. (2006)81.

Table 2. Biomechanical Properties (Mean ± SEM) of Natural Healing, Patellar Tendon Autograft, Tissue Engineered Construct and Normal Central-Third PT NH PTA TEC1 Normal2

(n = 7) (n = 7) (n = 7) (n = 8) Max Force (N) 121.8 ± 13.0 b,c 154.2 ± 20.1 b,c 339.3 ± 10.9 c 470.7 ± 23.8 Stiffness (N/mm) 52.0 ± 4.3 a,b,c 74.2 ± 4.6 b,c 141.6 ± 3.2 159.8 ± 11.1 Max Stress (MPa) 34.9 ± 4.9 b,c 23.7 ± 3.3 b,c 72.0 ± 1.8 c 100.7 ± 5.6 Modulus (MPa) 286.0 ± 36.8 b,c 243.4 ± 17.1 b,c 441.1 ± 3.1 c 861.4 ± 98.5 1TEC repair tissue biomechanical properties were previously obtained81 2Normal central-third PT biomechanical properties were previously obtained80, 81 a. Significantly less than PTA repair (p < 0.05) b. Significantly less than TEC repair (p ≤ 0.05) c. Significantly less than normal central-third (p ≤ 0.05)

Tendon Repair Mid-Substance Histology PTA and NH repair mid-substance tissues showed several similarities (Fig. 5). First, PTA and NH central-third repair tissues contained regions of neo-vascularization and hypercellularity, as did the native struts. Second, PTA and NH samples contained regions with aligned tissue showing some evidence of a crimp pattern along with both rounded and elongated cell nuclei.

Lastly, a band of hypercellular tissue was present at both the graft- and natural healing-native strut interfaces for PTA and NH samples, respectively.

Histological examination also revealed differences between PTA and NH samples. While both central-third repairs appeared hypercellular relative to normal, NH repairs appeared more hypercellular than PTA repairs (Fig. 5). Additionally, collagen type III staining was primarily localized to the graft-native tissue interface in PTA repair samples (Fig. 5D) but was found throughout the repair tissue in NH samples (Fig. 5B).

Patellar and Tibial Insertion Site Histology Although PTA and NH repairs showed similar results at the patellar and tibial insertions, neither repair matched the zonal insertion appearance of the native struts. In the native strut, the tendon passed through visible layers of fibrocartilage (FC) and mineralized fibrocartilage (MFC) before anchoring into the underlying bone (Fig. 6, bottom row). By contrast, neither the PTA nor

NH repair tissue regenerated a normal zonal insertion site (distinct regions of FC and MFC were not evident) at the patella or tibia. Instead, a fibrous scar (FS) tissue was evident at the insertion sites and the fibers in the PTA and NH repairs generally remained either parallel to or oriented away from the bone surface (Fig. 6, middle row and top row, respectively). H&E staining also revealed a tidemark in the native struts separating regions of fibrocartilage and mineralized fibrocartilage (Fig. 6, I and K) that was not present in PTA (Fig. 6, E and G) nor NH samples

(Fig. 6, A and C). IHC staining for collagen type II showed that both the PTA (Fig. 6, F and H) and NH (Fig. 6, B and D) tissues contained areas of fibrocartilage-like tissue similar to that found in the native struts (Fig. 6, J and L). However, the fibrocartilage region of the native struts contained rows of chondrocytes separated by collagen fibers anchoring into the underlying bone

(Fig. 6, bottom row) while the fibrocartilage-like regions of the PTA and NH samples contained disorganized chondrocytes with no evidence of fibers actually passing through the insertion site and anchoring into bone (Fig. 6, middle row and top row, respectively).

Figure 5 (above). H&E staining and IHC staining for collagen type III (Col3) of the tendon mid- substance for both NH and PTA after 12 weeks of healing. Col3 (brown) is primarily localized to the central-third repairs tissue (R) for natural healing (NH) samples and localized to the graft- native strut (NS) interface for PTA samples. Neo-vascularization (arrows) is visible in the native struts (NS) and central-third repair (R) regions for both tissues (the central-third repair is labeled as R for both NH and PTA samples). Scale bar = 200um.

Figure 6 (next page). Histological images of patellar and tibial tendon-to-bone insertion sites. Insertion site images are shown with H&E staining and IHC staining for type II collagen (Col2). After 12 weeks of healing, proper insertion sites were not regenerated by PTA or NH. The insertion sites in the native struts (NS) appear unaffected by the surgical procedure. H&E images are labeled to reflect: B, bone; T, tendon; FS, fibrous scar; MFC/FC, mineralized fibrocartilage/fibrocartilage zones of the insertion site, respectively (note: these zones are not as distinct in the patella as they are in the tibia); areas resembling neo-vasculature are circled. Scale bar = 200µm.

T T T

FS B

MFC/FC

B T

FS MFC/FC

Discussion Our first hypothesis, that after 12 weeks of recovery, the PTA would produce repair tissue with biomechanical properties superior to NH was only correct for stiffness. The statistical equivalence of maximum force, maximum stress, and modulus for NH and PTA repairs is likely due to poor tendon-to-bone integration at the patellar and tibial insertions. Relative to NH, treatment with the PTA increased insertion site failure frequency from ~50% to ~83%.

Therefore, it is possible that the PTA repair tissue mid-substance was biomechanically superior to that for NH but, because normal insertion sites were not regenerated, the PTA repairs failed at their weak link before the mid-substance tissue could be loaded to failure to determine its biomechanical properties.

Although our histological findings were discouraging, the results of this study are consistent with those of other investigators who found that at 12 weeks: 1) tendon re-attachment to bone did not regenerate a normal insertion site115, 132 and 2) samples failed at the soft tissue-to- bone attachment site, similar to what we presented for PTA repairs132. It was not until 26 weeks that 1) a tidemark became evident and tendon fibers were found anchoring into the surrounding bone for the length of the bone tunnel and 2) samples failed by tendon pull-out from the clamp or by mid-substance rupture.115

There are several possible reasons for why the NH and PTA repairs failed to regenerate normal insertion sites. 1) The rabbits were only given 12 weeks to recover and, as mentioned above, animals may need a longer recovery period to allow soft tissue-to-bone healing to mature.115 Our future studies will examine the effects of longer recovery times on both repair tissue biomechanical properties and insertion site formation. 2) The tendon-to-bone insertion site may not be exposed to the optimal mechanical environment for proper healing. While early protected passive mobilization can enhance the strength of repaired tendons166, early

immobilization may provide a protective environment for tendon integration into bone that drives the healing response toward regeneration rather than scar tissue formation.106, 115, 139 3)

The biological cues necessary to promote bone ingrowth and tendon incorporation may have been absent or present at insufficient levels.139 These biological cues include, but are not limited to, bone morphogenic proteins132, 139, collagen types I and II106, and alkaline phosphatase activity106. Future studies will examine how incorporating chemically and/or mechanically pre- conditioned biological augmentations at the insertion sites might enhance tendon integration into bone.

Our second hypothesis, that the PTA repair tissue biomechanical properties would be superior to our previous repairs using TECs but inferior to the normal central-third patellar tendon was rejected. While we expected the average force-displacement curve for NH to be significantly lower than our best TEC repair81, we were surprised to find that the average PTA force-displacement curve was also significantly less than our best TEC repair81(Fig. 4; Table 2).

PTA and NH repairs both exceeded the peak in vivo force required for activities of daily living

(100N)83, 84, which is one of our criteria for a successful repair. However, neither PTA repair nor

NH matched the tangent stiffness of the normal PT in the functional range of loading (Fig, 4, lower left corner), which is another one of our criteria for success. Overall, the repair tissue biomechanics of our mechanically pre-conditioned MSC-collagen sponge TECs were approximately twice the biomechanical properties of both PTA repairs and NH (Table 2). The biomechanical data for PTA repair go against the in vitro-to-in vivo predictor, that increased implant stiffness leads to increased repair tissue stiffness, found for TEC-based repairs81, 130. This leads us to propose that there may be a threshold of implant stiffness where the in vitro-to-in vivo correlation holds true but, beyond that level, implant stiffness inhibits integration with the

surrounding tissues. One potential explanation for this discrepancy is that the implanted tissue

(PTA or TEC) needs to match the biomechanical properties of the native struts before it can be integrated with them. It might seem as if the PTA and NS have the same initial biomechanical properties, however, when the central-third PT is removed, the native struts are overloaded and stress enhancement can reduce biomechanical properties137, 167-169. So, in the case of a PTA versus a MSC-collagen sponge TEC, the initial phase of healing for the PTA may involve catabolic mechanisms to reduce the implant biomechanical properties while that for the TEC may involve anabolic mechanisms to increase implant biomechanical properties. However, as discussed below, linear stiffness is not the only factor distinguishing the MSC-collagen sponge

TEC from the PTA.

Three factors that may explain why the TEC produced superior repair tissue biomechanics relative to the PTA are the cell population, the porosity and the compliance of the respective implants. The TEC contains bone marrow-derived MSCs that have been mechanically pre-conditioned in culture. We have demonstrated that, when compared to static culture, mechanical stimulation of MSC-collagen sponge TECs increases mRNA expression levels of collagen types I and III140, which each play an important role in tendon healing. The resident cell population of the PTA, on the other hand, is likely composed primarily of mature tenocytes with lower metabolic activity.170 While it is possible that the tenocytes may revert back to their highly metabolic state as tenoblasts, it is also possible that the graft tissue undergoes a phase of necrosis and hypocellularity, similar to what is seen during ligamentization of grafts for ACL replacement.32, 94, 135 Additionally, the porous microstructure of the MSC-collagen sponge TEC may offer the advantage of rapid cellular infiltration and easier integration with the native struts.

It is also possible that without early integration at the graft-native strut interfaces, the PTA was

stress shielded by the native struts and thereby not exposed to sufficient forces to promote significant tissue remodeling.171-173 Additionally, because the TEC is more compliant than the

PTA, it may take less force to stretch the implant and overcome the detrimental effects immobilization can have on tendon tissues.170, 174

Our study is not without limitations. 1) Our model does not replicate graft replacement or surgical reconstruction of a tendon or ligament tear. Our model does, however, provide a reproducible and accessible system to test tissue engineering strategies to promote tendon healing and integration into bone. 2) The current sample size did not allow us to make a definitive statistical conclusion about the maximum force sustained by PTA repairs and NH samples at 12 weeks. However, both repair methodologies are statistically inferior to TEC-based repairs and vast improvements are needed to produce a viable repair using the PTA.

This study demonstrates that a soft tissue PTA generates repair tissue that is generally equivalent to NH but inferior to both TEC repair and normal tendon. Biomechanical results show

PTA repairs possess inadequate tissue stiffness in the functional range of loading and fail near the insertion into bone. Histological results demonstrate incomplete integration of the PTA soft tissue into bone and affirm that the tendon-to-bone insertion is the weak link of the repair. Using the model presented here, our laboratory can now investigate potential strategies to improve the integration of tendon into bone. Future studies will examine novel ways to improve tendon-to- bone insertional repair, including MSC-gel TECs that have previously produced ectopic bones spicules in vivo.80

Chapter 4

Effect of Recovery Time and the Role of the Native Cell Population on Insertion Site Formation and Repair Tissue Biomechanics of the Patellar Tendon Autograft* Kirsten R.C. Kinneberg1, Marc T. Galloway2, David L. Butler1, Jason T. Shearn1

1. School of Energy, Environmental, Biological and Medical Engineering, Biomedical Engineering Program, University of Cincinnati, Cincinnati, OH; 2. Cincinnati Sportsmedicine, Cincinnati, OH *Manuscript is in preparation and will be submitted for publication in the journal Tissue Engineering.

Grafting procedures have a pivotal function in orthopaedic and sports medicine.

Autografts and allografts are commonly used to repair orthopaedic injuries when the damage is either too great for direct surgical repair or the tissue has poor inherent healing capacity.

Tendons, ligaments, menisci and cartilage are all commonly replaced with grafts. However, orthopaedic allografts are most commonly used for ACL reconstruction.107 Typical graft replacements for the ACL include the bone-patellar tendon-bone (BPTB) complex and the semitendinosus tendon of the hamstrings (HS).

All graft replacements, whether auto- or allogeneic, undergo a similar process of healing and incorporation when used for ACL repair.136 There are two primary sites of healing and incorporation: the tendon-to-bone insertion sites and the tendon body (also known as the tendon mid-substance)135, 136. Healing at the insertion sites occurs in the form of tissue integration between the tendon graft and bone. Integration of the graft material in the femoral and tibial bone tunnels is essential because, if a firm attachment is not obtained at the insertion sites, the graft cannot be exposed to sufficient mechanical loads to ensure the long-term survival and functionality of the graft. While the bone-to-bone healing that takes place with BPTB grafts occurs relatively quickly, the soft tissue-to-bone healing required for a HS graft can take up to 12

weeks to show signs of maturation and may never regain the transitional zones characteristic of a direct insertion115. Graft healing in the tendon mid-substance transforms the implant into a structure resembling the native ACL. This process enables the graft to better function as a ligament in the joint space.135

Despite the similar healing process for both graft types, allografts generally incorporate slower than autografts.94, 135-137 While it may seem like allografts are slower to incorporate because they are de-cellularized and lack a native cell population, the first phase of graft healing involves tissue necrosis and subsequent tissue repopulation with inflammatory cells, fibroblast- like cells and vascular proliferating cells. One explanation for the lag in allograft healing is an increased and prolonged immune response within the graft135, 136. In order to determine the effect of the native cell population, the foreign matrix of the allograft needs to be removed from the repair scenario.

The objective of this study was to understand the role of the native cell population on graft incorporation. To do this, we utilized our model of patellar tendon autograft healing and tested both a patellar tendon autograft (PTA) and a de-cellularized PTA (dcPTA). Additionally, because the model involves soft tissue-to-bone healing, which typically requires a long time period to mature, we also investigated the affect of recovery time on insertion site formation and repair tissue biomechanical properties. We hypothesized that: 1) removing the native cell population would allow for sooner infiltration of extrinsic cells and lead to more rapid osteotendinous integration of dcPTA repairs relative to PTA repairs, and 2) osteotendinous integration and biomechanical properties of both dcPTA and PTA repairs would improve with recovery time.

