Behavior of passaged chondrocytes in collagen-glycosaminoglycan scaffolds : effects of cross-linking, mechanical loading, and genetic modification of the scaffold

BEHAVIOR OF PASSAGED CHONDROCYTES IN

COLLAGEN-GLYCOSAMINOGLYCAN SCAFFOLDS: EFFECTS OF

CROSS-LINKING, MECHANICAL LOADING, AND GENETIC

MODIFICATION OF THE SCAFFOLD

by

Cynthia R. Lee

B.S. Bioengineering

University of California, Berkeley, 1997

S.M. Mechanical Engineering

Massachusetts Institute of Technology, 1999

Submitted to the Department of Mechanical Engineering in Partial

Fulfillment of the Requirements for the degree of Doctor of Science

at the

Massachusetts Institute of Technology

February 2002

© Massachusetts Institute of Technology, 2002

Signature of Author

Certified by

Professor of

4/

September 2001

w Mvyon Spector, Thesis Supervisor

/ Senior Lecturer Mechanical Engineering Orthoydic Surgery (B materials), Harvard Medical School

Certified by

MASSAC HUSETTS N OF TECHNOLOGY

MAR

2

9 2002

. Grod nsky, Thesis Supervisor Pr ssor of Electrical, Me ical, and Bioengineering

Ain A. Sonin Chairman, Departmental Committee on Graduate Students

BEHAVIOR OF PASSAGED CHONDROCYTES IN

COLLAGEN-GLYCOSAMINOGLYCAN SCAFFOLDS:

EFFECTS OF CROSS-LINKING, MECHANICAL

LOADING, AND GENETIC MODIFICATION OF THE

SCAFFOLD

by

Cynthia R. Lee

Submitted to the Department of Mechanical Engineering in partial fulfillment of the requirements for the Degree of Doctor of Science, February 2002.

ABSTRACT

Tissue engineering is a promising solution to the problematic healing of cartilage defects. The purpose of this thesis was to establish a foundation for the development of a collagen-glycosaminoglycan (CG) scaffold for articular cartilage tissue engineering by exploring the behavior of passaged chondrocytes in the CG scaffolds under the influence of a variety of environmental factors.

Using in vitro studies, the first two parts of the thesis evaluated the effects of the physical environment on the behavior of adult, passaged chondrocytes in the CG scaffold. Scaffold cross-linking procedures increased cross-link density and scaffold stiffness and increased resistance to cell-mediated degradation and contraction as follows: dehydrothermal treatment (DHT) < ultraviolet irradiation (UV) < gluteraldehyde (GTA)

< carbodiimide (EDAC). EDAC scaffolds also provided for the highest levels of cell

proliferation and protein and GAG synthesis throughout a 4-week culture. Static mechanical compression (0-50% strain) applied to cell-seeded EDAC cross-linked scaffolds decreased rates of protein and GAG synthesis while dynamic compression (3% sine amplitude, 0.1 Hz) increased rates of biosynthesis over a 24-hour period. These results were similar to those of prior studies of loading of intact cartilage explants. Unlike the explant studies, however, dynamic compression failed to increase the accumulation of matrix molecules within the construct compared to unloaded ("free-swelling") controls because of a large increase in the release of newly synthesized macromolecules into the media.

To evaluate the in vivo performance of the chondrocyte-seeded EDAC cross-linked CG scaffold, repair tissue formed 15 weeks after implantation of a 4-week in vitro cultured construct was evaluated. The majority of the repair tissue was hyaline and fibrocartilaginous. However, it displayed decreased levels of type II collagen and GAG staining compared to normal articular cartilage, and had a compressive stiffness that was 20-fold lower than normal.

Finally, in anticipation of future work utilizing gene therapy to improve cartilage repair, the CG scaffolds were modified for direct delivery of genetic material to cells in situ. Scaffold cross-linking and plasmid pH altered the ability of the CG scaffolds to carry plasmid DNA to local cells. EDAC cross-linked scaffolds and scaffolds prepared

with plasmid at a neutral pH bound the lowest amount of DNA but, over two to eight week in vitro culture periods, these scaffolds led to higher levels of gene expression compared to non- or DHT cross-linked scaffolds and plasmid preparations at an acidic

pH (pH 2.5).

Although current knowledge is not sufficient to successfully repair articular cartilage wounds, the understanding of the responsiveness of passaged chondrocytes in

CG scaffolds gained in this thesis can be used to further the development of this system.

In brief, the construct that is recommended for future investigation as an implant for articular cartilage repair is a chondrocyte-seeded, EDAC cross-linked CG scaffold cultured in vitro under dynamic compression prior to implantation.

Thesis Supervisors:

Myron Spector... Senior Lecturer Mechanical Engineering Massachusetts Institute of Technology and Professor of Orthopaedic Surgery (Biomaterials) Harvard Medical School Alan J. Grodzinsky...Professor of Electrical, Mechanical, and Bioengineering Massachusetts Institute of Technology Thesis Committee:

Loannis Yannas ... Professor of Mechanical Engineering and Materials Science Massachusetts Institute of Technology Gordana Vunjak-Novakovic ... Principal Research Scientist Massachusetts Institute of Technology and Adjunct Professsor of Chemical Engineering Tufts University

ACKNOWLEDGEMENTS

I am deeply grateful to my advisors, Professors Myron Spector and Alan J. Grodzinsky, and my thesis committee members, Drs. Vunjak-Novakovic and Yannas for all of their valuable advice and guidance. Professors Spector and Grodzinsky have been great mentors during the past four years. They allowed me the freedom to explore my own ideas, while giving me structure and guidance. Through Professor Spector's Orthopaedic Research Laboratory at Brigham's and Women's Hospital, I gained valuable exposure to the world of orthopaedics, both research and clinical. Through Professor Grodzinsky's Continuum Electromechanics Laboratory at MIT, I remained rooted in engineering (mechanical, biomedical, electrical, aeronautical, chemical, etc...).

To all of the people in the ORL and Al's gang, thank you many times over for everything you taught me - especially Eliot Frank's lessons on the Dynastat; Han-Hwa Hung's rigorous training in lab safety, maintenance, organization, supply ordering and everything in between; Steve Treppo's teachings of biochemistry; Andy Loening's tutorials in enzymes and cell isolation; Moonsoo's words of wisdom with the Incustat; Dr. Hsu's crash courses in knee anatomy, cartilage harvesting and suturing; Sandra Zapatka-Taylor's wealth of knowledge in all aspects of histology; Xiuying Zhang's lessons in Western blotting; John Kisiday's various cell culture and collagen tricks - and did for me - all those animal surgeries (Dr. Hsu); ALL those samples processed embedded, sectioned, and stained (Sandra); ALL those gels for Western blots and autoradiography (Han-Hwa); the troublesome SMA Western blots (Robyn Marty-Roix); and caring for my cultures whilst I traveled (John). Thank you also to Professor Yannas and members of the Fiber and Polymers Laboratory (Lila Chamberlain, Mark Spilker, Toby Freyman, and, in the end, Brendan Harley) for use of the collagen-GAG technology and cell culture space, and for giving me my first "home" when I came to MIT. Thank you Ray Samuel, MD/PhD, for introducing me to gene therapy and bringing me into the project with the "gene-seeded scaffolds" (Chapter 5!). Thank you also to all the students who worked with me as UROPs and helped with various experiments - Wendy Liu, Katherine Oates, Juwell Wu, Julie Watts, Aileen Wu, and Karen Riesenburger.