Materials and Methods All procedures were approved by the University of Cincinnati Institutional Animal Care and Use Committee. Twenty-seven one year old, skeletally mature, female New Zealand white rabbits were used for this study. At surgery, each rabbit was anesthetized, as previously described165, and both knees were aseptically prepared. The patellar tendon (PT) was exposed through an anteromedial incision. A central-third defect was created by excising a full-length, full-thickness segment measuring 3mm in width from the central portion of the PT. The excised tissue would serve as the patellar tendon autograft. PTA samples from the right limb were stored in phosphate buffered saline (PBS) soaked gauze (PTA) while those from the left limb were de- cellularized by three consecutive freeze-thaw cycles in liquid nitrogen (1 minute) and sterile PBS

(3 minutes) (dcPTA). Freeze-thaw cycling has been shown to kill 95-100% of cells.175 While the dcPTA was prepared, bone defects were created at the proximal and distal insertion sites

(approximately 3mm wide, 5 mm long and deep enough to induce bleeding from the cortical bone) using a pneumatic sagittal saw (MicroAire Surgical Instruments, Charlottesville, VA). In the right limb, the excised PTA was secured into the defect site with four sutures (Prolene 50,

Ethicon, Somerville, NJ): two each near the patella and tibia. The dcPTA was sutured into the left limb defect site using the same method. The sutures served to connect the PTA tissues with the adjacent native struts (NS), not to either bone end. Incisions were closed with a continuous subcutaneous suture followed by simple interrupted sutures in the skin. After recovering from anesthesia, animals were returned to their individual cages and allowed unrestricted cage activity. At six, 12, and 26 weeks post-surgery, rabbits were anesthetized (n = 9 per time point) as described above and then euthanized by intracardiac administration of Euthasol®

(1.0cc/4.54kg; Virbac Animal Health, Inc., Fort Worth, TX) followed by bilateral pneumothorax.

Hind limbs were disarticulated at the hip and removed using a scalpel. For biomechanical

evaluation, whole limbs were stored at -20°C until testing. For histological evaluation, patella-

PT-tibia samples were placed in 10% neutral buffered formalin and stored at room temperature

(25°C) until further processing.

Biomechanical Evaluation Limbs were thawed overnight at 4°C before testing. On the day of testing, limbs were dissected to isolate the patella, patellar tendon (PT) and tibia. Whole PT and repair tissue dimensions were obtained, as described in Chapter 3. Briefly, the length and width of both the whole PT and central-third repair tissue were measured at the medial/central/lateral and proximal/central/distal portions of the tendon, respectively, using digital calipers (accurate to

0.1mm). Tendon length was measured on the posterior surface from tibial insertion to patellar insertion. Whole PT and repair tissue thickness were measured using a light force (<0.15 N) digital micrometer (accurate to 0.01mm; IDCtype Mitutoyo Digimatic Indicator, MTI Corp.,

Aurora, IL). Due to the ellipsoid shape of the PT, whole PT width, thickness and cross-sectional area (width*thickness) are reported by region (proximal/central/distal). After recording tissue dimensions, the tibia/fibula complex was cut roughly 2.5cm distal to the tibial tuberosity to create a bone block. Patellar and tibial bone blocks were fixed into custom designed grips using polymethylmethacrylate cement (Dentsply International , York, PA).81, 164 The testing grips for each sample were secured into a Plexiglas tank mounted on a materials testing system (Model

8501, Instron, Inc., Canton, MA). The tank contained phosphate buffered saline (PBS; pH 7.4) heated to 37°C. Repair tissues were preloaded, preconditioned for 50 cycles to 3% strain at 1Hz using a sinusoidal strain pattern and then failed in uniaxial tension at a constant strain rate of

20%/sec while force and displacement were continuously recorded.78, 81, 144, 164

Histological Evaluation Tissue processing has been described previously107 but a brief protocol is outlined below.

(patella and proximal PT, tibia and distal PT) were processed by cutting eight serial sections

Statistical Analysis Data were normal and homoscedastic. At 6 weeks post-surgery, two animals had bilateral

PT ruptures and one had a unilateral rupture of the PTA. At 12 weeks post-surgery, one animal

had bilateral PT ruptures and three had unilateral ruptures, two of the PTA and one of the dcPTA. At 26 weeks post-surgery, three animals had bilateral PT ruptures and two had unilateral ruptures, one each of the PTA and dcPTA. Ruptured samples were not included in the statistical analysis. Multiple one-way ANOVAs were performed to determine the effect of the native cell population and recovery time. Post-hoc testing was conducted using Fisher‟s protected LSD. The significance level was set at p < 0.05.

Results Repair Tissue Dimensions (Table 3) Increases in repair tissue dimensions were generally observed from 6 to 12 weeks post-surgery.

However, repair tissue dimensions typically remained the same between 12 and 26 weeks or decreased with time.

Effect of Recovery Time: For dcPTA repairs, whole PT width in the proximal region and central- third repair tissue cross-sectional area were both significantly affected by recovery time (p <

0.05). Whole PT width in the proximal region increased from 6 to 12 weeks post-surgery (p <

0.05). However, it was not different between 6 and 26 week repairs or between 12 and 26 week repairs. Central-third repair tissue cross-sectional area decreased from 12 to 26 weeks post- surgery (p < 0.05). However, it was not different between 6 and 12 week repairs or between 6 and 26 week repairs. For PTA repairs, whole tissue cross-sectional area in the proximal region and central-third repair tissue width were both significantly affected by recovery time (p < 0.05).

Whole tissue cross-sectional area in the proximal region and central-third repair tissue width at 6 weeks post-surgery were both less than values for 12 and 26 week repairs (p < 0.05). However, neither parameter varied between 12 and 26 weeks post-surgery.

Effect of Graft Type: At 6 weeks post-surgery, both whole tissue thickness and cross-sectional area in the proximal region were greater in dcPTA repairs as compared to PTA repairs (p <

0.05). At 12 weeks post-surgery, no tissue dimensions were different between dcPTA and PTA repairs. At 26 post-surgery, isolated central-third cross-sectional area was larger for PTA repairs versus dcPTA repairs (p < 0.05).

Table 3. Repair Tissue Dimensions (Mean ± SD) for Whole and Central-Third PT Repair Tissues 6 Weeks 12 Weeks 26 Weeks Variable Region1 dcPTA PTA dcPTA PTA dcPTA PTA n = 5 n = 4 n = 5 n = 5 n = 3 n = 3 Whole Tissue CSA Proximal 34.4 ± 4.3 26.3 ± 2.6a,b,c 40.3 ± 10.4 38.2 ± 6.3 36.1 ± 4.6 38.3 ± 5.7 2 (mm ) Central 37.6 ± 4.9 29.5 ± 2.6 43.6 ± 5.7 49.1 ± 17.1 37.8 ± 3.2 32.8 ± 5.7 Distal 32.1 ± 9.4 23.5 ± 7.6 30.3 ± 5.8 34.5 ± 13.1 26.5 ± 3.7 28.1 ± 6.4 Width Proximal 9.4 ± 0.7a 9.2 ± 0.9 11.1 ± 0.9 10.9 ± 0.8 10.3 ± 0.7 10.6 ± 1.1 (mm) Central 11.3 ± 1.5 10.8 ± 1.2 13.6 ± 0.7 14.3 ± 2.4 12.7 ± 1.8 12.9 ± 0.8 Distal 10.1 ± 1.4 9.2 ± 0.8 9.7 ± 1.1 10.4 ± 1.4 9.7 ± 0.1 11.3 ± 1.5 Thickness Proximal 3.7 ± 0.3 2.9 ± 0.6c 3.4 ± 0.7 3.5 ± 0.9 3.5 ± 0.2 3.6 ± 0.4 (mm) Central 3.4 ± 0.4 2.7 ± 0.7 3.1 ± 0.1 3.3 ± 0.9 3.0 ± 0.2 2.6 ± 0.5 Distal 3.2 ± 0.9 2.5 ± 0.7 3.2 ± 0.5 3.4 ± 1.3 2.7 ± 0.4 2.5 ± 0.4 Length (mm) 22.3 ± 2.1 22.1 ± 1.7 22.5 ± 2.3 23.9 ± 3.0 22.7 ± 1.9 22.3 ± 2.9 Central-Third Repair 2 CSA (mm ) 7.7 ± 0.9 7.4 ± 1.4 8.1 ± 0.9 9.0 ± 2.6 6.3 ± 0.4a,d 7.6 ± 0.4 Width (mm) 2.8 ± 0.3 2.8 ± 0.1 3.0 ± 0.2 3.0 ± 0.2 2.8 ± 0.3 3.1 ± 0.02 Thickness (mm) 2.7 ± 0.4 2.7 ± 0.6 2.8 ± 0.5 3.0 ± 0.8 2.3 ± 0.2 2.5 ±0.2 Length (mm) 22.7 ± 2.0 22.7 ± 1.5 22.7 ± 1.8 24.0 ± 4.2 22.2 ± 2.7 22.1 ± 3.4 1 If a region is not indicated, the dimension listed reflects either the patellar tendon or central- third repair as a whole (i.e. the average value for the three regions). a Less than 12 week repair of same graft material b Less than 26 week repair of same graft material c Less than dcPTA repair at same time point d Less than PTA repair at same time point

Repair Tissue Biomechanical Properties (Table 4; Fig. 7) Repair tissue biomechanical properties typically increased from 6 to 12 weeks post-surgery and then remained constant or decreased between 12 and 26 weeks

Effect of Recovery Time: For dcPTA repairs, all biomechanical properties were affected by recovery time (p < 0.05; Fig. 7). The maximum force sustained by dcPTA repairs at 6 weeks was less than 12 week repairs (p < 0.05). However, maximum force did not vary between 6 and 26 week repairs or between 12 and 26 week repairs. Repair tissue linear stiffness and maximum stress at 6 weeks were both less than values for 12 and 26 week repairs (p < 0.05). However, neither linear stiffness nor maximum stress varied between 12 and 26 week repairs. Repair tissue modulus increased across all time points (p < 0.05). Modulus at 6 weeks was less than values for both 12 and 26 week repairs (p < 0.05) and the modulus at 12 weeks was also less than 26 week repairs (p < 0.05). For PTA repair, no biomechanical properties were affected by recovery time.

Based on the similar patterns in biomechanical data demonstrated by the PTA and dcPTA repairs, it was not expected that PTA and dcPTA repairs would have dissimilar statistical outcomes. Power analysis was conducted on all biomechanical comparisons of PTA repairs. For maximum force between 6 and 12 week repairs, statistical power was 64%. Additionally, for maximum force between 6 and 26 week repairs, statistical power was 54%. Therefore, without an increased sample size, we cannot definitively conclude if these properties are not different.

Effect of Graft Type: At all time points, biomechanical properties were not different for dcPTA and PTA repairs (Table 4, Fig. 7).

Repair Tissue Failure Mechanisms

Of the 25 limbs tested, almost 50% primarily failed at the patellar insertion. The remaining repairs primarily failed at either the tibia or both the patella and tibia together (i.e. the anterior aspect failed at the patella while the posterior aspect failed at the tibia). Additionally, there was one mid-substance failure (dcPTA at 12 weeks). There were five cases where bone avulsion occurred: two cases at 6 weeks and one case at 12 weeks occurred at the patella whereas two cases at 26 weeks occurred at the tibia.

Table 4. Biomechanical Properties (Mean ± SD) of Patellar Tendon Autograft and De-Cellularized Patellar Tendon Autograft Repairs 6 Week 12 Week 26 Week Response dcPTA PTA dcPTA PTA dcPTA PTA Measure n = 5 n = 4 n = 5 n = 5 n = 3 n = 3 Maximum 101.5 ± 42.8 110.8 ± 38.1 170.0 ± 31.4 168.6 ± 35.6 153.8 ± 18.1 164.2 ± 30.4 Force (N) Stiffness (N/mm) 48.0 ± 14.7 50.7 ± 13.8 70.2 ± 15.9 65.5 ± 18.9 75.8 ± 11.1 67.3 ± 9.9 Maximum Stress (MPa) 13.0 ± 4.8 15.6 ± 6.6 20.9 ± 2.6 19.8 ± 6.0 24.4 ± 4.0 21.4 ± 3.2 Modulus 138.0 ± 24.5 159.7 ± 55.4 195.3 ± 36.0 178.1 ± 47.7 267.5 ± 60.3 193.6 ± 26.2 (MPa)

Figure 7. Maximum force and stiffness plots as percent of normal. Maximum force and stiffness sustained by dcPTA and PTA repairs relative to normal central-third PT (data obtained previously80, 81; n = 8) at 6, 12 and 26 weeks post-surgery.

Tibial Insertion Site Histology There were four key events that progressed with time for both dcPTA and PTA repairs

(Fig. 8). First, at 6 weeks post-surgery, a bone nodule was evident on the anterior face of the tibia distal to the insertion site. The bone nodule was surrounded by tendon on both its anterior and posterior surfaces. By 12 weeks, the bone nodule had lengthened and by 26 weeks it was no longer a distinguishable structure.134 Second, at 6 weeks post-surgery, the articular cartilage

lining the tibial tubercle no longer separated the PT from the bone (as opposed to in the native strut, NS, where the articular cartilage and PT are two distinct structures). At 12 weeks, there was evidence of integration between the tendon and tibial tubercle and chondrocytes were present at both the proximal aspect of the tubercle and the proximal aspect of the insertion site region. By 26 weeks, cartilage-like tissue was present between the tendon and tibia and separation between the tissues started originating at the proximal aspect of the insertion site.

Third, integration at the osteotendinous junction improved with time. The amount and organization of fibrocartilage increased with time and, by 26 weeks a tidemark was evident in both dcPTA and PTA repair insertion sites. Additionally, tendon fiber organization increased with time and, by 26 weeks, parallel rows of fibrochondrocytes and tendon fibers were present at the insertion site. However, the fibers of the insertion site were not completely continuous with the tendon body as seen in the native strut (NS). Fourth, cellular morphology changed from mostly rounded at 6 weeks to a mix of rounded and elongated at 12 weeks. By 26 weeks, the cells were mostly elongated. Despite the change in morphology, the tendon of both dcPTA and

PTA repairs was hypercellular through 26 weeks post-surgery. The presence of elongated cells at

12 weeks was partnered by initial observations of crimp in the tendon.

There was also one important observation that was consistent across all repairs at all time points: Osteotendinous integration was relatively enhanced closer to the native struts than in the center of the defect site. Moving from the center outward, the fibrocartilage became more organized and the invaginations with bone became deeper. The defect site, a transitional area and the NS were generally found in successive levels.

T C

MFC/F MFC/FC

MFC/FC

MFC/FC FS FS

T T

MFC/FC

B B

Figure 8. Tibial insertion sites for de-cellularized PTA (dcPTA), PTA and native struts (NS) at 6, 12 and 26 weeks post-surgery. H&E stained images are labeled to reflect: B, bone; T, tendon; FS, fibrous scar (note: a region of fibrous scar is not present in 6 week dcPTA repairs); MFC/FC, mineralized fibrocartilage/fibrocartilage zones of the insertion site, respectively; areas resembling neo-vasculature are circled. Scale bar = 200µm.

Discussion

In this study, we examined tendon graft healing in an extra-articular environment.

Therefore, the biological processes mediating healing in this model could either follow those described for ACL graft healing and incorporation32, 135-137 or those described for tendon healing4. Graft healing and incorporation can be described in three overlapping phases: 1) early graft healing (up to about 4 weeks post-surgery) involving tissue necrosis and hypocellularity, inflammatory cell infiltration, and initial matrix remodeling between 1-2 weeks post-surgery caused by an influx of cells replacing the native cell population; 2) revascularization and proliferation phase (between 4 and 12 weeks post-surgery) where maximal remodeling activity occurs, mechanical strength deteriorates until roughly 6-8 weeks post-surgery, and enhanced revascularization occurs from about 4-12 weeks post-surgery; 3) remodeling and ligamentization phase (begins around 12 weeks post-surgery with unknown endpoint) where remodeling of the graft is continued toward the morphology and strength of the native ACL and cellular morphology slowly changes from the ovoid shape of the metabolically active tenoblast to the elongated spindle shape of the less active tenocyte between roughly 3 and 24 weeks.32, 135-137

Tendon healing has also been described in three similar overlapping phases4. However, tendon healing can be achieved through both intrinsic (proliferation of epitenon and endotenon cells) and extrinsic (infiltration of tendon sheath and inflammatory cells) cellular mechanisms4. The tissue necrosis and re-population aspects of ACL graft remodeling essentially eliminate the

feasibility of intrinsic cellular contributions. Intrinsic healing yields improved biomechanical properties whereas extrinsic healing generates scar tissue. Extrinsic healing tends to predominate in rotator cuff injuries where fibrous scar tissue is typically found at the tendon-to-bone insertion. Based on the results of this study which demonstrate no biomechanical benefit of the native cell population, we propose that extra-articular graft healing undergoes a healing and incorporation process similar to that described for intra-articular ACL grafts. However, further study is needed to validate this theory.

The increased width and cross-sectional area measurements typically observed at 12 weeks post-surgery for dcPTA and PTA repairs are likely due to cellular mechanisms occurring during the proliferative phase of healing135. While increases in tissue width and cross-sectional area have been reported for the early phase of graft healing, it is likely that our 6 week time point was too late to observe these changes. However, cellular infiltration has been shown to elicit an increase in cross-sectional area and a deterioration of biomechanical properties.176 This response is likely due to increased expression of collagen type III by cells that originated extrinsically.177

The increased biomechanical properties of dcPTA repairs at 12 weeks post-surgery could indicate that the tissue is in the remodeling phase of healing. This was corroborated histologically by the presence of both elongated cell nuclei and regions of crimp in 12 week repairs135. As the grafts progress through remodeling, large diameter fibrils are commonly replaced by smaller diameter fibrils which provide less mechanical strength135. This shift in collagen fibril diameter could help explain why biomechanical properties did not improve from

12 to 26 weeks post-surgery. Additionally, a persistent increase in collagen type III expression might contribute to the threshold of biomechanical properties seen in vivo.135, 177

The results of this study agree with previous studies that show biomechanical properties reach a threshold in animal models and do not progress past this point regardless of recovery time.135, 137, 178, 179 For ACL reconstruction in animal models, it has been reported that, even at one year (or longer) after surgery, repair tissue structural properties do not achieve more than 50-

60% of the native tissue.134, 135, 137, 178, 179 The biomechanical properties of dcPTA and PTA repairs were not different at any time point and therefore the repair results were pooled before being contrasted against normal central-third PT. The collective repair tissue biomechanical properties of dcPTA and PTA repairs reached 22%, 36% and 34% of normal central-third PT maximum force at 6, 12 and 26 weeks post-surgery, respectively.80, 81 The collective autograft repairs also achieved 31%, 42% and 45% of normal central-third PT stiffness at 6, 12 and 26 weeks post-surgery, respectively.80, 81 We propose that animals reach this threshold in repair capacity because of the limited physical loads to which they are exposed. Unlike humans who receive specialized rehabilitation, animals are generally allowed unrestricted cage activity with or without a period of immobilization or casting. It is possible that without implementing physical therapy into our paradigm of repair, we may not surpass the current level of success.

The results of this study also agree with reports that tendon-to-bone insertion sites begin to show signs of maturation at 12 weeks and demonstrate some osteotendinous integration accompanied by a tidemark at 26 weeks.115 It is unclear whether a zonal insertion site would be regenerated even with longer durations of recovery. Our laboratory, along with others131, 132, 139, have attempted multiple augmentation strategies to improve osteotendinous integration with little success to date.

One limitation of this study is the sample size. Although dcPTA and PTA repairs showed similar patterns in their biomechanical properties, only dcPTA repairs were affected by recovery

time (p < 0.05). Power analysis revealed that with a reasonable increase in sample size (n = 12), it may be possible to demonstrate that PTA repair tissue maximum force improved with recovery time. This was not true for the other biomechanical properties. A second limitation is the user- bias introduced when isolating central-third repair tissues. Although the differences in isolated central-third repair tissue cross-sectional area for dcPTA and PTA repairs could be attributed to variations in thickness, it is possible that the discrepancies in repair tissue width for PTA samples were influenced by the user (i.e. central-third repair tissue was cut too thin when removing the native struts for biomechanical testing). Lastly, our model does not replicate graft replacement or surgical reconstruction of a tendon or ligament tear. However, our model does provide a reproducible and accessible system to test tissue engineering strategies to promote tendon healing and integration into bone.

The objective of this study was to understand the role of the native cell population on graft incorporation. We hypothesized that osteotendinous integration and biomechanical properties of both dcPTA and PTA repairs would improve with recovery time. Osteotendinous integration improved with recovery time for both dcPTA and PTA repairs (based on subjective observation). However, biomechanical properties were only affected by recovery time for dcPTA repairs (p < 0.05). Repair tissue modulus increased with time (p < 0.05) while the remaining biomechanical properties increased from 6 to 12 weeks (p < 0.05) but did not show continued improvement up to 26 weeks. Despite the changes in biomechanical properties demonstrated by dcPTA repairs, biomechanical properties did not vary between dcPTA and PTA repairs at any time point. Therefore, in this model of tendon healing, the native cell population offered no histological or biomechanical benefit and it is likely the PTA tissues went through healing and

incorporation processes similar to that described for intra-articular graft healing. However, additional testing is needed to validate this theory.

Chapter 5

MSC-Collagen Gel Augmentation to Enhance Osteotendinous Integration of a Patellar Tendon Autograft* Kirsten R.C. Kinneberg1, Marc T. Galloway2, David L. Butler1, Jason T. Shearn1

1. School of Energy, Environmental, Biological and Medical Engineering, Biomedical Engineering Program, University of Cincinnati, Cincinnati, OH; 2. Cincinnati Sportsmedicine, Cincinnati, OH *Manuscript is in preparation and will be submitted for publication in the journal Tissue Engineering. .

Osteotendinous integration is required for many orthopaedic surgical procedures including reattachment after tendon avulsion, tendon transfer, and ligament reconstruction.

Healing of the tendon-to-bone junction is important for joint stability and function.114, 132, 139

However, current repairs initially develop a fibrous scar at the osteotendinous junction in place of the zonal insertion seen in the uninjured tissue.115, 132 In fact, the high rate of failure for rotator cuff repair is often attributed to deficient tissue integration at the tendon-to-bone insertion site.132

A normal zonal insertion transitions from tendon to fibrocartilage, mineralized fibrocartilage and bone.25, 26, 132, 133 The complex enthesis connecting tendon to bone serves to mediate load transfer and minimize stress concentrations formed between the relatively compliant tendon and the relatively stiff bone. 25, 26, 132, 133 The limiting factor in regenerating this tissue appears to be bone ingrowth.115, 131, 132, 139 Accordingly, many attempts have been made to promote bone ingrowth and facilitate rapid healing of this tissue. Rodeo et al. have used both bone morphogenic protein-2 (BMP-2)131 and an osteoinductive bone protein extract132 to create growth factor-collagen sponge implants. Their implants increased bone production adjacent to the tendon and improved repair tissue biomechanical properties at various time points relative to an untreated sponge implant. However, when the increases in strength produced by the bone

protein extract-collagen sponge implant were normalized by tissue volume, no differences were found.132 This finding is consistent with the theory that growth factor treatments produce large amounts of poor quality scar tissue rather than regenerated tendon.138, 139

The first objective of this study was to use mesenchymal stem cells (MSCs) in a collagen type I gel carrier to promote bone ingrowth at the tendon-to-bone insertion site. We postulated that implanting a biologic augmentation (BA) composed of key constituents of an insertion site, namely collagen type I to span from tendon to bone and MSCs capable of differentiating into chondrocytes, osteoblasts and fibroblasts, would improve osteotendinous integration between a patellar tendon autograft (PTA) and the underlying bone. Our ultimate goal was to restore the continuity between the collagen fibers in tendon and the subchondral bone in the patella and tibia because, as tissue integration progresses, attachment strength improves131.

The second objective of this study was to understand how cellular activity and mRNA expression levels in vitro influence repair outcomes in vivo. To do this, MSC-collagen gel BAs were created with two different cell densities, in two different culture systems and for two lengths of time in culture. Using an alkaline phosphatase (ALP; early marker for bone) activity assay and quantitative real-time PCR (qRT-PCR) to determine mRNA expression levels for both osteo- and chondrogenic genes of interest, a treatment group was identified with increased ALP activity and mRNA expression levels relative to the BA implanted in the first iteration of surgery. This BA was then implanted in the second iteration of surgical implantation. By monitoring cellular activity and expression levels in vitro and evaluating insertion site formation in vivo, we hope to gain an understanding of how cellular expression at implantation affects surgical outcomes for MSC-collagen gel TECs. Further, if we can use in vitro culture conditions

to enhance the osteo- and/or chondrogenic capacity of our TECs, perhaps the osteotendinous junction will develop and mature more quickly.

Our hypotheses were as follows: For both iterations of surgical implantation, MSC- collagen gel BAs will improve 1) tendon integration with bone at the patellar and tibial insertion sites as compared to PTA alone, and 2) repair tissue biomechanical properties relative to PTA at

12 weeks post-surgery. Additionally, MSC-collagen gel BAs with superior cellular activity and mRNA expression levels (second iteration of surgical implantation) will enhance insertion site formation and repair tissue biomechanical properties over the initial BA repairs (first iteration of surgical implantation) and PTA repair alone.

Experimental Design

All procedures were approved by the University of Cincinnati Institutional Animal Care and Use Committee. A total of 18 one year old, skeletally mature, female New Zealand white rabbits were used for this study. Two iterations of surgical implantation were performed in this study and 9 animals were dedicated to each (Fig. 9). Autologous mesenchymal stem cells

(MSCs) were isolated from the bone marrow of each animal. MSCs were subcultured to passage two before being used to make MSC-collagen gel biologic augmentations (BAs). MSC-collagen gel BAs were fabricated with a cellular density of 100K cells/ml, in a silicone dish, for two weeks before being implemented in our patellar tendon autograft (PTA) model of tendon healing.

At surgery, the central-third patellar tendon was removed from both knees of each rabbit. Bone defects were created at the patellar and tibial insertion sites. In the right limb, a MSC-collagen gel BA was placed in each insertion site posterior to the PTA (PTA+BA). In the left limb, the

excised PTA was sutured back into the defect site with no BAs (PTA). After 12 weeks of healing, limbs were collected for either biomechanical (n = 7) or histological (n = 2) analyses.

To understand how cellular activity and mRNA expression levels in vitro influence repair outcomes in vivo, a second iteration of surgery was performed utilizing a MSC-collagen gel BA with a different activity/expression level profile than the BA implanted originally. The BA for the second iteration of surgery was selected through an in vitro optimization study (BAs were created using autologous cell lines from animals dedicated to the second iteration of surgical implantation). We determined the effects of cellular density, cell culture system, and time in culture on ALP activity and mRNA expression of osteo- and chondrogenic genes in vitro using real time quantitative reverse-transcriptase PCR (qRT-PCR). ALP activity was evaluated using a colorimetric kit (AnaSpec, San Jose, CA). Using qRT-PCR, we examined both osteo- and chondrogenic genes because the osteotendinous junction has sections of both and, although bone ingrowth has been cited as a limiting factor for insertion site integration115, 131, 132, 138, 139, we could not rule out the influence of chondrogenic factors. For the osteogenic pathway, we tested for the protein collagen type I and the transcription factor runx2.180-182 For the chondrogenic pathway, we tested for the protein collagen type II and the transcription factor sox9.183-185

Collagen type I predominates in the fibrocartilage zone (FC) of the insertion site and collagen type II predominates in the mineralized fibrocartilage zone (MFC).25, 43, 44 Runx2 is a critical transcription factor for osteoblast differentiation and activation of runx2 is associated with increased expression of osteoblast-specific genes, such as alkaline phosphatase and osteocalcin, and increased bone formation.180-182 Sox9 activates COL2A1 transcription and acts cooperatively with other Sox genes to regulate chondrogenesis in vivo. 183-185

The biologic activity of cells is largely determined by both cell-cell and cell-matrix interactions. Cell-cell interactions are regulated by cell density and it has been demonstrated that both cellular density and cellular shape modulate MSC phenotype.80, 81, 130, 140-143 Therefore, we wanted to test both high and low cellular densities to gauge how MSCs would react in the collagen gel matrix. At the low end, we chose 100K cells/ml because, in our experience, this is near the lower threshold of contraction for MSCs inoculated in collagen gel (in our silicone dish culture system). This was also the cellular density of MSC-collagen gel BAs used for the first iteration of in vivo implantation. At the high end, we chose 600K cells/ml because this was roughly the highest cellular density that could be maintained for two weeks in the silicone dish without pre-mature failure (i.e. the MSC-collagen gel contracting so much that it tore off the restraining post in the well of the silicone dish). Another regulator of cellular density is matrix contraction; as cells contract the collagen gel matrix around them, the cellular density inherently increases. To determine if cell-mediated contraction would influence cell-cell interactions, MSC- collagen gel BAs were fabricated in two different cell culture systems, in a silicone dish186-188 and around a taut suture in a glass boat 80, 189, 190, which each allow for different amounts of contraction. MSC-collagen gel BAs were also cultured for 7 and 14 days because time in culture can have a significant impact on cellular behavior.141, 147-149

The MSC-collagen gel BA treatment group with mRNA expression levels superior to those for BAs utilized in iteration 1 of surgical implantation was chosen for iteration 2 of surgical implantation (Fig. 9). Based on the differences in gene expression levels between the two MSC-collagen gel BAs, the treatment group utilized in iteration 1 of surgical implantation with relatively lower gene expression levels will be referred to as BAs with low-activity (BA-

LA); the treatment group with relatively higher expression levels will be referred to as BAs with

high-activity (BA-HA). MSC-collagen gel BA-HAs were implanted, as described above, in iteration 2 of surgical implantation. After 12 weeks of healing, limbs were collected for either biomechanical (n = 7) or histological (n = 2) analyses.