Thank you also to all of my friends here in Cambridge - Helen and Camille, you made adjusting to life on the east coast so much easier - I guess us California girls won't

melt in the eastern summers or turn to popsicles during the winters after all (tho' it sure felt like it at times). My roommates, Chris(tina), Alan, and Tim suffered through grad school at MIT with me and provided drama as a reminder that there was life beyond the sterile confines of the lab. Of course, I could not have kept my sanity through four plus years at MIT without some fun, whether it be playing field hockey with the Minutewomen (and Minutemen), soccer with the Kickbacks, mountain biking with Helen and Chris, kayaking with Andy, triatholoning with Robin Evans (when she wasn't crashing into one thing or another!), or random adventures with Linda Bragman, Bodo Kurz, and Parth Patwari (yes Al, when you were gone, we did play...).

And, last, but not least, thank you to my family who have always given me the love and support to help me succeed (even when they thought I was majoring in recreation)!

CONTENTS

CHAPTER . GENERAL INTRODUCTION... 10

1.1. STRUCTURE AND FUNCTION OF ARTICULAR CARTILAGE ... 10

1.2. CARTILAGE DEGENERATION ... 11

1.3. SURGICAL TREATMENT OF CARTILAGE WOUNDS ... 12

1.4. TISSUE ENGINEERING OF ARTICULAR CARTILAGE ... 14

1.4.1. General Approach... 14

1.4.2. Review of Previous Work ... 15

1.4.2.1. Scaffolds... 15

1.4.2.2. Growth Factors... 16

1.4 .2.3. C ells... . 16

1.4.2.4. Regulation by physical forces ... 17

1.4.2.5. The Collagen-Glycosaminoglycan System... 18

1.5

CHAPTER 2: EVALUATION OF CROSS-LINKING METHODS FOR COLLAGEN-GLYCOSAMINOGLYCAN SCAFFOLDS ... 21

2.1. INTRODUCTION... 21

2.2. MATERIALS AND METHODS... 22

2.2.1. Collagen-Glycosaminoglycan Scaffold Fabrication... 22

2.2.1.1. Freeze-Drying... 22

2.2.1.2. Cross-Linking Methods... 22

2.2.2. Physical Characterization of Scaffolds ... 23

2.2.2.1. Swelling Ratio Determination ... 23

2.2.2.2. Compression Testing of Scaffolds ... 24

2.2.2.3. Glycosaminoglycan Content of Scaffolds... 24

2.2.3. Cell-Seeded Assays... 25

2.2.3.1. Chondrocyte Isolation and Culture ... 25

2.2.3.2. Cell Seeding and Culture of CG Scaffolds... 25

2.2.3.3. Measurement of Cell-Mediated Contraction...26

2.2.3.4. Radiolabel Incorporation... 26

2.2.3.6. D N A A nalysis ... 26

2.2.3.7. Glycosaminoglycan Content of Cell-Seeded Scaffolds ... 27

2.2.3.8. Immunohistochemistry... 27

2.2.4. Statistical A nalysis ... 27

2

2.3.1. Physical Characterization of Unseeded Scaffolds ... 28

2.3.1.1. Sw elling R atio ... 28

2.3.1.2. Com pressive Stiffness ... 30

2.3.1.3. GAG Content of Scaffolds ... 30

2.3.2. Cell-Seeded Assays... 31

2.3.2.1. Cell-mediated Contraction ... 31

2.3.2.2. D ry W eights ... 33

2.3.2.3. Radiolabel Incorporation... 34

2.3.2.4. D N A A nalysis ... 36

2.3.3. GAG Content of Seeded Scaffolds... 37

2.3.4. Immunohistochemistry ... 39

2

CHAPTER 3: EFFECTS OF MECHANICAL COMPRESSION ON BIOSYNTHETIC ACTIVITY OF PASSAGED CHONDROCYTES IN TYPE II COLLAGEN-GLYCOSAMINOGLYCAN SCAFFOLDS ... 48

3

3.2.

3.2.1. Collagen-Glycosaminoglycan Scaffolds ... 49

3.2.2. Cell Culture and Cell-Seeding of Scaffolds ... 49

3.2.3. Mechanical Compression of Cell-Seeded Scaffolds...49

3.2.3.1. Static Compression - Dose Response ... 49

3.2.3.2. Static Compression - Kinetics ... 50

3.2.3.3. Dynamic Compression...51

3.2.4. Protein and GAG Biosynthesis... 51

3.2.5. Analysis of Macromolecules Released to the Medium... 52

3.2.6. Newly Synthesized Proteoglycan Size and Affinity for Hyaluronic Acid .. 52

3.2.7. Statistical A nalysis ... 53

3

3.3.1. Static Compression - Dose Response ... 53

3.3.2. Static Compression - Kinetics ... 55

3.3.3. Dynamic Compression ... 56

3.3.4. Newly Synthesized Macromolecules Released to the Medium...57

3.3.5. Proteoglycan Analysis... 59

3 .4 . D ISC U SSIO N ... 60

CHAPTER 4: REPAIR OF CANINE CHONDRAL DEFECTS IMPLANTED WITH AUTOLOGOUS CHONDROCYTE-SEEDED TYPE II COLLAGEN SCAFFOLDS ... 65

4.2.