Iteration 1 of Surgical Implantation (n = 9) MSC-collagen gel BAs fabricated with a cellular density of 100K cells/ml, in a silicone dish, cultured for 14 days; MSC-collagen gel BAs implanted in vivo. Animals allowed 12 weeks for recovery. Repair tissue biomechanical properties (n = 7) and histologic outcomes (n = 2) assessed.

In Vitro Optimization MSC-collagen gel BAs fabricated with: Cellular Density Culture System Culture Period

100K cells/ml Silicone dish 7 days

600K cells/ml Glass boat 14 days

Gene expression levels normalized to BA from iteration 1 of surgical implantation. MSC-collagen gel BA treatment group with mRNA expression levels superior to those for BAs utilized in iteration 1 of surgical implantation was chosen for iteration 2 of surgical implantation. Based on mRNA expression levels:

MSC-collagen gel BAs from iteration 1 of surgical implantation = BAs with relatively low activity: BA-LA BAs selected by vitro optimization = BAs with relatively high activity: BA-HA

Iteration 2 of Surgical Implantation (n = 9) MSC-collagen gel BA-HAs implanted in vivo. Animals allowed 12 weeks for recovery. Repair tissue biomechanical properties (n = 7) and histologic outcomes (n = 2) contrasted against both PTA+BA-LA and PTA repairs. Figure 9. Experimental Design Overview for MSC-Collagen Gel BA Surgical Implantation and In Vitro Optimization.

Materials and Methods Mesenchymal Stem Cell Harvest MSCs were isolated from the iliac crest of each animal (n = 18) and processed as previously described.81, 130, 140, 164, 165, 186 Briefly, each aspirate was mixed with 25ml of complete advanced-DMEM (complete ADV-DMEM; ADV-DMEM supplemented with 5% FBS, 1%

Gluta-MaxTM, 1% Antibiotic-Antimycotic) and then centrifuged at 2000rpm for 5 minutes. After performing this step twice for each aspirate, the remaining pellet was re-suspended in 10ml complete ADV-DMEM, counted and plated at 35-45x106 cells per 75mm2 T-flask. Cells were cultured for 10-14 days until they became 70-90% confluent. At this point, they were subcultured to passage two (P2).

MSC-Collagen Gel Biologic Augmentation Creation Collagen gel was prepared at a concentration of 2.6mg/ml186, 188, 191 according to the manufacture‟s protocol (InvitrogenTM protocol 5024) with the exception that de-ionized water was replaced with a mixture of phosphate buffered saline (PBS), Dulbecco‟s Modified Eagle‟s

Medium (DMEM) and MSCs at a concentration of either 100K or 600K cells/ml. For MSC- collagen gel BAs cultured in a silicone dish, a 1.4ml volume of the cell-gel mixture was pipetted into the well of a silicone dish and allowed to set in an incubator (37°C, 5% CO2) for 2-4 hours before being supplemented with up to 1ml MSC feeding media (complete ADV-DMEM supplemented with 5% L-ascorbic acid (Gibco®)). For MSC-collagen gel BAs cultured in a glass boat, a monofilament synthetic absorbable suture (MaxonTM 4-0; Henry Schein, Inc.,

Melville, NY) was tied around a v-shaped K-wire to keep the suture taut and laid across the long axis of a hemispherical glass boat, as previously described.80, 189, 190 A 350µl volume of the cell- gel mixture was pipetted into the boat around the suture. MSC-collagen gel BAs were allowed to

set in an incubator (37°C, 5% CO2) for 2-4 hours before being supplemented with 35-40ml of

MSC feeding media. Based on the cell number-to-media volume ratio, MSC feeding media was changed every third day for the glass boats (35-40ml) and every second day for the silicone dishes (1-2ml). At the end of the two week culture period, BAs were transported to surgery, collected for ALP activity assay, or collected for real time quantitative reverse-transcriptase PCR

(qRT-PCR).

Alkaline Phosphatase Activity and MTT Assays Note: At the time of submission of this dissertation, the alkaline phosphatase activity was normalized by cell number using MTT optical density. The ALP activity will be re-tested and normalized by a more accurate double stranded DNA (dsNDA) assay. However, until our spectraphotometer is fixed, the ALP activity data cannot be re-tested and validated. The results included in this dissertation are those normalized by MTT. Alkaline phosphatase (ALP) activity was assessed using the SensoLyteTM pNPP Alkaline

Phosphatase Colorimetric Assay Kit (AnaSpec, San Jose, CA). Briefly, MSC-collagen gel BAs were cut in half; one half was used for ALP activity and the other was used for MTT. Briefly, for

ALP activity, BAs were homogenized in lysis buffer, centrifuged and the supernatant was collected for analysis. All reagents were made according to the manufacture‟s protocol. The p-

Nitrophenyl phosphate (pNPP; colorimetric phosphatase substrate) reaction mixture was incubated (25oC) with the sample supernatant for 30 minutes before the 96-well plate was read at

OD 405nm. For MTT, paired halves were washed with PBS to remove serum proteins and then incubated with MTT solution (1mg/ml in PBS) on an orbital shaker for 3 hours at 37oC and 5%

CO2. At the end of this incubation period, BAs were washed with PBS to remove remnant MTT contamination and then formazan product was obtained by incubating BAs with methoxyethanol on an orbital shaker for 3 hours at room temperature (25oC). The 96-well plate was read at OD

590nm. ALP activity (OD 405nm) was normalized by MTT (OD 590nm) to obtain ALP activity/cell.

Real Time qRT-PCR BAs dedicated to qRT-PCR were collected in 1ml of TRIzol® Reagent in an RNase free conical tube. An RNase-free pestle was used to break up the gels and allow full saturation with

TRIzol® Reagent. After incubating at room temperature for five minutes, samples were stored at

-80°C. As described below, MSC-collagen gel TECs cultured in a silicone dish were too large to fit in the bone trough and therefore were appropriately trimmed (the regions contracted around the restriction posts were excluded for in vivo implantation and they were excluded from samples collected for real-time qRT-PCR).

RNA was isolated from the TRIzol® Reagent suspension according to the manufacture‟s protocol. RNA pellets were dissolved in 15-20µL of RNase-free water; re-suspension volumes were decided based on the size of the RNA pellet. The total amount of RNA isolated from each sample was quantified using a Qubit® Fluorometer and Qubit® RNA Assay Kit. The appropriate amount of RNA was then converted to cDNA using the High Capacity RNA-to-cDNA kit from

Applied Biosystems and a VeritiTM Thermal Cycler (Applied Biosystems). Real-time qPCR assays were performed in a StepOnePlusTM thermocycler (Applied Biosystems) using TaqMan gene expression assay kits (Applied Biosystems) for the genes listed in Table 5. Each reaction contained 1.0µL TaqMan® Gene Expression Assay Mix (20X), 10.0µL TaqMan® Fast

Universal PCR Master Mix (2X; no AmpErase® UNG) and appropriate volumes of cDNA template and RNase-free water to make a total reaction volume of 20.0µL. Volumes of cDNA varied based on the concentration of RNA re-suspended in RNase-free water (all reagents from

Applied Biosystems). Sealed 96-well plates were subjected to the following thermal cycling

conditions: 95°C for 20 sec, followed by 40 cycles of 95°C for 1 sec and 60°C for 20 sec. Each sample was run in duplicate and all the samples for each cell line were run on the same day, using the same master mix, to avoid variability. Prior to conducting real time qRT-PCR on the experimental groups, a housekeeping panel of six genes was run on previously collected test samples (see Table 5). For both the housekeeping test panel and the experimental groups, raw Ct values generated by the TaqMan® gene expression assays were analyzed using StepOneTM software (Applied Biosystems). Data are reported as fold change relative to BA-LA (iteration 1 of surgical implantation) using the 2^-[delta][delta]Ct method.

Table 5. Gene Names*, TaqMan® Assay Identification (ID) Numbers, and Amplicon Length for All Tested Genes Amplicon Gene Name Assay ID Length Proteins COL1A1 Oc03396073_g1 70 COL2A2 Oc03396134_m1 70 Transcription Factors RUNX2 Oc02386741_m1 91 SOX9 Oc04096872_m1 61 Endogenous Controls GAPDH Oc03823402_g1 82 18S Hs99999901_s1 187 28S Hs0365441_s1 98 HPRT1 Oc03399466_g1 141 PPIA Oo03396990_g1 98 ACTB Oc03824857_g1 106 *COL1A1, collagen type I; COL2A2, collagen type II; GAPDH, Glyceraldehyde 3-phosphate dehydrogenase; 18S, eukaryotic 18S rRNA; 28S, eukaryotic 28S rRNA; HPRT1, hypoxanthine phosphoribosyltransferase 1; PPIA, peptidylprolyl isomerase A (cyclophilin A); ACTB, Beta- actin

MSC-Collagen Gel Biologic Augmentation Implantation At surgery, each rabbit was anesthetized using an anesthesia cocktail of ketamine/ acepromazine (40mg/kg and 5mg/kg, respectively) and isoflurane gas (as needed). Under aseptic conditions, each patellar tendon was exposed through an anteromedial incision. A 3 mm wide, full-length, full-thickness, central-third defect was created in each PT, producing a soft tissue, patellar tendon autograft (PTA). Bone defects were created at the proximal and distal insertion sites using a pneumatic sagittal saw (MicroAire Surgical Instruments, Charlottesville, VA). In the right limb, a MSC-collagen gel BA was placed in the bone trough of both the patella and tibia. For iteration 1 of surgical implantation, MSC-collagen gel BA-LAs cultured in a silicone dish were too large to fit in the bone trough so the BA-LAs were trimmed to exclude the regions contracted around the two restraining posts. For iteration 2 of surgical implantation, MSC- collagen gel BA-HAs were carefully isolated using a scalpel blade to hold the gel and a forceps to slowly pull out the suture. After positioning the BAs in the bone troughs, the excised PTA was then secured back into the defect site on top of the BAs (PTA+BA; Iteration 1, PTA+BA-LA;

Iteration 2, PTA+BA-HA) using four sutures (Prolene 5-0, Ethicon, Somerville, NJ): two each near the patella and tibia. Each suture incorporated both the native strut and the BA to help keep the BAs in place. In the contralateral limb, the PTA was sutured back into the defect site in the same manner as the right limb but with no BAs in the bone troughs (PTA). The incisions were closed and the animals were allowed to recover before being returned to their individual cage.

Animals were allowed unrestricted cage activity. After 12 weeks of recovery, each rabbit was anesthetized as described above and then euthanized by intracardiac administration of Euthasol®

(1.0cc/4.54kg; Virbac Animal Health, Inc., Fort Worth, TX) followed by bilateral pneumothorax.

Hind limbs were disarticulated at the hip and removed using a scalpel. For biomechanical evaluation, whole limbs were stored at -20°C until testing. For histological evaluation, patella-

PT-tibia samples were placed in 10% neutral buffered formalin (NBF) and stored at room temperature (25°C) until further processing.

Biomechanical Evaluation One day prior to testing, limbs were removed from the freezer and placed in a cold water bath to thaw. On the day of testing, all extraneous tissues were removed leaving only the patella, patellar tendon (PT) and tibia. Whole PT and repair tissue dimensions were obtained, as described in Chapter 3. Briefly, the length and width of both the whole PT and repair tissue were measured at the medial/central/lateral and proximal/central/distal portions of the tendon, respectively, using digital Vernier calipers (accurate to 0.1mm). Tendon length was measured on the posterior surface from tibial insertion to patellar insertion. Whole PT and repair tissue thickness were measured using a light force (<0.15 N) digital micrometer (accurate to 0.01mm;

IDCtype Mitutoyo Digimatic Indicator, MTI Corp., Aurora, IL). Due to the ellipsoidal shape of the PT (in both the coronal and sagittal planes), whole PT width, thickness and cross-sectional area (width*thickness) are reported by region (proximal/central/distal). After recording tissue dimensions, the tibia/fibula complex was cut roughly 2.5cm distal to the tibial tuberosity to create a bone block. Patellar and tibial bone blocks were fixed into custom designed grips using polymethylmethacrylate cement (Dentsply International , York, PA).81, 164 The testing grips for each sample were secured into a Plexiglas tank mounted on a materials testing system (Model

20%/sec while force and displacement were continuously recorded.78, 81, 144, 164

Histological Evaluation Patella-PT-tibia samples were processed for histological evaluation as described in

Chapter 3.165 Briefly, tissues were fixed in 10% NBF for 60-72 hours before being sectioned in the transverse plane to isolate a) the patella and proximal PT, b) the tendon mid-substance, and c) the tibia and distal PT. Patellar and tibial bone ends were decalcified in 10% formic acid

(Shandon‟s TBD 2; Fisher Scientific) until an endpoint test with 5% ammonium oxalate did not precipitate calcium oxalate (roughly two weeks for patellas and three weeks for tibias). Tissue sections were then dehydrated through a gradient of alcohols and xylene before being embedded in a paraffin block. The bone ends (patella and proximal PT, tibia and distal PT) were processed by cutting eight serial sections (4µm thick) in the sagittal plane at 1000µm intervals through the sample. These tissues were used to examine tendon integration into bone and insertion site formation. The mid-substance samples were processed by cutting eight serial sections in the coronal plane at 150µm intervals. These tissues were used to examine PTA integration with the tendon native struts (NS). One section from each interval was stained with hematoxylin and eosin (H&E). Repair tissue organization, cellularity, neovascularization, repair tissue integration into the native tendon struts, and tendon-bone insertion site formation were evaluated.

Statistical Analysis Biomechanical and dimensional data for PTA+BA-LA, PTA+BA-HA and PTA repairs were analyzed collectively. Data were normal and homoscedastic. Differences between groups were assessed using a one-way ANOVA with repair treatment set as a fixed factor. Post-hoc testing was conducted using Fisher‟s protected LSD. The significance level was set at p < 0.05.