4.2.1. A nim al M odel ... 66

4.2.2. Type II Collagen Scaffolds ... 68

4.2.3. Cell Culture and Preparation of Cell-Seeded Implants... 68

4.2.4. H istom orphom etry... 69

4.2.5. M echanical Testing ... 70

4

4.3.1. Histology and Immunohistochemistry of Cell-Seeded Scaffolds... 72

4.3.2. General Gross and Histological Observations ... 72

4.3.3. Histomorphometric Evaluation of Reparative Tissue ... 74

4.3.4. Mechanical Properties of Repair Tissue... 76

4

CHAPTER 5: FABRICATION OF GENE-SEEDED COLLAGEN-GLYCOSAMINOGLYCAN SCAFFOLDS... 81

5.1. INTRODUCTION ... 81

5.2. MATERIALS AND METHODS ... 82

5.2.1. Preparation of Collagen-GAG Scaffolds ... 82

5.2.2. Addition of Plasmid DNA to Preformed CG scaffolds... 83

5.2.3. Electron Microscopy of the CG Scaffolds ... 83

5.2.4. Plasmid DNA Content of GSCG Scaffolds... 83

5.2.5. Plasmid DNA Leaching Studies ... 84

5.2.6. Transfection of Canine Articular Chondrocytes in Gene-Supplemented CG Scaffolds ... 84

5.2.7. Stability of Chondrocyte Transfection ... 84

5 .3 . R ESULTS... 85

5.3.1. Morphology and Ultrastructure of the GSCG Scaffolds... 85

5.3.2. Loading of Plasmid DNA ... 87

5.3.3. Release K inetics ... 88

5.3.4. In Situ Transfection of Chondrocytes Seeded into GSCG Scaffolds...91

5.3.5. Stability of Chondrocyte Transfection ... 93

5 .4 . D ISCU SSION ... 93

CHAPTER 6: CONCLUSIONS ... 96

CHAPTER 7: LIMITATIONS AND FUTURE WORK...100

REFERENCES ... 104

Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure

LIST OF FIGURES

Anatomy of articular cartilage... 10

Approved clinical procedures for treatment of articular cartilage defects ... 13

A porous collagen-GAG scaffold... 15

Inverse swelling ratio of cross-linked scaffolds... 28

Compressive stiffness of cross-linked scaffolds ... 29

Compressive stiffness vs. inverse swelling ratio... 29

GAG content of unseeded scaffolds... 31

Cell-mediated contraction ... 32

Correlations for normalized cell contraction... 33

Dry weight of seeded scaffolds ... 34

Protein and GAG biosynthesis ... 35

DNA content of scaffolds... 36

GAG content of seeded CG scaffolds ... 38

Type II collagen immunostaining of cell-seeded scaffold ... 39

Proposed cross-linking mechanisms ... 44

Type II collagen western blot... 46

Chambers for compressive loading ... 50

Biosynthetic dose response to static compression... 54

Kinetics of radiolabel incorporation... 55

Effects of compression on macromolecular accumulation ... 56

Total rates of biosynthesis... 58

Size distribution of newly synthesized proteoglycans ... 59

Surgical creation of chondral defects ... 67

Light micrographs of cell-seeded type II collagen scaffolds cultured for four w eeks... . 72

Gross appearance of joint surfaces at necropsy ... 73

Histological sections from the center of repair tissue filling the canine chondral defects 15 weeks after implantation ... 75

Indentation stiffness of repair tissue and articular cartilage...76

Histomorphometric comparison of repair tissue filling defects subjected to different treatm ents ... 78

Electron micrographs of gene-seeded collagen-GAG (GSCG) scaffolds.... 86

Total DNA loaded into scaffolds ... 87

D N A leaching profiles ... 89

D N A release rate ... 90

In situ transfection of chondrocytes in GSCG scaffolds ... 92

Figure 4.3. Figure 4.4. Figure 4.5. Figure 4.6. Figure 5.1. Figure 5.2. Figure 5.3. Figure 5.4. Figure 5.5.

LIST OF TABLES

Table 2-1. Summary of cross-linking treatments ... 23

Table 5-1. Plasm id DNA release rates ... 91

LIST OF EQUATIONS

Equation 2-1. Calculation of swelling ratio from wet and dry weights ... 23

CHAPTER 1: GENERAL INTRODUCTION

Normal joint function depends on the low joint friction and absorption and transmission of loads afforded by healthy articular cartilage. For over two and a half centuries it has been known that articular cartilage has limited natural healing [Hunter, 1743] and that injury to articular cartilage can lead to acute and chronic pain. Despite this knowledge, relatively little progress has been made towards improving cartilage healing. Recently, the development and advances of tissue engineering - the ex vivo growth of tissue - have given renewed hope for articular cartilage repair. To take full advantage of any tissue engineering scheme, a thorough understanding of the role of various environmental factors in the natural and engineered development of articular cartilage is necessary. The focus of this thesis is the development of a tissue engineering system employing a collagen-glycosaminoglycan scaffold with passaged adult chondrocytes.

1.1. STRUCTURE AND FUNCTION OF ARTICULAR CARTILAGE

Articular cartilage is aneural, glassy-like (hyaline)

a highly specialized connective tissue. It is the avascular, connective tissue lining the ends of long bones (Figure 1.1).

proteoglycan

cartilage

Type 11 collagen n y chondrocyte

Figure 1.1. Anatomy of articular cartilage. Articular cartilage lines the ends of long bones in

diarthroidal joints (i.e. hip, knee, elbow, knuckles). The solid components of the extracellular matrix of articular cartilage are primarily rope-like type II collagen fibers and negatively-charged proteoglycans. The chondrocyte is the cell responsible for the synthesis and organization of the matrix.

Articular cartilage serves as a natural bearing material that absorbs and transmits loads across diarthroidal joints (i.e. knee, hip, knuckles, etc.). Healthy articular cartilage supports compressive, tensile, and shear loads while providing low friction between articulating joint surfaces.

The mechanical properties of articular cartilage arise from the seemingly simple, though unique structure and composition of the tissue. Articular cartilage is primarily water (65-75% wet weight), with an arcuate collagen network (50-90% dry weight as type II collagen with smaller amounts of type VI, IX, X and XI collagens) and highly negatively charged proteoglycans (5-10% dry weight) [Muir, 1995] making up the solid components of the matrix. The collagen fibers provide tensile and shear strength while the proteoglycans and interstitial fluid impart compressive strength. Even small changes in the composition or organization of the extracellular matrix can profoundly alter the mechanical properties of articular cartilage. Thus, a crucial aspect in the natural and engineered regeneration of articular cartilage is the realization of the precise cartilage matrix substance and architecture.

1.2. CARTILAGE DEGENERATION

The specialized articular cartilage matrix is maintained by chondrocytes. Mature chondrocytes are sparsely distributed throughout the tissue (approximately 10,000 cells/mm3 in adult human cartilage [Muir, 1995]) and have relatively low mitotic and biosynthetic activity. As a result, once damaged, articular cartilage has limited capacity for repair. In particular, even relatively small changes to the integrity of the cartilage matrix will alter its mechanical properties, making damaged cartilage prone to more severe degradation. As the extracellular matrix of cartilage degenerates, bone wears on bone, resulting in pain, deformity and loss of joint motion. This condition, known as arthritis, is the leading cause of disability in the United States with some 43 million Americans suffering from arthritis in 1998, at a cost to the US economy of $65 million in medical care and lost wages [AAOS, 2000]. By 2020, it is estimated that as many as 60 million people will suffer from arthritis [AAOS, 2000].

Osteoarthritis, or "wear and tear" degeneration, is the most common form of arthritis, affecting 30 million Americans in 1998 [AAOS, 2000]. Although the etiology

of osteoarthritis is largely unknown, it is generally accepted that unrepaired focal lesions that may initially be due to trauma can extend themselves and predispose the joint to more wide-spread degeneration due to excess friction and uneven joint loading [Grande et al., 1989; Minas and Nehrer, 1997]. The avascular and aneural nature of the cartilage matrix, coupled with the limited mobility, proliferative, and biosynthetic activity of the mature chondrocyte severely impairs the healing of defects limited to the cartilage ("partial thickness" or "chondral"). While these partial thickness lesions do not heal [Ghandially et al., 1977; Mankin, 1982], wounds penetrating the cartilage and underlying bone ("full thickness" or "osteochondral") are more likely to fill with repair tissue. This repair tissue, however, lacks the mechanical integrity of normal articular cartilage and tends to deteriorate over time [Shapiro et al., 1993].

1.3. SURGICAL TREATMENT OF CARTILAGE WOUNDS

Left untreated, isolated cartilage wounds may eventually progress to severe cartilage degeneration and joint pain. The only relief for severely arthritic joints is total joint replacement (TJR). While TJR is successful for most patients, problems exist, particularly for young, active patients. Early treatment of cartilage lesions to restore the integrity of the joint surface may prevent osteoarthritic degeneration of the joint and future TJR. In an effort to relieve pain and preempt TJR, it has been estimated that each year 1 million Americans have surgery to treat articular cartilage injuries [AAOS, 2000].

Several surgical techniques aim to improve cartilage healing [Minas and Nehrer,

1997; Newman, 1998] by promoting filling of the cartilage wound. Since cartilage

lesions which penetrate the subchondral bone can fill with repair tissue, techniques such as microfracture (Figure 1.2a) and abrasion arthroplasty have been developed to draw blood and multi-potent stem cells from the underlying subchondral bone into the wound to promote filling of the defect. These techniques increase the amount of repair tissue, providing at least temporary pain relief. The inferior mechanical properties of the repair tissue, however, predispose the repair tissue and surrounding cartilage to degeneration under the demanding joint loads [Minas and Nehrer, 1997; Temenoff and Mikos, 2000].

Transplantation procedures such as autologous cell implantation (ACI), osteochondral transplantation (OCT) and mosaicplasty aim to fill the defect with articular

cartilage synthesized in situ by transplanted chondrocytes (ACI) or with mature cartilage translocated from "non-weight-bearing" locations (OCT and mosaicplasty). In ACI (Figure 1.2b), chondrocytes are harvested from "healthy," "non-weight-bearing" cartilage, expanded in monolayer culture, and reinjected into the cartilage wound under a periosteal flap. This procedure has been shown to alleviate pain [Brittberg et al., 1994; Minas and Nehrer, 1997], but the retention of the cells within the defect [Breinan et al.,

1998], the degeneration induced by suturing of the periosteal flap [Breinan et al., 2000;

Breinan et al., 1997], the quality and durability of the repair tissue synthesized by the cultured cells [Breinan et al., 2001], and the integration of the repair tissue with the host tissue [Breinan et al., 1998; Brittberg et al., 1994] are problematic. With OCT [Outerbridge, 1995 #481] and mosaicplasty [Matsusue et al., 1993] (Figure 1.2c), an

OKA Stone M.D.

a

b

Figure 1.2. Approved clinical procedures for treatment of articular cartilage defects. (a) In

microfracture, the surgeon creates tiny holes in the underlying subchondral bone to stimulate bleeding into the defect (image downloaded from www.stoneclinic.com). (b) In autologous chondrocyte implantation (ACI), chondrocytes are harvested from a cartilage biopsy, expanded in monolayer culture, resuspended, and injected under a periosteal flap (image downloaded from

www.drmendbone.com). (c) In mosaicplasty, multiple cartilage-bone plugs are harvested from one (less load-bearing) site and transplanted to fill the larger defect (image downloaded from

intact cartilage matrix and underlying subchondral bone are transplanted into the defect. Major disadvantages to these procedures include donor site morbidity [Temenoff and Mikos, 2000], limited availability and suitability (surface contour and load-bearing capacity) of graft tissue [Hunziker, 1999; Laurencin et al., 1999; Temenoff and Mikos, 2000], and instability and degeneration of the graft [Laurencin et al., 1999].

Although the above procedures improve pain and joint function for many patients [Minas and Nehrer, 1997; Temenoff and Mikos, 2000], thus far no surgical solution has been able to fully regenerate articular cartilage and proven to be a long-term solution.

1.4. TISSUE ENGINEERING OF ARTICULAR CARTILAGE

An alternative, or possible augmentation, to the above-mentioned surgical techniques, is tissue engineering of articular cartilage. Tissue engineering, as coined by Langer and Vacanti involves the use of cells, scaffolds, and/or regulators to grow functional tissue ex vivo [Langer and Vacanti, 1993]. More recently, "regenerative medicine" has been adopted to refer to procedures in which the majority of the regeneration process occurs in vivo. This thesis is motivated by the latter approach. However, for the purpose of this write-up, the term "tissue engineering" will be used. 1.4.1. General Approach

The relatively simple structure of articular cartilage (avascular, aneural, single cell population), limited self- and surgically-induced repair, and the clinical consequences for cartilage repair, have made tissue engineering of articular cartilage an area of intense research over the past decade [Temenoff and Mikos, 2000]. In addition to the three traditional pillars of tissue engineering: 1) the scaffold to serve as an analogue of the natural extracellular matrix, 2) the cells, and 3) regulators and/or growth factors [Langer and Vacanti, 1993], it is important to also consider the role of physical forces in the development of musculoskeletal tissues, such as articular cartilage. Due to the complex physical loads to which orthopaedic tissues are subjected to in vivo, the optimal tissue engineering schemes for articular cartilage may involve a combination of in vitro and in

vivo stages.

1.4.2. Review of Previous Work

1.4.2.1.Scaffolds

A wide array of materials has been used in various in vitro and in vivo studies for

articular cartilage engineering. Candidate scaffolds must be biocompatible and accommodate cell proliferation and matrix synthesis. The scaffold serves as an analog of the natural three-dimensional extracellular matrix, and can also be used as a carrier of cells, growth factors, and/or genetic material.

Scaffold composition and porosity is known to affect cell viability, attachment, morphology, and macromolecular biosynthesis [Coutts et al., 1994; Frenkel et al., 1997; Grande et al., 1997; Nehrer et al., 1997b]. Materials that are most often studied in cartilage tissue engineering include hydrogels made up of collagen [Kawamura et al.,

1998; Rich et al., 1994; Wakitani et al., 1994], agarose [Buschmann et al., 1992; Cook et

al., 1997; Rahfoth et al., 1998], and synthetic peptides [Kisiday et al., 2001]; sponge-like scaffolds manufactured from collagen [Frenkel et al., 1997; Grande et al., 1997; Nehrer et al., 1998b; Nehrer et al., 1997a; Nixon et al., 1993; Sams et al., 1995; Toolan et al., 1996; Toolan et al., 1998], polyglycolic acid [Freed et al., 1994; Freed et al., 1993; Grande et al., 1997], and/or polylactic acid [Chu et al., 1997; Coutts et al., 1994; Freed et al., 1993]; and materials with a naturally-occurring porous structure, such as coral [Shahgaldi, 1998] and devitalized articular cartilage [Toolan et al., 1998].