Real-time qRT PCR data are reported as fold change relative to BA-LA (iteration 1 of surgical implantation) using the 2^-[delta][delta]Ct method (relative fold changes were determined by

StepOneTM software (Applied Biosystems)). Relative fold change data were non-normal. Thus, each treatment group was contrasted against the BA-LA using a non-parametric Mann-Whitney test with p < 0.05 to determine if any of the treatment groups had gene expression levels superior to the BA-LA.

Results The first objective of this study was to use mesenchymal stem cells (MSCs) in a collagen type I gel carrier to promote bone ingrowth at the tendon-to-bone insertion site.

Iteration 1 of Surgical Implantation

Treatment with BA-LA Did Not Improve PTA Repair Tissue Biomechanical Properties or

Osteotendinous Integration. PTA+BA-LA repairs sustained 91%, 87%, 88% and 87% of PTA repair tissue maximum force, stiffness, maximum stress and modulus, respectively (Table 6; Fig.

12). Histologically, PTA+BA-LA repairs showed no evidence of organized fibrocartilage-like tissue at the tendon-to-bone insertion site; the tendon was disorganized near bone and showed very little integration, if any (Fig. 13). PTA+BA-LA repairs actually appeared less integrated at the insertion than PTA repairs (Fig. 13).

Table 6. Biomechanical Properties (Mean ± SD) of Patellar Tendon Autograft and PTA+BA Repairs and Normal Central-Third PT PTA PTA+BA-LA PTA+BA-HA Normal Variable n = 12 n = 7 n = 5 n = 81 Maximum Force (N) 164.0 ± 53.0 149.7 ± 51.1 140.5 ± 30.6 470.7 ± 23.8 Stiffness (N/mm) 69.9 ± 17.8 61.1 ± 12.3 55.7 ± 17.1 159.8 ± 11.1 Maximum Stress (MPa) 21.0 ± 8.0 18.6 ± 5.7 17.7 ± 4.7 100.7 ± 5.6 Modulus (MPa) 200.2 ± 49.6 173.9 ± 26.0 164.7 ± 61.1 861.4 ± 98.5 1Normal central-third PT biomechanical properties were previously obtained80, 81

In Vitro Optimization

Alkaline Phosphatase Activity (Fig. 10) Alkaline phosphatase activity (ALP) was significantly affected by culture system, cellular density, and days in culture (p ≤ 0.007). ALP activity per cell was higher for the glass boat culture system versus the silicone dish (p < 0.001), higher for cellular density of 100K vs. 600K

(p = 0.005) cells/ml and higher at day 7 versus day 14 (p = 0.007).

Gene Expression – Relative Fold Change (Fig. 11) Only three treatment groups showed improvement over MSC-collagen gel BAs fabricated with

100K cells/ml and cultured in a silicone dish for 14 days (BAs designated to have low activity,

BA-LA). MSC-collagen gel BAs created with 100K or 600K cells/ml and cultured in a glass boat for 14 days both had significantly higher expression of collagen types I and II relative to BA-LA

(p < 0.05). MSC-collagen gel BAs fabricated with 600K cells/ml and cultured in a glass boat for

7 or 14 days both had significantly higher expression of collagen type II and sox9 relative to BA-

LA (p < 0.05).

MSC-Collagen Gel BAs fabricated with 100K cells/ml and Cultured in a Glass Boat for 14 Days

Chosen for Iteration 2 of Surgical Implantation. Although MSC-collagen gel BAs fabricated with either 100K or 600K cells/ml and cultured in a glass boat for 14 days had similar gene expression levels, ALP activity/cell was higher for BAs with 100K cells/ml relative to 600K cells/ml.

Figure 10. Alkaline phosphatase activity per cell. ALP activity (OD 405nm) was normalized using an MTT assay for cell viability (OD 590). Statistical analysis revealed ALP activity/cell was higher for BAs 1) fabricated in the glass boat versus the silicone dish, 2) fabricated with 100K cells/ml versus 600K cells/ml, and 3) cultured for 7 days versus 14 days.

Figure 11. Gene expression of MSC-collagen gel BAs normalized to GAPDH. Data is shown (using a log-base 2 scale) for BAs fabricated with 100K and 600K cells/ml and cultured in a silicone dish (DISH) or in a glass boat (BOAT) for 7 or 14 days. Fold changes in gene expression are relative to the treatment group utilized for iteration 1 of surgical implantation: MSC-collagen gel BAs fabricated with 100K cells/ml and cultured in a silicone dish for 14 days (fold change is 1 for every gene in this group). * Indicates statistical difference for same gene relative to MSC-collagen gel BAs fabricated with 100K cells/ml and cultured in a silicone dish for 14 days. COL1A1, collagen type I; COL2A1, collagen type II.

Iteration 2 of Surgical Implantation Neither treatment with BA-LA nor BA-HA improved PTA repair tissue biomechanical properties or osteotendinous integration (Table 6; Fig. 12). Additionally, neither PTA+BA-LA, PTA+BA-

HA nor PTA repairs regenerated a zonal insertion site (Fig. 13). Repair tissue biomechanical and histological data are presented below for PTA+BA-LA, PTA+BA-HA and PTA repairs.

Repair Tissue Dimensions (Table 7) With the exception of whole tissue thickness in the distal patellar tendon, treatment with

MSC-collagen gel BA-LA or BA-HA did not affect repair tissue dimensions when compared to

PTA. In the distal patellar tendon, treatment with BA-HA significantly reduced thickness as compared to PTA repair (p = 0.018).

Table 7. Repair Tissue Dimensions (Mean ± SD) for Whole and Central-Third PT Whole Tissue Variable Region1 PTA PTA+BA-LA PTA+BA-HA n = 12 n = 7 n = 5 CSA (mm2) Proximal 31.6 ± 6.0 30.1 ± 9.2 42.9 ± 2.5 Central 34.0 ± 7.5 30.1 ± 6.5 38.0 ± 2.9 Distal 28.8 ± 5.3 24.6 ± 4.3 25.8 ± 4.6 Width (mm) Proximal 10.1 ± 1.5 9.7 ± 1.6 11.7 ± 1.6 Central 12.5 ± 2.1 11.6 ± 1.3 14.3 ± 2.2 Distal 10.3 ± 0.9 9.8 ± 0.7 10.9 ± 0.8 Length (mm) 23.3 ± 2.4 22.7 ± 3.0 23.7 ± 2.3 Central-Third Repair CSA (mm2) 8.1 ± 1.6 8.1 ± 1.0 8.1 ± 1.0 Width (mm) 3.0 ± 0.1 3.0 ± 0.1 3.0 ± 0.1 Length (mm) 22.8 ± 2.7 23.1 ± 3.2 23.5 ± 3.5 1 If a region is not indicated, the dimension listed reflects either the patellar tendon or central- third repair as a whole (i.e. the average value for the three regions).

Repair Tissue Biomechanical Properties and Failure Mechanisms Treatment with MSC-collagen gel BA-LA or BA-HA did not increase PTA repair tissue biomechanical properties at 12 weeks post-surgery (Table 6; Fig. 12). PTA+BA-LA repairs failed primarily at the patella whereas PTA+BA-HA repairs failed primarily at the patella and tibia together (i.e. the anterior face failed at the patella while the posterior face failed at the tibia).

PTA repairs also failed primarily at the patella. There were no mid-substance tendon failures.

There were five cases of bone avulsion: one for PTA+BA-LA repairs, two for PTA+BA-HA repairs, and three for PTA repairs.

Figure 12. Force-displacement curve for PTA and PTA+BA repairs. Patellar tendon autograft (PTA) and PTA repair supplemented with biologic augmentations (BAs) with either relatively low biologic activity (PTA+BA-LA) or relatively high biologic activity (PTA+BA-HA) at 12 weeks post-surgery. These failure curves are contrasted against a normal central-third PT failure curve obtained in a previous study (Normal; n = 880, 81). In vitro force and displacement (IVF and IVD, respectively) are labeled based on data also obtained in a previous study.83, 84

Patellar and Tibial Insertion Site Histology PTA+BA-HA repairs (Fig. 13, top right) appeared better integrated at the tendon-to-bone insertion site than PTA+BA-LA repairs (Fig. 13, top left) at 12 weeks post-surgery but similar to

PTA repairs (Fig. 13, lower left). Both PTA+BA-HA and PTA repairs showed some organization at the tendon-to-bone interface with shallow invaginations of tendon and bone with short rows of organized fibrochondrocyte-like cells (Fig. 13, top row). Conversely, in PTA+BA-

LA repairs there was no evidence of organized fibrocartilage-like tissue; the tendon was disorganized near bone and showed very little integration, if any (Fig. 13). However, all treatments produced fibrous scar (FS) tissue at the insertion; neither treatment produced a zonal insertion site at 12 weeks (Fig. 13). Interestingly, for all treatment groups, tendon was better integrated with bone in regions adjacent to the native strut as compared to regions closer to the middle of the defect site. Moving from the center outward, the fibrocartilage became more organized and the invaginations with bone became deeper. In each defect site, the tendon was hypercellular and neo-vasculature was present. The insertion site in the native strut (NS; Fig. 13, lower right) did not appear disrupted by the central-third PT repair: tendon fibers in the NS ran through a region of fibrocartilage and anchored into the underlying bone. The fibrocartilage was organized with columnar rows of fibrochondrocytes separated by parallel tendon fibers.

FS FS B B

B T

B FS MFC/FC

Figure 13. Tibial insertion of a PTA+BA repairs, PTA repair and native struts (NS) at 12 weeks. PTA+BAs with relatively low biologic activity (PTA+BA-LA), PTA+BAs with relatively high biologic activity (PTA+BA-HA), PTA, and native strut (tidemark is drawn in for emphasis). A zonal insertion site was not regenerated by either repair treatment (PTA+BA-LA or PTA+BA- HA) or the PTA. H&E stained images are labeled to reflect: B, bone; T, tendon; FS, fibrous scar; MFC/FC, mineralized fibrocartilage/fibrocartilage zones of the insertion site, respectively; areas resembling neo-vasculature are circled. Scale bar = 200µm.

Discussion Two-dimensional cultures of human MSCs (hMSCs) have demonstrated that cellular shape regulates lineage commitment143. Specifically, hMSCs only differentiated into osteoblasts when cultured at a low density and given space to spread out. It is thought that the role of cellular shape on lineage commitment is mediated at least partially by the actomyosin cytoskeleton143.

The need for undifferentiated MSCs to have space in order to commit to the osteogenic lineage may explain why MSC-collagen gel BAs with a lower cellular density (100K cells/ml) had higher ALP activity than BAs with a higher cellular density (600K cells/ml). Additionally, the higher ALP activity in BAs cultured in the glass boat may have contributed to the enhanced tendon-to-bone integration seen in PTA+BA-HA repairs relative to PTA+BA-LA repairs at 12 weeks post-surgery. Unfortunately, in this repair model, enhanced osteotendinous integration did not translate to superior repair tissue biomechanics.

In contrast to promoting MSCs differentiation into osteoblasts, when MSCs are utilized for cartilage tissue engineering, the cells are often formed into high density pellets141 or seeded onto scaffolds at a high density192. It appears as if the high density culture is required to maintain sufficient cell-cell interactions.141, 192 The interaction between cells is important for chondrogenic differentiation and maintenance of the chondrogenic phenotype in vitro.141 For the MSC-collagen gel BAs studied here, an upregulation of collagen type (col2) was more associated with cell culture system than initial cellular density. If the glass boat culture system allows for a higher degree of contraction, as compared to the silicone dish, it could be that col2 expression is regulated by some mechanism of contraction in vitro and/or requires cell-mediated matrix compaction to achieve a higher degree of cellular density than provided initially. The process of cell-mediated contraction and increasing cellular density may not be mutually exclusive in

regulating col2 expression. Further study is needed to assess contraction of BAs cultured in a glass boat versus a silicone dish.

In this study, neither treatment group increased repair tissue biomechanical properties relative to PTA at 12 weeks post-surgery. It is possible that using BAs at the tendon-bone interface limits rather than promotes osteotendinous integration. This may be due to a variety of factors including: 1) the BA reduces the amount of compressive force between the PTA and bone defect thus inhibiting integration of the tissues, 2) the BA still does not possess the correct biological cues to promote integration in the environment of the interface, and 3) the BAs did not remain in place. However, the results of this study agree with those of Thomopoulos et al193 who demonstrated that using a fibrin clot at the bone-tendon interface in a rat rotator cuff injury did not affect repair tissue biomechanical properties except at 3 weeks where there was a decrease in material properties. Therefore, placing augmentations in between the tendon and bone may not be a viable option for promoting osteotendinous integration. However, additional studies by our laboratory have demonstrated that osteotendinous integration improves with recovery time and therefore it is possible that the augmented repairs simply needed more time to mature.

There are several limitations to consider in this study. First, as mentioned above, we cannot be sure that the BAs remained in place after being implanted. BAs created with

CellTrackerTM CM-DiI (Molecular Probes, Invitrogen) labeled cells were implanted for both

PTA+BA-LA and PTA+BA-HA treatment groups, but the dye was not visible 12 weeks after surgery. Future studies will examine a lentiviral propagation protocol to induce green fluorescent protein (GFP) transduction of our MSCs. Transduction with GFP will provide a more sustainable cell marker for ascertaining the presence or absence of the implanted cells. Second, the gene expression data was highly variably (high standard deviations for all genes of interest). Although

three cell lines were tested, a fourth line (n = 4) will be added to help deduce outliers and/or reduce variability. Additionally, preliminary data (not shown) suggests that cell lines may behave differently in monolayer culture (at passage 2). Variability between cell lines in culture may contribute to the variability of the BAs. Third, gene expression was only analyzed for two structural proteins and two transcription factors. As we gain a more in deepth understanding of how tendons and tendon-to-bone insertion sites develop, and as more genetic tools become available for the rabbit, BAs need to be tested for more genes of interest. Fourth, sample size for in vivo testing was low. However, the coefficient of variation ranges from 19 to 32% for structural properties and 14 to 38% for material properties. These values are reasonable for biologic samples. It is unlikely that adding additional samples would produce significant differences between the treatment groups. Fifth, correlations have yet to be established between

1) MSC-collagen gel BA contraction and mRNA expression levels for both osteo- and chondrogenic genes of interest and/or 2) MSC-collagen gel BA mRNA expression in vitro and repair tissue biomechanical properties in vivo. Further study is needed to fully understand the impact of cellular pre-conditioning in vitro on repair outcomes in vivo.