100 AM

Figure 1.3. A porous collagen-GAG scaffold. This scanning electron microscope image shows the

porous structure of a typical scaffold resulting from freeze-drying of a type I collagen-GAG co-precipitate used in the studies described in Chapters 2 and 5 of this thesis.

While each of these scaffolds has particular advantages and disadvantages, there are general characteristics of the different types of scaffolds. For example, hydrogels are easy to fabricate and allow for uniform and predictable cell-seeding of the gels. The gels, however, have poor mechanical properties and are difficult to handle during implantation surgery. In contrast, the porous scaffolds have higher mechanical integrity and can be sutured or press-fit into cartilage defects. The porous nature of such scaffolds, however, result in a lower retention of newly synthesized macromolecules.

1.4.2.2. Growth Factors

Tissue engineering scaffolds can also be used as carriers for therapeutic proteins or genes for the proteins. Various growth factors (i.e. fibroblast growth factor-2, transforming growth factor-$, insulin-like growth factor-1, and osteoprogenitor factor-1) have been used to modulate chondrocyte phenotype, proliferation, and biosynthesis rates. Such growth factors may be tethered to insoluble scaffolds to permit localized dosing in vivo. Alternatively, advancing research in gene therapy may utilize tissue engineering scaffolds for direct, localized delivery of specific genes to cells for prolonged in situ production of growth factors in therapeutic quantities. Recently, it has been reported that polymer scaffolds loaded with genetic material can be used to transfect cells over prolonged periods [Bonadio et al., 1999; Lauffenburger and Schaffer, 1999]. Furthermore, biodegradable scaffolds can be designed to deliver the desired proteins or genes gradually over a period of time dependent on the degradation rate of the scaffold. 1.4.2.3. Cells

Due to the low metabolic rate of native, mature chondrocytes, most approaches to tissue engineering of articular cartilage involves transplantation of cells along with the scaffold. Traditionally, autologous articular chondrocytes are used [Breinan et al., 2000; Breinan et al., 1998; Breinan et al., 1997; Brittberg et al., 1994; Frenkel et al., 1997; Kawamura et al., 1998], but allogenic chondrocytes [Rahfoth et al., 1998; Toolan et al.,

1998], chondrocytes from other cartilaginous tissues [Bouwmeester et al., 1997; Chu et

al., 1997; Klein-Nulend et al., 1998], and chondroprogenitor cells [Johnstone and Yoo,

1999; Nixon et al., 1999; Wakitani et al., 1997] have also been used.

The low cell density of cartilage requires the expansion of the harvested cell population if autologous articular chondrocytes are used. Chondrocytes expanded in

monolayer cultures (passaged cells) demonstrate increased rates of proliferation [Green,

1971], but de-differentiate and lose their characteristic phenotype - namely they lose their spherical morphology and synthesize type I rather than type II collagen [Benya and Shaffer, 1982]. Fortunately, the chondrocyte phenotype can be re-induced (from these de-differentiated passaged cells) or induced (from progenitor cells) by the appropriate matrix environment [Benya and Shaffer, 1982], growth factors [Benya and Padilla, 1990; Benya et al., 1988; Benya and Padilla, 1993; Borge et al., 1997; Jakob et al., 2001; Johnson et al., 2001; Kulyk and Reichert, 1992; Martin et al., 2001a; Martin et al.,

2001b], or other culture variables [Domm et al., 2000].

1.4.2.4.Regulation by physical forces

Although the precise mechanisms by which physical loads effect biological responses have yet to be resolved, it has been well-established that mechanical loads regulate chondrocyte behavior both in vivo [Grumbles et al., 1995; Kiviranta et al., 1988; Paukkonen et al., 1986] and in vitro [Buschmann et al., 1995; Buschmann et al., 1999; Gray et al., 1989; Grodzinsky et al., 1998; Jin et al., 1999; Kim et al., 1995; Kim et al., 1994; Lee et al., 1981; Martin et al., 2000; Quinn et al., 1998a; Quinn et al., 1998b; Quinn et al., 1999; Ragan et al., 1999b; Sah et al., 1989]. Specifically, static compression [Buschmann et al., 1995; Kim et al., 1996; Ragan et al., 1999a] and high-impact compressive loading [Loening et al., 1999; Quinn et al., 1998b] impair chondrocyte metabolism, while dynamic compression [Kim et al., 1994; Quinn et al., 1998a; Sah et al., 1989] and shear [Jin et al., 1999] at moderate amplitudes can increase biosynthetic activity.

Comparatively little research has been done on chondrocytes in tissue engineering systems. It has been established that chondrocytes in non-native scaffolds such as agarose [Buschmann et al., 1995] and alginate gels [Ragan et al., 2000] respond to mechanical compression in a similar manner to chondrocytes in cartilage explants, but only once the cells become encapsulated in extracellular matrix. In the development of cartilage constructs, dynamic compression of chondrocyte-seeded agarose gels [Mauck et al., 2000] and laminar fluid flow on chondrocyte-seeded PGA fibrous meshes [Martin et al., 2000] have been used over extended periods of time to increase cartilage-like matrix production.

1.4.2.5.The Collagen-Glycosaminoglycan System

Porous collagen-glycosaminoglycan (CG) scaffolds, developed for the purpose of dermal tissue engineering [Yannas and Burke, 1980; Yannas et al., 1980; Yannas et al.,

1989], have also shown promise in promoting the healing of peripheral nerve

[Chamberlain et al., 1998; Chamberlain et al., 2000] and conjuctiva [Hsu et al., 2000]. Previous studies in our laboratory have shown that the CG scaffold may also be used to promote articular cartilage repair. In vivo studies using a canine model for articular cartilage repair, indicate that implantation of a CG scaffold, either alone or seeded with cells [Breinan et al., 2000; Nehrer et al., 1998b] improved healing of surgically created cartilage defects compared to untreated or ACI-treated defects. The repair tissue formed after implantation of the scaffolds, however, was principally fibrocartilage rather than the desired hyaline cartilage.

The cell-seeded scaffolds utilized autologous chondrocytes that were expanded in monolayer culture prior to seeding into the CG scaffolds. Although the extent of de-differentiation and re-de-differentiation were not quantified, in vitro studies have shown that passaged articular chondrocytes will proliferate and synthesize glycosaminoglycans, and may also synthesize type II collagen when seeded into the CG scaffolds [Nehrer et al.,

1998a; Nehrer et al., 1997a; Nehrer et al., 1997b].

In order to further improve cartilage repair stemming from implantation of cell-seeded CG scaffolds, a more thorough understanding of the chondrocyte behavior in the scaffolds is necessary. In particular, the physical properties of the scaffold and the response of the seeded chondrocytes to mechanical forces are important to the progression of cartilage-specific tissue engineering systems. Additionally, it may be of great interest to use the CG scaffolds to transfect chondrocytes and/or infiltrating stem cells in situ.