The objective of the study was to determine if implanting MSC-collagen gel BAs at the

PTA insertion sites would improve tendon integration with bone. We hypothesized that implanting a BA at both the patellar and tibial insertions would improve PTA repair tissue biomechanical properties over PTA repair alone at 12 weeks post-surgery. However, this was not the case. Although PTA+BA-HA repairs appeared better integrated at the insertion sites histologically versus PTA+BA-LA repairs, this did not translate into superior biomechanical properties. It is possible that enhancing the biological component of the wound site may be insufficient. In development, there is no evidence of the insertion site until roughly 1 week

postnatally.25, 26 The muscle contraction and movements which occur after birth are thought to trigger formation of the zonal insertion site.25, 26 Future studies will continue to explore both biological and mechanical augmentation strategies to improve osteotendinous integration of the

PTA.

Chapter 6

Chondroitin-6-Sulfate Incorporation and Mechanical Stimulation Increase MSC-Collagen Sponge Construct Stiffness* Kirsten R.C. Kinneberg1, Victor S. Nirmalanandhan1,2, Natalia Juncosa-Melvin1,3, Heather M. Powell4,5, Steven T. Boyce4,6, Jason T. Shearn1, David L. Butler1

1Tissue Engineering and Biomechanics Laboratories, Department of Biomedical Engineering, University of Cincinnati, Cincinnati, Ohio; 2Office of Therapeutics, Discovery and Development, KU Medical Center, Kansas City, Kansas; 3Surgical Energetics, Inc., Cincinnati, Ohio; 4Department of Research, Engineered Skin Laboratory, Shriners Hospitals for Children, Cincinnati, Ohio; 5Department of Materials Science and Engineering, Department of Biomedical Engineering, The Ohio State University, Columbus, Ohio; 6Department of Surgery, University of Cincinnati, Cincinnati, Ohio *Published in the Journal of Orthopaedic Research, vol. 28, pp. 1092-99, 2010.

Abstract Using functional tissue engineering principles, our laboratory has produced tendon repair tissue which matches the normal patellar tendon force-displacement curve up to 32% of failure.

This repair tissue will need to withstand more strenuous activities, which can reach or even exceed 40% of failure force. To improve the linear stiffness of our tissue engineered constructs

(TECs) and tissue engineered repairs, our lab is incorporating the glycosaminoglycan chondroitin-6-sulfate (C6S) into a type I collagen scaffold. In this study, we examined the effect of C6S incorporation and mechanical stimulation cycle number on linear stiffness and mRNA expression (collagen types I and III, decorin and fibronectin) for mesenchymal stem cell (MSC)- collagen sponge TECs. The TECs were fabricated by inoculating MSCs at a density of 0.14x106 cells/construct onto pre-cut scaffolds. Primarily type I collagen scaffold materials, with or without C6S, were cultured using mechanical stimulation with three different cycle numbers (0,

100, or 3000 cycles/day). After two weeks in culture, TECs were evaluated for linear stiffness and mRNA expression. C6S incorporation and cycle number each played an important role in

gene expression, but only the interaction of C6S incorporation and cycle number produced a benefit for TEC linear stiffness.

Introduction Tendon and ligament injuries remain the most common and significant musculoskeletal injuries. Each year, over 16 million patients in the US present with soft connective tissue injuries to tendon, ligament, and capsular structures.88 The knee accounts for roughly 23% of activity- related injuries.107 Left untreated, many of these injuries result in significant dysfunction and disability for the patient.81, 94, 107

Repair outcome after tendon and ligament injury varies depending on the type of treatment and the extent of injury. Direct repair is limited by the intrinsic healing capacity of the tissue and the extent of tissue disruption.1 With poor healing and extensive damage, surgeons may use a graft to replace the tissue. Autografts are limited by availability and impaired recovery due to harvest site morbidity and pain.94 Allografts are also used, but suffer from high cost, limited availability, potential disease transmission, and immune rejection.944 Overall, grafts can lose strength over time and fail to fully incorporate into bone.81, 94

Frequent injuries and the challenges of traditional repair have led some researchers to consider novel treatments such as tissue engineered constructs. Tissue engineered constructs

(TECs) are designed to aid in natural tissue regeneration or replacement and eventually degrade.

TECs are commonly composed of a biodegradable polymeric scaffold and cells. Synthetic polymeric scaffolds can be processed into unique 3D geometries, possess relatively good mechanical strength, and have a controllable degradation rate. However, without surface modification, these scaffolds do not support extensive cellular adhesion, infusion and/or proliferation.194-196 By contrast, natural polymeric scaffolds, such as type I collagen, are highly

biocompatible and can support cell adhesion and proliferation. Unfortunately, these scaffolds exhibit lower mechanical strength than synthetic polymeric scaffolds.196 Physical cross-linking can increase collagen stiffness, but can also make the scaffold brittle and disturb surface markers for cell adhesion and migration, thus limiting tissue ingrowth and remodeling. 197

Using the principles of functional tissue engineering,126, 198 our laboratory is designing more effective collagen-based TECs for tendon repair.80, 81, 130, 140, 146 We first established the functional range of normal tendon loading for activities of daily living (ADLs) by recording in vivo forces in numerous tendons in the rabbit82-84 and goat models85. Peak in vivo forces in the goat patellar tendon (PT) reached 32% of normal PT failure force.85 While our current tissue engineered repairs, which use mechanically preconditioned mesenchymal stem cell (MSC)- collagen sponge TECs, can sustain this load81, tendons can function at up to 40% of failure during more strenuous activities.198 To accommodate these activities, our goal is to produce repair tissue that matches the normal PT failure curve up to 40% of normal failure loads.

One strategy to improve TEC linear stiffness and tissue engineered repair is to incorporate the glycosaminoglycan (GAG) chondroitin-6-sulfate (C6S) into the collagen scaffold. Adding C6S into our collagen scaffold can potentially: 1) improve sponge biomechanics150, 151 and 2) create a more homogenous scaffold by “linking” discontinuous fibrils57, 156. While not the predominant GAG in tensile-load bearing tendons57, C6S does bind to decorin45, which is essential for proper collagen fibrillogenesis in tendon.65, 66 C6S incorporation thus has the potential to positively regulate type I collagen fibrillogenesis.45

Our study objectives were to determine the individual and interactive effects of incorporating C6S and mechanical stimulation on in vitro linear stiffness and mRNA expression

(collagen type I, collagen type III, decorin and fibronectin) of MSC-collagen sponge TECs. We

hypothesized that: 1) Under static culture conditions, incorporating C6S will increase linear stiffness and gene expression; 2) Mechanical stimulation of TECs without C6S will improve linear stiffness and gene expression; and 3) Combining C6S and mechanical stimulation will synergistically increase both TEC linear stiffness and gene expression levels.

Materials and Methods Experimental Design A collagen sponge scaffold fabricated with and without C6S (COL-C6S and COL, respectively; Engineered Skin Lab, Shriners Hospitals for Children, Cincinnati, OH) was evaluated. COL and COL-C6S scaffolds were analyzed for average pore diameter (4 per sample, n=5 per group) using scanning electron microscopy (SEM) and for relative crosslink density (n =

3 per scaffold) using differential scanning calorimetry (DSC). MSC-collagen sponge TECs were created using banked cell lines harvested from ten (n=10) skeletally mature, female New Zealand

White rabbits. MSCs were sub-cultured to passage two (P2) using previously described techniques.80, 140 For each scaffold type, COL and COL-C6S, three treatment levels of mechanical stimulation were tested: static culture, mechanically stimulated with 100 cycles/day and mechanically stimulated with 3000 cycles/day. TECs were evaluated for biomechanics and mRNA expression (collagen types I and III, decorin, and fibronectin). For each test condition, ten TECs (one per cell line) were dedicated to each response measure, biomechanics and gene expression (Table 8).

Table 8. Experimental Design Mechanical Response Measure Scaffold Stimulation Biomechanics Biochemistry COL Static n = 9 n= 10 100 cycles/day n= 10 n= 10 3000 cycles/day n= 10 n= 10 COL-C6S Static n = 9 n= 10 100 cycles/day n= 10 n= 10 3000 cycles/day n= 10 n= 10

Collagen Scaffold Fabrication Collagen sponge scaffolds, with and without chondroitin-6-sulfate, were fabricated at the

Engineered Skin Lab.199 Briefly, comminuted bovine collagen (Kensey-Nash) was solubilized in acetic acid (0.55% wt./vol.) and homogenized by rapid mixing at 5200 rpm. C6S (0.05% wt./vol.) was co-precipitated with the collagen solution through the slow addition of a C6S-acetic acid solution during homogenization to ensure even dispersion of the C6S.199 The collagen-C6S mixture was injected into custom designed casting frames and frozen by submersion in a 95%

EtOH bath. The frozen sheet of collagen-C6S was then lyophilized and dehydrothermally cross- linked at 140ºC for 24 hours. Collagen sponges (COL; 0.6%wt./vol.) were fabricated, as described, without the addition of C6S.

Scanning Electron Microscopy (SEM) Collagen scaffold morphology was examined by scanning electron microscopy (FEI

Sirion), and average pore diameter was determined using ImageJ software. Samples were collected from dry collagen sponges, sputter coated with gold-palladium and imaged in secondary electron mode (5 kV). From the images (4 per sample, 5 samples per group), the diameter of at least 25 pores from each scaffold type was measured.

Differential Scanning Calorimetry (DSC)

As a relative measure of crosslink density, peak denaturation temperature (Tg) of COL and COL-C6S scaffolds was quantified (n = 3 per scaffold). Samples (5-10 mg in dry wt) were sealed in an aluminum cup. An empty cup served as a reference. DSC was performed from 25 to

150°C at a ramp rate of 10°C/min (TA Instruments DSC 2920). Scaffolds formed with acid- soluble collagen served as a negative control.

Tissue Engineered Construct Preparation COL and COL-C6S constructs were prepared as previously described.81 COL and COL-

C6S sponges were cut to fit in wells of a silicone dish. Two 4mm holes were created to allow the scaffold to be secured over the posts in each well. Scaffolds were soaked in 70% ethanol for 24 hours, rinsed with PBS (Gibco) and MSC growth media, and placed in the silicone dishes. MSC were inoculated (0.14x106 cells) on the scaffold in 0.4ml of MSC growth media. All TECs were incubated for two weeks and fed three times weekly (ADV-DMEM, 5% L-ascorbic acid 2- phosphate, 5% FBS, 1% GlutaMAX ™, 1% AB/AM).

Mechanical Stimulation After two days of static culture, TECs allocated for mechanical stimulation were placed into our pneumatic system. 80, 81, 130, 140, 146 Static culture TECs remained in the incubator. TECs were stimulated to a peak strain of 2.4%, at 1Hz, for eight hours a day with either 100 or 3000 cycles/day. 82-85 After two weeks in culture, static and stimulated (2 days of acclimation, 12 days of stimulation) TECs were prepared for biomechanical testing (n=10) or gene expression analysis

(n=10). Biomechanical testing samples were stored at -80ºC; gene expression analysis samples were treated with RNAlater (QIAGEN Inc; Valencia, CA) for 6 hours, snap frozen in liquid

100

nitrogen and placed in a -80°C freezer to prevent RNA degradation.

Biomechanical Evaluation of Constructs TECs were removed from the freezer, allowed to thaw to room temperature, and re- hydrated in PBS prior to testing. The width and thickness of each construct were measured using calipers. TEC ends were secured into testing grips and loaded into the testing system

(TestResources, Inc.; Shakopee, MN). Constructs were failed in tension at a strain rate of

10%/sec. Force-elongation plots were used to determine linear stiffness.

Biochemical Evaluation of Constructs RNA was extracted using RNeasy mini kit (QIAGEN Inc.).140 First-strand cDNA was generated using a conventional reverse transcriptase reaction (MuLV reverse transcriptase,

Applied Biosystems; Foster City, CA). Using rabbit specific primers (collagen type I, collagen type III, decorin, fibronectin and glyceraldehyde-3-phosphate dehydrogenase (GAPDH)200, 201) reverse-transcribed RNA was amplified using conventional polymerase chain reaction (PCR).

Products were verified using electrophoresis and SYBR safe DNA gel stain (Invitrogen

Molecular Probes; Eugene, OR). Messenger RNA levels were then quantified in duplicate using quantitative real-time PCR (qRT-PCR) and normalized to GAPDH expression. GAPDH was used as the housekeeping gene because previous studies by our laboratory have shown GAPDH levels are not affected by mechanical stimulation.140 Primer sequences used for gene expression analysis were published previously.140

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Statistical Analysis Differences in linear stiffness and gene expression were compared using a one-way

ANOVA with C6S incorporation and cycle number as fixed factors. Data were normal and heteroscedastic. Tamhane‟s analysis was used for post-hoc testing. The significance level was set at p < 0.05. Note: qRT-PCR found no RNA for COL-C6S constructs at 100 cycles/day and no collagen type III mRNA in static COL and COL-C6S constructs. These groups were excluded from the analysis.

Results TEC width and thickness averaged 8.2 ± 0.6 mm and 1.8 ± 0.4 mm, respectively (mean ±

SD). These dimensions were not affected by C6S incorporation or mechanical stimulation. C6S incorporation and mechanical stimulation cycle number produced both independent and interactive effects on stiffness and gene expression (Table 9). Independently, each factor significantly altered mRNA expression of collagen type I (p < 0.001), collagen type III (p <

0.001), decorin (p ≤ .005) and fibronectin (p < 0.001). Interactively, the two factors altered TEC linear stiffness (p = 0.006) and mRNA expression of collagen type I (p < 0.001), decorin (p =

0.001) and fibronectin (p < 0.001). GAPDH expression levels were consistent under static culture conditions (0.001 ± 0.0003 and 0.0009 ± 0.0002, COL and COL-C6S, respectively; mean

± SEM), with 100 cycles/day (0.0011 ± 0.0003 and 0.0012 ± 0.0003) and 3000 cycles/day of mechanical stimulation (0.001 ± 0.0003 and 0.003 ± 0.0003).

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Effect of C6S Incorporation on Inherent Scaffold Structure and Biomechanics Adding C6S to the collagen sponge left the scaffold structure unchanged (Fig. 14A and

B) with no significant differences in pore diameter (55.4 ± 2.4 um and 53.1 ± 8.9 um, COL and

COL-C6S, respectively; mean ± SD). C6S incorporation did not significantly improve relative crosslink density (83.0 ± 1.4°C and 80.2 ± 1.4°C, COL and COL-C6S, respectively; mean ± SD).

In addition, no significant differences were found in the as-fabricated stiffness of COL and COL-

C6S sponges (0.027 ± 0.0045 N/mm and 0.025 ± 0.003 N/mm, respectively; mean ± SEM).

Figure 14. Scanning electron microscopy of scaffold materials. Incorporation of C6S had no effect on the structure or pore size of COL and COL-C6S scaffolds [(A) and (B), respectively]. However, CD-COL had larger pores with a wider pore size distribution (C).

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Effect of C6S Incorporation on TEC Linear Stiffness and Gene Expression Static culture: C6S incorporation did not alter TEC linear stiffness (Fig. 15) but did significantly affect nearly all mRNA expression levels (Fig. 16 and 17). The addition of C6S significantly increased collagen type I (p = 0.049), decorin (p = 0.006) and fibronectin (p < 0.001) expression.

Neither COL nor COL-C6S constructs expressed collagen type III.

100 cycles/day: C6S incorporation produced a significant increase in TEC linear stiffness (p =

0.016; Fig. 15). However, adding C6S resulted in no detectable RNA expression.

3000 cycles/day: C6S incorporation had no effect on TEC linear stiffness but did significantly alter gene expression patterns (Fig. 16 and 17). Adding C6S significantly increased collagen type

I expression (p = 0.002) but significantly decreased expression levels of collagen type III (p <

0.001), decorin (p = 0.041), and fibronectin (p = 0.009).

Effect of Cycle Number on TEC Linear Stiffness and Gene Expression COL constructs: Cycle number had no effect on linear stiffness but did significantly affect mRNA expression of collagen types I and III, decorin, and fibronectin (Fig. 16 and 17). Collagen type I expression was highest for constructs stimulated with 3000 cycles/day (p < 0.001), followed by statically cultured constructs and those stimulated with 100 cycles/day (p = 0.016).

Collagen type III expression was also highest for constructs stimulated with 3000 cycles/day (p <

0.001). Decorin expression in constructs cultured statically and stimulated with 3000 cycles/day were both higher than those stimulated with 100 cycles/day (p ≤ 0.046) but were not different than each other. Fibronectin expression was highest for constructs cultured statically (p ≤ 0.012) and constructs stimulated with 3000 cycles/day had higher expression than those stimulated with

100 cycles/day (p = 0.001).

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COL-C6S constructs: Increasing cycle number from 100 to 3000 cycles/day significantly decreased TEC linear stiffness (p < 0.001). However, mechanical stimulation with 100 or 3000 cycles/day did not improve construct stiffness above those cultured statically (Fig. 15).

Introducing mechanical stimulation also affected COL-C6S mRNA expression levels (Fig. 16 and 17). Stimulating COL-C6S constructs with 3000 cycles/day increased expression of collagen type I (p = 0.001) but decreased expression of decorin (p = 0.005) and fibronectin (p < 0.001) when compared to static controls. Collagen type III was only expressed by constructs stimulated with 3000 cycles/day.

Interactive Effects of C6S Incorporation and Cycle Number on TEC Linear Stiffness and Gene Expression C6S incorporation and cycle number interacted to significantly affect both construct linear stiffness and mRNA expression levels of collagen type I and fibronectin (Fig. 15-17). The combination of C6S incorporation and stimulation with 100 cycles/day significantly increased

TEC stiffness above COL constructs stimulated with both 100 and 3000 cycles/day (p = 0.016 and 0.002, respectively). Adding C6S in conjunction with 3000 cycles/day of stimulation significantly increased mRNA expression of collagen type I above all treatment groups involving

COL constructs (p ≤ 0.002). Under static culture conditions, the addition of C6S produced the highest fibronectin expression (p < 0.001). Additionally, in combination with 3000 cycles/day of stimulation, the addition of C6S significantly increased fibronectin expression above COL constructs stimulated with 100 cycles/day (p = 0.029) but decreased expression when compared to static COL constructs (p = 0.022).

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Table 9. Biomechanics and Gene Expression Data for MSC-Collagen Sponge TECs [Mean (SEM)] Cultured for Two Weeks Statically and with Mechanical Stimulation. Scaffold Cycle Number Stiffness Collagen Collagen Decorin Fibronectin Material (Cycles/Day) (N/mm) Type I Type III As 0.027 COL - - - - Fabricated (0.005) 0.053 0.022 No 0.026 0.276 0 (0.006) (0.004) Expression (0.006) (0.064) 0.050 0.004 0.024 0.007 4.40e-4 100 (0.004) (0.003) (0.010) (0.002) (9.33e-5) 0.041 0.076 5.09 0.052 0.030 3000 (0.005) (0.005) (0.735) (0.012) (0.005) As 0.025 COL-C6S - - - - Fabricated (0.003) 0.057 0.070 No 0.710 1.94 0 (0.009) (0.022) Expression (0.218) (0.225) 0.080 100 No RNA No RNA No RNA No RNA (0.006) 0.032 0.600 0.022 0.005 0.004 3000 (0.003) (0.086) (0.003) (0.001) (0.001)

Figure 15. Linear stiffness normalized by static control (mean ± SEM). The addition of C6S increased TEC stiffness when constructs were stimulated with 100 cycles/day. However, mechanical stimulation did not significantly improve linear stiffness above static controls. c/d, cycles/day. *p<0.05.

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Figure 16. mRNA expression of collagen types I and III normalized by GAPDH (mean_SEM). The addition of C6S increased collagen type I expression when TECs were cultured statically and stimulated with 3,000 cycles/day. Mechanical stimulation with 3,000 cycles/day increased collagen type I expression. c/d, cycles/day; ‡, no expression; †, no RNA. *p<0.05

Figure 17. mRNA expression of decorin and fibronectin normalized by GAPDH (mean ± SEM). The addition of C6S increased decorin and fibronectin expression when TECs were cultured statically but reduced expression when TECs were stimulated with 3,000 cycles/day. c/d, cycles/day; ‡, no expression; †, no RNA. *p<0.05

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Discussion This study was designed to examine how incorporating chondroitin-6-sulfate (C6S) into the scaffold and mechanical stimulation into the culture period affect linear stiffness and mRNA expression levels of MSC-collagen sponge tissue engineered constructs (TECs). Each factor uniquely affected the biochemical and biomechanical responses of the TEC. Incorporating C6S increased nearly all of the mRNA levels under static culture and increased linear stiffness when

TECs were exposed to 100 cycles/day of mechanical stimulation. The benefits of incorporating

C6S and mechanical stimulation may be attributed to altered cell-matrix and matrix-matrix interactions.

Gene expression levels produced by COL-C6S constructs may be correlated with the nutrient levels available to the cells. Increased mRNA levels of COL-C6S constructs cultured statically could be due to a higher concentration of nutrients reaching the cells. The negatively charged C6S incorporated into the collagen sponge should make the scaffold swell and increase the water content. If media is pulled into the scaffold, the cells may be provided with a greater influx of nutrients, allowing them to be more metabolically active. However, when mechanical stimulation with 3000 cycles/day was introduced, mRNA expression for nearly all genes decreased. This may be attributed to the fact that when TECs are strained during stimulation, the

TEC thickness decreases during each stimulation cycle due to Poisson‟s effect. The media could be forced out of the scaffold and limit the cells‟ ability to use the nutrients. Therefore, the benefit of incorporating C6S appears to be attenuated in the presence of mechanical stimulation.

Cellular adaptation to mechanical loading is not only affected by the loading frequency, or cycle number, but also by cell-matrix interactions.202 The reduced mRNA levels or lack of

RNA for COL and COL-C6S constructs, respectively, when exposed to 100 cycles/day may be due to either reduced cell viability or reduced cell activity. In an attempt to reduce strain or

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because they were being stress shielded by the stiff construct matrix, the cells may have released integrins, and potentially detached.202 On the other hand, when cells are engaged in persistent sub-threshold interactions, they can become insensitive to activation203 and potentially quiescent.

In contrast, stimulation with 3000 cycles/day may have surpassed this sub-threshold level of stimulation and contributed to increased mRNA expression of fibrillar genes. However, if matrix deposition occurred, it was likely balanced by some level of matrix degradation, as evidenced by unaltered linear stiffness with 3000 cycles/day of stimulation.

The increase in linear stiffness produced by C6S incorporation when TECs were exposed to 100 cycles/day of stimulation may be attributed to altered cell-matrix interactions. If type I collagen and C6S deform differently in response to uniaxial tension, this would alter cell- matrix interactions and potentially cellular adaptation to mechanical stimulation. Additionally,

C6S molecules within the scaffold have the potential to form interfibrillar bonds that act as a link between discontinuous collagen fibrils.57, 156 Collagen-C6S interactions could aid in distributing the mechanical signal throughout the construct, consequently producing a more homogenous, rather than localized, response to the stimulus. These matrix-matrix interactions may contribute to the cell-mediated effects discussed above.

Our current results disagree with several of the findings from a previous study in our laboratory which demonstrated that mechanical stimulation (2.4% strain, 8 hours/day for 12 days) of MSC-collagen sponge TECs improved linear stiffness and mRNA expression of collagen types I and III.80, 81, 130, 140, 146 Using a commercially-derived collagen sponge (CD-COL;

Kensey-Nash, Exton, Pa), the TECs showed orders of magnitude higher mRNA expression levels than the current TECs created using COL and COL-C6S sponges. To help understand these differences, we compared the structure of CD-COL sponges with those of COL and COL-

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C6S sponges. Scanning electron microscopy revealed that, despite using the same source of raw material, the structure of the CD-COL sponge was qualitatively and quantitatively different than that of COL and COL-C6S (Fig. 14). CD-COL sponges were comprised of thick reticulations of collagen with larger pores and a wide pore size distribution while both the COL and COL-C6S sponges contained thin collagen reticulations with smaller and more uniform pores (Fig.14).

Additionally, the average pore diameter within the CD-COL sponge (142.5±16.1um; mean±SD) was approximately three times that of the COL and COL-C6S groups (55.4±2.4um and

53.1±8.9um, respectively). However, peak denaturation temperature, a measure of relative crosslink density, was not significantly altered by processing the CD-COL (80.4±4.5°C; mean±SD), COL (83.0±1.4°C) and COL-C6S sponges (80.2±1.4°C). These results suggest that other factors, like pore size and pore size distribution, may be important to control in scaffold materials. A study is currently underway to identify the optimal pore size for rabbit MSCs on a collagen-C6S sponge to ensure sufficient linear stiffness and enhanced expression of relevant genes important in tendon repair.

The influence of MCS-collagen sponge TEC mRNA expression levels at the time of surgery on repair tissue biomechanics remains unclear. Collagen types I and III are both important in tendon healing with the ratio of collagen type III to collagen type I increasing early in tendon healing before eventually decreasing during the remodeling phase.204 Decorin, the predominant proteoglycan in tensile load-bearing tendon, mediates type I collagen fibrillogenesis and matrix assembly65, 66 and participates in fibril-to-fibril force transfer156. Fibronectin plays a key role in ECM-cell interactions such as adhesion, migration, growth and differentiation52 and also serves to mediate post-translation collagen fibril modifications and assembly53. Despite understanding the function of these factors, we still do not know the magnitude and timing of

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expression needed to dramatically improve repair biomechanics in our PT defect model. A future paired in vitro-in vivo study should help us to better understand how in vitro mRNA expression relates to repair biomechanics in vivo.

There are several limitations to consider. 1) Since TEC aspect ratio was approximately

2:1, end effects could have influenced the mechanical properties. However, low aspect ratio has been shown to have a greater impact on failure properties than the sub-failure linear stiffness which we monitored.205 2) Stimulation-based increases in TEC stiffness have been attributed to increased collagen fibril alignment and ECM deposition by MSCs. However, TEC architecture and cellular contribution to TEC stiffness were not examined in this study. Future studies will implement methods such as SEM, TEM or FTIR to evaluate not only TEC architecture (pore size, pore size distribution, etc.) but also potential inter-fibrillar bonds between collagen and

C6S. We will also compare acellular and cellular TEC stiffness at various time points in the culture period to understand how cells affect TEC biomechanics. 3) Gene expression data was only collected after two weeks in culture. Consequently, we do not yet understand the temporal changes in mRNA expression. We plan to add time points for gene expression in future studies to help understand the development of TECs in culture. 4) Type I collagen production was not evaluated. Since the sponge scaffold is collagen-based, it is difficult to differentiate newly synthesized collagen from matrix collagen. Although immunohistochemical methods are available to stain for type I pro-collagen, the concern is that the collagen may not be integrated into the matrix. Future studies may incorporate radio-labeling to quantify the collagen produced by the MSCs. 5) Cellular viability (living vs. dead) and activity (proliferating vs. quiescent) were not assessed in our constructs. Understanding the viability and activity of our constructs would help clarify whether the lower mRNA expression produced by COL and COL-C6S constructs

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was due to a lack of viable cells or reduced cellular activity.

In conclusion, this study demonstrates that incorporating C6S and applying 100 cycles/day of mechanical stimulation increases the linear stiffness of MSC-collagen sponge

TECs. While C6S incorporation and cycle number each play an important role in gene expression of COL and COL-C6S TECs, their impact is not mutually exclusive. Instead, their interaction was found to produce a benefit for TEC linear stiffness. However, these in vitro results need to be paired with an in vivo study before we can conclude whether these treatments will have a significant impact on tendon healing.

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Chapter 7

Discussion and Conclusions

Chapters 3 through 6 described the progressive findings of a functional tissue engineering approach to improve tendon-to-bone integration and tendon mid-substance healing. The results of Chapter 3 laid the foundation for two parallel paths of research: one examining tissue engineering strategies to improve tendon integration to bone at the insertion site and the other aiming to improve tendon mid-substance healing. The collective findings from these studies are discussed here.

Linear stiffness of tissue engineered constructs (TECs) at the time of surgery has a significant impact on repair outcome but needs to be weighted with other parameters of the

TEC/implant such as matrix density, cellular phenotype and cellular metabolic state/activity level. The results presented in this dissertation demonstrate that a soft tissue patellar tendon autograft (PTA) with linear stiffness matching that of native tendon did not perform as expected according to the previously established correlation between TEC stiffness and repair tissue stiffness81, 130. Taking construct stiffness out of the equation, key differences between a mesenchymal stem cell (MSC)-collagen sponge TEC and the soft tissue PTA include matrix density and cellular activity levels.