1.5. SPECIFIC AIMS

To lay the groundwork for future development of the CG scaffolds for articular cartilage tissue engineering, the purpose of this thesis was to explore the interactions of passaged chondrocytes with the porous CG scaffolds. To better understand how the physical properties of the CG scaffold affect chondrocyte behavior, the effects of scaffold

cross-linking and mechanical loading of cell-seeded scaffolds were investigated in vitro and used to design an implant for in vivo study. Additionally, to determine how the scaffold could be used in future work involving therapeutic growth factors, experiments were conducted to determine the feasibility of transfected chondrocytes in situ with plasmid DNA attached to the CG scaffold. The working hypotheses were:

1. Increasing cross-link density would decrease chondrocyte-mediated contraction and increase chondrocyte proliferation and biosynthesis;

2. Mechanical compression of cell-seeded CG scaffolds could be used to stimulate chondrocyte biosynthetic activity;

3. In vitro culture of cell-seeded CG constructs prior to implantation into chondral defects would improve repair of the cartilage lesions; and

4. CG scaffolds loaded with genetic material could be used to transfect

chondrocytes over an extended period of time.

To test these four hypotheses, in vitro and in vivo experiments were conducted with the following specific aims:

L.a. To determine the effects of dehydrothermal (DHT), ultraviolet (UV), glutaraldehyde (GTA), and carbodiimide (EDAC) cross-linking treatments on compressive stiffness and glycosaminoglycan (GAG) content of unseeded CG scaffolds;

1.b. To quantify the extent of in vitro cell-mediated contraction of scaffolds with different compressive stiffnesses by passaged, adult canine chondrocytes;

1.c. To evaluate in vitro chondrocyte proliferation and protein and GAG accumulation rates in the cross-linked scaffolds;

2.a. To evaluate the effects of 24 hours of static and dynamic compression on protein and GAG synthesis by passaged adult chondrocytes seeded in

EDAC cross-linked scaffolds;

2.b. To quantify protein and GAG accumulation rates in cell-seeded CG scaffolds subjected to up to 24 hours of static or dynamic compression in vitro;

3.a. To evaluate the histological make-up of repair tissue formed in vivo 15 weeks after the implantation of a 4 week in vitro cultured chondrocyte-seeded CG construct for comparison to previous results in the same animal model; and

3.b. To compare the mechanical indentation properties of the repair tissue to normal articular cartilage;

4.a. To assess the effects of the pH of the solution of plasmid DNA added to the CG scaffold on the leaching characteristics of DNA from

gene-supplemented CG (GSCG) scaffolds;

4.b. To evaluate the effects of GSCG scaffold cross-linking on plasmid leaching characteristics; and

4.c. To measure transfection levels of chondrocytes cultured in GSCG scaffolds for 2, 4, or 8 weeks in vitro.

CHAPTER 2: EVALUATION OF CROSS-LINKING

METHODS FOR COLLAGEN-GLYCOSAMINOGLYCAN

SCAFFOLDS

2.1. INTRODUCTION

Previous studies investigating the behavior of passaged adult canine chondrocytes in collagen-glycosaminoglycan (CG) scaffolds have noted significant dimensional changes in the cell-seeded scaffolds over time [Lee et al., 2000a; Nehrer et al., 1997a]. That articular chondrocytes have the potential for contraction, was recently supported by immunohistochemical findings of a contractile muscle actin, a-smooth muscle actin

(SMA), in human [Kim and Spector, 2000] and canine [Wang et al., 2000] articular

chondrocytes in situ and in the reparative cartilaginous tissue in healing defects in a canine model [Wang et al., 2000]. This finding represents a potential problem in the application of such scaffolds for tissue engineering. As the scaffold contracts, there is a reduction in the pore volume that could restrict cell proliferation. Additionally, in vivo deformation of the construct could result in a loss of contact between the implanted device and the host tissue, thereby decreasing the chances for successful integration of the repair tissue.

Recognizing the potential importance of chondrocyte-mediated contraction of scaffolds employed for tissue engineering, one objective of this experiment was to evaluate the ability of various cross-linking methods to increase the cross-linking density and stiffness of the CG scaffold. The hypothesis was that increased levels of cross-linking could be used to sufficiently increase scaffold stiffness to thwart chondrocyte-mediated contraction. The cross-linking treatments evaluated here were: dehydrothermal treatment (DHT) [Weadock et al., 1983; Weadock et al., 1995; Weadock et al., 1996; Yannas et al., 1989]; ultraviolet irradiation (UV) [Weadock et al., 1983; Weadock et al., 1995; Weadock et al., 1996]; glutaraldehyde (GTA) [Petite et al., 1994; Weadock et al., 1983; Yannas et al., 1989], and carbodiimides (EDAC) [Olde Damink et al., 1996;

Osborne et al., 1998; Osborne et al., 1999; Weadock et al., 1983]. In addition to affecting stiffness, different cross-linking protocols can affect scaffold degradation rate [Petite et

al., 1994; Weadock et al., 1983; Weadock et al., 1996], cell proliferation [Osborne et al.,

1998; Petite et al., 1994; Weadock et al., 1983] and biosynthesis [Chevallay et al., 2000].

Therefore, another objective of this experiment was to evaluate the effects of the selected cross-linking treatments on the proliferative and biosynthetic activity of adult canine articular chondrocytes seeded in the scaffolds. The hypothesis was that the changes in physical and chemical properties afforded to the CG scaffolds by the different cross-linking methods would affect adult canine chondrocyte proliferation and protein and

GAG biosynthesis within the scaffolds.

2.2. MATERIALS AND METHODS

2.2.1. Collagen-Glycosaminoglycan Scaffold Fabrication

2.2.1. 1.Freeze-Drying

The porous CG scaffold was produced by freeze-drying a co-precipitate of type I bovine tendon collagen (Integra Life Sciences, Plainsboro, NJ) and shark chondroitin-6-sulfate (Sigma Chemical, St. Louis, MO) as previously described [Yannas et al., 1989] and outlined in detail in Appendix A. The final concentration of the slurry was 5.0 mg collagen/ml and 0.44 mg chondroitin sulfate/ml. The scaffolds have been previously reported to have a porosity of approximately 87% and an average pore size of 84 pm [Nehrer et al., 1997a]. Nine-millimeter diameter disks (approximately 3.5 mm thickness) were used for subsequent mechanical tests and cell culture experiments.