The porous structure and MSC components of the TEC may prove to be key advantages over the PTA. The results presented in this dissertation lead us to propose that the sponge matrix allowed for more rapid infiltration of cells because, as compared to the PTA, it required less

MMP/catabolic breakdown to make space for cellular (and potentially vasculature and neuronal) infiltration. The process of cellular infiltration has been shown in ex vivo studies to significantly

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reduce the modulus of patellar tendon (PT) fascicles.206 Thus, as the PTA matrix is being digested chemically, it is likely that the matrix is also being disrupted by cellular infiltration. In contrast, we do not yet understand how cellular infiltration affects the biomechanical properties of our collagen sponge scaffolds. With regard to cellular activity, the primary cell type of the

TEC is the MSC while the primary cell type of the PTA is likely the quiescent tenocyte. The results of Aim 1, along with previous studies by our laboratory demonstrating a biomechanical benefit of cell-based repairs over cell-free repairs, indicate the presence of an initial cell population improves repair outcome. However, the results of Aim 2 contradict this for PTA repair. In Aim 2, the presence of cells (the native cell population) produced no biomechanical or histological benefit for PTA repair. We have now seen results for the PTA contradict two established correlations for tissue engineered repairs in our central-third model of tendon healing. This begs the question: Do MSC-collagen sponge TECs and PTAs heal by different mechanisms? And, if so, how do we/can we manipulate the PTA to heal as well as the MSC- collagen sponge TECs?

While ex vivo studies could be used to understand how cellular infiltration affects collagen sponge biomechanical properties and how cellular activity levels change with time for each implant, it is unlikely that the ex vivo condition will mimic the in vivo environment. A better test to understand the effects of matrix density on cellular infiltration and healing might be to assay both PTA and MSC-collagen sponge repairs for both catabolic (i.e. expression levels and/or gelatin zymography of MMPs, histological presence of macrophages, etc.) and anabolic markers (i.e. expression level of collagen type I, histological evaluation of healing rate, etc.) of healing at early time points. If the MSC-collagen sponge TEC does undergo a more rapid transition from inflammation to remodeling, understanding the biological cues/pathways

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involved in that process will help enable us to promote the cues/pathways of interest in future applications. However, if the MSC-collagen sponge TEC and the PTA undergo similar rates of healing and incorporation, matrix density may not be an influential TEC parameter. Either way, the knowledge gained would help influence the design of future TECs. Additionally, to assess cellular activity and wound healing in the defect site over time, tissue sections could be histologically examined using a traditional H&E stain, Ki-67 stain for proliferating cells207, and a variety of stains for vasculature (such as von Willebrand factor, laminin, and collagen type IV)208 and pericytes/myofibroblasts (such as alpha-smooth muscle actin)208.

Understanding mechanisms of repair is critical to the success of tissue engineered repairs because we cannot change the healing response, we can only work to enhance it. In addition to the methods discussed within this dissertation, two approaches worthy of further consideration are supplementation with growth factors and modification of the repair tissue mechanical environment in vivo. It has been proposed that biological cues may dominate fetal development while biomechanical factors may dictate postnatal development and maturity.25 However, mechanisms of development may correlate to mechanisms of repair.209

Biological cues regulated by cell-signaling molecules, such as members of the transforming growth factor-beta (TGF-β) superfamily, include cellular proliferation, differentiation, chemotaxis as well as matrix synthesis/degradation. Of particular interest is TGF-

210, 211 β3. TGF-β3 is one of three isoforms of TGF-β and has been implicated in the “scarless”

211, 212 healing observed in the fetal and neonatal environment. In contrast, TGF-β1 and TGF-β2, the other two isoforms of TGF-β, have been associated with scar tissue formation in adult

212 healing. In the adult rat, either exogenous neutralization of TGF-β1 and TGF-β2 or exogenous

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212 administration of TGF-β3 reduced scar formation in healing cutaneous wounds. Based on the success demonstrated in the adult rat model, treatment with TGF-β3 has been applied to tendon- to-bone healing in rotator cuff repair where scar formation at the tendon to bone interface is a

210, 211 limiting factor in repair. However, the results are varied. In one study, TGF-β3 was applied with a calcium-phosphate (Ca-P) paste to the tendon-bone interface for rotator cuff repair in the

210 rat model. At 4 weeks, augmentation with TGF-β3 increased load to failure when compared to the Ca-P paste alone, and at 2 weeks, had produced a larger volume of new bone than non- treated controls. However, it was proposed that augmenting with TGF-β3 would reduce scar tissue formation in healing but the authors made no mention of this in their results.210 In a second study of adult rat rotator cuff repair, TGF-β3 was applied using a heparin/fibrin-based delivery

211 system. Contrary to their hypothesis, the authors found that augmentation with TGF-β3 actually increased scar formation.211 Despite this disappointing result, they also found that TGF-

β3 accelerated all phases of the healing process (inflammation, proliferation, and remodeling) and led to increased structural and mechanical properties of the healing insertions.211 The authors raise an excellent point211 that we must also keep in mind as we advance our tissue engineering approaches to tendon healing and insertion site formation: The action of growth factors and other biological augmentations may be quite different in the milieu of adult healing than they are prenatally. The discrepancy in behavior could be due to a variety of factors including the cell types and growth factor receptors present in each tissue and the pathways available for

210, 211 activation.

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Biologic enhancement at the insertion site may be insufficient. Re-establishing the mechanical loading environment may be essential for insertion site formation. Perhaps the most compelling evidence for the role of mechanical forces in tendon development and healing, as it relates to this dissertation, is the fact that the fibrocartilagenous region between tendon and bone does not appear until roughly 1 week postnatally.25, 26 As muscle forces and motion develop, so do the tensile stresses that produce proteins associated with tendon (i.e. collagen type

I) and the compressive stresses that produce proteins associated with cartilage (i.e. collagen type

II). However, when these forces are removed, fibrocartilage undergoes drastic remodeling.213

When tendons wrap around a bony pulley, their composition is altered in response to the compressive force against the bone.213 These “wrap-around” tendons contain a region with chondrocytes and elevated GAG content (similar to cartilage).213 However, when this “wrap- around” tendon is translocated and made to run directly from muscle to bone, the cartilage cells disappear and the GAG content is reduced.213 Additionally, the capacity of the “wrap-around‟ tendon to restore its native composition of chondrocytes and GAG diminishes the longer it remains translocated.213 These biochemical changes in response to mechanical loading may represent similar changes occurring in our soft tissue PTA model. When the excised PTA tissue is secured into the defect site, the PTA tissue is only reattached to the native struts. The footprint of the tendon is not restored at either the patella or tibia and the mechanical stresses between the bones and tendon are likely lost until integration occurs. The longer it takes to restore both the tendon-to-bone insertion site and the mechanical loading environment of native fibrocartilage, the longer it may take for the fibrocartilage to regenerate. Therefore, in addition to augmentation with biological cues, it may be necessary to restore the mechanical stresses at the insertion site in order regenerate the presence of an organized fibrocartilage region.

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Understanding how our TECs and BAs develop in vitro will enhance our ability to produce superior outcomes in vivo. MSC-collagen gel BAs were designed with developmental biology in mind. When determining the genes of interest, both early and late markers of osteo- and chondrogenesis were chosen because: 1) we do not fully understand how our TECs and BAs develop in culture, and 2) we do not know what profiles will produce a benefit in vivo. However, now that biological tools, such as real time PCR, gene array analysis and laser capture analysis, have become more accessible, an understanding of the spatial and temporal patterns of gene expression is more attainable. Taking advantage of these biological tools will provide a major stepping stone for advancing tissue engineered repairs.

As our understanding of normal patellar tendon development and normal patellar tendon healing progress, we can begin to screen for important similarities and differences between the two. We know that normal development produces the ideal outcome while normal healing does not. By eliminating factors that have similar expression profiles, we can distinguish markers which are unique to each process and isolate those critical to normal development. The factors which differentiate normal development from natural healing can then be screened for in our

MSC-collagen sponge TECs and MSC-collagen gel BAs. Using in vitro screening to limit testing in vivo will not only save time, money and resources, but also help us avoid the slow- progressing, iterative process typically employed by tissue engineers.

As we make progress in the patellar tendon model, we can start to apply these findings to not only larger animal models but also more clinically relevant injury models. While the excised central-third patellar tendon is not a clinically relevant injury, it is a reproducible model of tendon healing. The progress we make in this model can be applied most readily to rotator cuff repair and anterior crucial ligament (ACL) reconstruction. For example, osteotendinous

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integration is a significant problem for rotator cuff healing. The studies described in this dissertation have attempted to improve osteotendinous integration using biologic augmentations

(BAs) between the tendon and bone. However, as described in Chapter 5, these BAs have limitations that can be applied to the clinically relevant situation. If the BAs do in fact act as a barrier for osteotendinous integration in our model of PT healing, it is likely that similar treatments in the rotator cuff will also inhibit integration. Additionally, if our future studies discern that enhanced biology is insufficient to restore osteotendinous integration such that both biological and mechanical factors are needed to restore the zonal insertion site, it is also likely that a similar treatment modality will be required in the clinically relevant situation. For healing of the ACL, studies have demonstrated differences in healing exist between the intra- and extra- articular regions of the bone tunnel.214 Understanding how the extra-articular PT integrates with bone could provide important information to help promote ACL graft integration with bone in the extra-articular region of the bone tunnel. Additionally, if our in vivo studies reveal that the implanted MSCs in our sponge TECs undergo less catabolic/MMP breakdown and indeed provide a biomechanical benefit to PT healing, MSC-collagen sponge TECs could be utilized as

“wrap-around” treatments for ACL repair. These MSC-collagen sponge TECs could help sustain autograft viability and potentially help reduce the catabolic breakdown of autograft tissue in the joint capsule. A further understanding of how the MSC-collagen sponge TEC is incorporated in vivo will aide in our ability to apply this treatment to more clinically relevant injury models.

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Chapter 8

Recommendations for Future Studies

This chapter will address some of the proposed future studies and unanswered research questions described in the previous chapters of this dissertation.

Is there a critical size defect for tendon healing and insertion site formation? The results of Chapter 4 indicate that osteotendinous integration was superior at the medial/lateral edges of the defect site, as compared to the central region. This result brings up the issue of critically sized defects in tendon healing and insertion site formation. Similar to studies that examine critically sized defects in bone, our laboratory should examine critically sized defects in the patellar tendon (PT). If a critical size defect is identified, below which the PT defect can heal to normal biomechanical and histological levels, subsequent studies could explore: 1) how that size relates to the size of the PT used for ACL replacement (perhaps a smaller portion of the PT could/should be used for ACL replacements); 2) if suturing the native struts together, as is done clinically, affects the critical size of the defect both in the mid- substance and at the insertion; and 3) methods to heal critically sized defects (such as the biologic augmentations presented in Chapter 5).

To expand on the idea of exploring methods to heal critically sized defects in the PT, the results presented in Chapter 5 indicate that biologic augmentation/enhancement may be insufficient to promote normal insertion site formation for the patellar tendon autograft (PTA).

Therefore, in addition to biologic augmentation, it may be necessary to restore the mechanical

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stresses at the insertion site in order regenerate the presence of an organized fibrocartilage region

(as discussed in Chapters 5 and 7).

The most important aspect of mechanical loading to artificially restore is the compressive force between tendon and bone. If tendon is compressed against bone, the animal‟s normal gait will restore the tensile loads of the tendon (and potentially promote formation of the insertion site as seen developmentally). Additionally, the compressive load is critically important for maintaining not only the presence of fibrocartilage but also its phenotypic organization and composition.

Internal and external mechanisms could be used to restore a compressive loading environment at the insertion site. Internal mechanisms might include incorporating a suture anchor technique used to clinically repair rotator cuff tendons (while potentially not feasible in the rabbit, this approach would be feasible in the larger sheep model) and/or implanting fixture plates over the top of the tendon. Both of these approaches could be paired with biologic augmentations (BAs) placed either between the PTA and bone or over the top of the tendon as an onlay. External mechanisms to restore compressive loading at the insertion site might include a static casting or a dynamic/cyclic apparatus applied to the healing area (potentially only for a set number of hours/day). When MSCs-agarose TECs were cultured under cyclic compression, mRNA expression of collagen type II increased from day 7 to day 14 and decreased thereafter.149

It is possible that a similar phenomena could occur in vivo. These external mechanisms could each be paired with both an internal suture anchor technique and/or BAs placed in the tendon- bone junction or as an onlay.

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Prior to optimizing the BAs and internal and/or external load restoration mechanisms, we first need to understand: Do the implanted mesenchymal stem cells (MSCs) remain in the defect site and incorporate into the healing tissue? If the implanted cells simply undergo apoptosis once they are implanted, it would seem there is little need for pre-conditioning them in vitro prior to implantation. However, if the cells contribute to the healing process, it becomes even more important to understand not only the mechanisms of normal healing but how these mechanisms diverge from normal development. As described in Chapter 7, by contrasting spatial and temporal gene expression patterns we can identify genes of interest unique to normal development and normal healing. Once we have isolated factors that are involved in the superior process of normal development, we can begin attempts to promote and/or screen for these factors in vitro in future TECs. This study would also help us understand what proteins, besides those studied in this dissertation (collagen types I and II and the transcription factors runx2 and sox9), are important for formation of a zonal tendon-to-bone insertion site.

In a similar fashion to how we can use our understanding of normal development to enhance tissue engineering approaches to tendon healing and insertion site formation, we can use our understanding of a “good” repair as well. For example, the results of Chapter 3 demonstrate that, after 12 weeks of recovery, MSC-collagen sponge tissue engineered constructs (TECs) produced repair tissue biomechanical properties roughly twice those produced by PTA. While there are many factors which distinguish the MSC-collagen sponge TEC from the PTA (such as matrix density and native cell population, all discussed previously), the fact remains that TEC- mediated repair was superior to PTA repair. By understanding how normal tendon develops and heals and how both “good” and “bad” tissue engineered repairs heal, similarities and differences

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in their pathways can be identified and used to improve future TECs for tendon healing and insertions site formation.

If future studies are undertaken to alter gene expression levels of MSC-based augmentations, we need to understand: How can external stimuli be used to alter MSC gene expression in vitro? Data presented in Chapter 5 demonstrates that gene expression tends to increase with time in culture (not a statistically significant conclusion). However, if our augmentations are to be implemented clinically, gene expression levels will need to reach expected levels much faster than 14 days. Potential options for external stimuli include chemical and/or mechanical stimulation. For example, when human bone marrow-derived MSCs were cultured in osteogenic media supplemented with 10% FBS, 0.1 mM dexamethasone, 0.05 mM ascorbic acid-2-phosphate, and 2 mM bglycerophosphate, mRNA expression of both runx2 and collagen type I were up-regulated in a two-stage process with a slight increase at day 1, relatively constant expression until day 7 and a large increase thereafter.215 Additionally, in vitro mechanical stimulation has been reported to induce fibrillar orientation216 and cell alignment217,

218; increase cell proliferation217, 219, DNA synthesis219, and secretion of growth factors (TGF-β, bFGF, and PDGF)220, 221 and collagen217. In order to produce elevated gene expression levels in a timely fashion, chemical and mechanical stimulation may have to be applied simultaneously.

123

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