2.2.1.2. Cross-Linking Methods

All scaffolds were sterilized and minimally cross-linked for 24 hours by

dehydrothermal treatment (DHT) [Yannas et al., 1989]. Certain scaffolds were further cross-linked as follows: a) by exposure to ultraviolet radiation (UV) for one hour (5 cm from a 258nm source rated at 4510 pW/cm2 _{at 5 cm), with the scaffolds being turned}

over after thirty minutes, b) immersion in a 0.25% glutaraldehyde solution in 0.05 M acetic acid for 24 hours at room temperature (GTA), or c) immersion in a carbodiimide solution (14 mM 1 -ethyl-3-(3-dimethylaminopropyl)carbodiimide hydrochloride and 5.5 mM N-hydroxysuccinimide; Sigma) for two hours at room temperature (EDAC). Excess glutaraldehyde was removed from the scaffold through a series of four washes in sterile

distilled water. Excess EDAC was rinsed from the scaffold using phosphate-buffered saline (PBS) followed by two washes in sterile distilled water.

Table 2-1. Summary of cross-linking treatments.

Group Cross-linking Treatment

DHT 24 hours dehydrothermal treatment at 105'C, 30 mtorr

24 hours DHT plus 30 minutes ultraviolet irradiation (4510 gW/cm2

X=258 nm) to each side of scaffold

24 hours DHT plus 24 hours immersion in 0.25% gluteraldehyde in

GTA 0.05 M acetic acid, followed by 4x15 minute rinse in sterile distilled,

deionized water

24 hours DHT plus 2 hours immersion in 14 mM EDAC/5.5 mM NHS

EDAC solution in distilled water, followed by 2 hour rinse in sterile phosphate buffered saline

2.2.2. Physical Characterization of Scaffolds

2.2.2. 1.Swelling Ratio Determination

Cross-link density for randomly coiled polymer networks is inversely related to the swelling ratio [Weadock et al., 1983]. Thus, as an approximate measure of the density of cross-links that were formed by the different cross-linking methods, the swelling ratios of the scaffolds were determined as described by Weadock et al. [Weadock et al., 1983]. Cross-linked CG samples placed in a water bath at 90'C for two minutes to swell and denature the collagen. Water within the pores was expelled by pressing the swollen scaffolds between sheets of filter paper with a 1.0 kg weight placed on top for 20 seconds. The sample was weighed and the weight recorded as wet weight (WW). Samples were then dried in an oven (1 100C) overnight and the dry weights (DW)

of the collagen scaffolds determined. The swelling ratio, defined as the inverse of the volume fraction of dry collagen (Vf), was calculated from the wet and dry weights and the densities of water (Pwater=.00 g/cm3) and collagen (p,=1.32 g/cm3) as follows:

r +DW (WW -D W

)7J]

r =- D- Eq. 2-1

2.2.2.2. Compression Testing of Scaffolds

Disks (9 mm in diameter) from each cross-linking group were hydrated in phosphate buffered saline (PBS) and stored at 40C prior to mechanical testing. On the day of testing, the thickness of the fully hydrated specimen was measured using a micrometer. Disks were then placed in a PBS-filled polymethylmethacrylate (PMMA) chamber mounted in the lower jaw of a Dynastat Mechanical Spectrometer (IMASS, Hingham, MA). A 50-gram load cell (Sensotec, Cleveland, OH) fitted with a 9.5-mm diameter PMMA cylindrical plunger was fixed in the upper jaw of the Dynastat and the distance between the plunger and the lower chamber set to the thickness of the hydrated scaffold.

Using displacement-feedback control, successive ramp-and-hold displacements were applied in radially-unconfined compression, giving sequential strain increments of

1-5% up to a maximum of 40% strain (see Appendix C). Radially-unconfined compression was used because it was desired to design a single chamber that would accommodate future testing of cell-seeded scaffolds which are known to contract in diameter over time. At each strain level stress relaxation was achieved in 30-75 seconds, depending on the magnitude of the displacement, and the equilibrium loads were recorded. Stress was computed as the load normalized to the initial unstrained disk area.

Compression testing of samples occurred within 48 hours of hydration (DHT and

UV) or the completion of cross-linking (GTA and EDAC). Additionally, DHT and GTA

samples stored in sterile PBS at 4C for six weeks were tested to determine if the stiffness changed over time.

2.2.2.3.Glycosaminoglycan Content of Scaffolds

To determine the amount of GAG tightly bound to the collagen network after cross-linking, the sulfated GAG content of unseeded scaffolds was determined by the dimethylmethylene blue (DMMB) dye assay (Appendix E.3). Unseeded, cross-linked scaffold disks were hydrated in PBS to rinse away unbound GAG. The disks were then lyophilized and solubilized overnight at 60'C with 1 ml of papain buffer (6 pg/ml papain and 10 mM cysteine-HCl in 0.1 M sodium phosphate and 5 mM Na2EDTA, PBE). An aliquot of the digest (20 pl) was mixed with 200 p1 of the DMMB dye in a 96-well

microplate and the absorbance at 520 nm was measured. Shark chondroitin-6-sulfate was used as the standard.

2.2.3. Cell-Seeded Assays

2.2.3.1. Chondrocyte Isolation and Culture

Chondrocytes were isolated from articular cartilage from the patellar, femoral and tibial surfaces of knee (stifle) joints of three adult mongrel dogs using sequential pronase (one hour at 37'C; 20 U/ml; Sigma Chemical, St. Louis, MO) and collagenase (overnight at 37*C; 200 U/ml; Worthington Biochemical) digestion as described by Kuettner, et. al [Kuettner et al., 1982] and detailed in Appendix D. Isolated chondrocytes were resuspended in culture medium (DMEM/F12, Gibco Life Sciences, Grand Island, NY) supplemented with 10% fetal bovine serum (FBS, Hyclone Technologies, Logan, UT), 25

gg/ml ascorbic acid (Wako Chemical, Osaka, Japan), and a

Penicillin/Streptomycin/Fungizone cocktail (Gibco) and plated in 75-cm2 _{flasks at a} density of 2 million cells/flask. The culture flasks were incubated at 37C with 5% CO2.

Cells were cultured to confluence, trypsinized, resuspended and replated into 75-cm2

flasks. Chondrocytes from the three different animals were isolated and cultured separately throughout.

2.2.3.2.Cell Seeding and Culture of CG Scaffolds

Third passage cells were collected by trypsinization and resuspended in complete medium at a concentration of 4x106 cells/ml. CG disks were incubated with 0.5 ml of cell suspension per disk for 1.5-2 hours on a rocking table (n=16 for each of the 3 animals). This targeted seeding density (2x 106 cells/disk) aimed to seed the cells at a near physiological density for adult cartilage (10,000 cells/mm3 _{[Muir, 1995]).}

Approximately 50% of the chondrocytes (lx106 cells/scaffold, 5,000 cells/mm3_{) attach to}

the scaffolds by this seeding method (Appendix J). Disks were then transferred to agarose-coated wells (12-well plates) with 1.0 ml of complete media per well and returned to the incubator overnight. The following day, an additional 0.5 ml of medium was added to each well. Media (1.5 ml) were changed every other day. Cultures were terminated after 2, 7, 15, or 29 days. Unseeded disks were cultured as controls.

2.2.3.3.Measurement of Cell-Mediated Contraction

The diameters of the seeded and unseeded scaffolds were measured 1, 3, 5, 7, 15, 21, and 29 days post-seeding. Contraction was calculated as the relative change in scaffold diameter from the day 1. Cell-mediated contraction (CMC) was determined by subtracting the contraction of the unseeded scaffolds from the contraction of the seeded scaffolds.

CMC = OriginalDiameter - Diameter OriginalDiameter - Diameter Eq. 2-2

OriginalDiameter OriginalDiameter nveage

2.2.3.4.Radiolabel Incorporation

On days 2, 7, 15, and 29, nine seeded (three per animal) and one to three unseeded scaffolds from each cross-linking group were terminated for biochemical analyses of synthesis rates and proliferation. During the last eight hours of culture, constructs were cultured in complete media supplemented with 10 gCi/ml of 3_{H-proline and 10 RCi/ml of}

3 5

S-sulfate, to assess rates of total protein and GAG synthesis, respectively. At the end of the labeling period, disks were washed (4x15 minutes at 40C) in PBS supplemented with unlabeled proline (1 mM) and sulfate (0.8 mM), lyophilized and digested with 1 ml papain buffer as described above for the digestion of the unseeded scaffolds for GAG determination. Radiolabel incorporation was determined by mixing 100 gl of the papain digest with 2 ml scintillation cocktail (EcoLume, Costa Mesa, CA) and measuring 3_{H and}

35S counts per minute (cpm) in a liquid scintillation counter (Rack-Beta 1211 LKB, Turku, Finland), with corrections for spillover (Appendix E.5). Counts were normalized to DNA content (see below).

2.2.3.5.Dry Weight Determination

In order to determine the net rates of matrix accumulation and degradation for the different scaffolds, the weights of the scaffolds were measured after lyophilization (DW). Both the total DW and the net change in DW were analyzed.

2.2.3.6.DNA Analysis

The DNA content of the constructs was measured using Hoechst 33258 dye (Appendix E.4). A 20 gl aliquot of the papain digest was mixed with 80 pl of phosphate buffered EDTA and 2 ml of Hoechst dye solution (10% Hoechst dye in 10 mM Tris, 1

mM Na2EDTA, and 0.1 M NaCl, pH 7.4) and assayed fluorometrically. Calf thymus DNA was used as the standard. The background fluorescence of the scaffold was accounted for by subtracting the values obtained for the unseeded scaffolds.

2.2.3.7.Glycosaminoglycan Content of Cell-Seeded Scaffolds

The GAG content of cell-seeded scaffolds was determined from the digests of the constructs using the DMMB assay described above for the unseeded scaffolds.

2.2.3.8.Immunohistochemistry

At each sacrifice time point (days 2, 7, 15, and 29), three seeded samples (one from each animal) and one unseeded sample were fixed in 10% neutral buffered formalin, dehydrated, and embedded in paraffin. Specimens were sectioned (7 gm thick) in cross-section and stained for type II collagen. Deparaffinized slides were prepared for immunostaining by digestion in 0.1% protease XIV (diluted in Tris-buffered saline, pH 7.4, TBS; Sigma) for 60 minutes and non-specific staining was blocked with application of 30% goat serum (Sigma) for 20 minutes. Sections were incubated with the primary antibody to either a-smooth muscle actin (Sigma) or type II collagen (II-116B3, prepared

by T. Linsenmayer and obtained from the Developmental Studies Hybridoma Bank,

University of Iowa, Iowa City, IA; diluted 1:20 in TBS) or TBS alone (negative control) for 2 hours, followed by incubation with biotinylated goat anti-mouse IgG antibody (Sigma; 1:200 in TBS) for 45 minutes. Endogenous peroxidase was quenched with 3% hydrogen peroxide (10 minutes) and sections were then incubated with ExtrAvidin-Conjugated Peroxidase (Sigma; 1:50 in TBS) for 20 minutes and developed with AEC (Zymed Laboratories, Inc., S. San Francisco, CA) and counter-stained with Mayer's hematoxylin for 20 minutes.

2.2.4. Statistical Analysis

There was no significant effect of animal on any of the measured parameters by one-way analysis of variance (ANOVA). Therefore, the data from all three animals were pooled and are reported as the mean ± standard error of the mean (SEM). ANOVA and Fisher PLSD post-hoc testing were performed using StatView (SAS Institute, Inc, Cary, NC).

2.3. RESULTS

2.3.1. Physical Characterization of Unseeded Scaffolds

2.3.1. 1.Swelling Ratio

The swelling ratio of the CG scaffolds, calculated as the inverse of the volume fraction of collagen, ranged from 5.3 ± 0.2 (mean ± SEM) for the DHT scaffolds to 3.2 ±

0.1 for the EDAC scaffolds (Figure 2.1). Taking the cross-link density to be proportional

to the inverse of the swelling ratio, the density of cross-links increased with the different cross-linking methods as follows: DHT<UV<GTA<EDAC (p<0.05 for each Fisher

PLSD post-hoc test). *4

0.35

-0.3

-0.25

-0.2

0.15

-0.1

0.05

-

0-DHT

UV

GTA

EDAC

Figure 2.1. Inverse swelling ratio of cross-linked scaffolds. Cross-link density is proportional to

the inverse of the swelling ratio and increased with UV, GTA, and EDAC cross-linking treatments of DHT cross-linked scaffolds. Mean ± SEM; n=4. All groups were significantly different from each other.

28

* '0

14.17

1400

1200-

1000-

800-

600-

400-

20

0-346

369

145

DHT

UV

GTA

EDAC

Figure 2.2. Compressive stiffness of cross-linked scaffolds. Unconfined compressive stiffness of

hydrated DHT cross-linked scaffolds was the lowest. UV and GTA cross-linking more than doubled scaffold stiffness, while EDAC cross-linking increased stiffness 9-fold over DHT cross-linking alone. Mean ± SEM; n=5-10. All groups significantly different except the pair marked NS.

1400

1200

U

1000~

800

-600

-

400-200

0-0.15

0.2

0.25

0.3

0.35

1/Swelling Ratio

Figure 2.3. Compressive stiffness vs. Inverse swelling ratio. Unconfined compressive stiffness

increase linearly with increasing cross-link density (R2

=0.80). The correlation improved to R2 =0.98

when the compressive stiffness of the GTA scaffold after 6 weeks storage was utilized (open diamond).

NS

F-f-I

1266

0.80

2.3.1.2. Compressive Stiffness

The equilibrium, uniaxial unconfined compressive stress-strain relation was remarkably linear (see Appendix C) with coefficients of determination above 95%. The slope of the linear fit to the stress-strain data was used to define the "apparent compressive modulus" of the scaffolds. This method yielded moduli ranging from 145 + 23 Pa (mean ± SEM) for the DHT samples to 1117 ± 109 Pa for the EDAC samples

(Figure 2.2). Compared to the minimally cross-linked DHT scaffolds, the compressive modulus was doubled by UV and GTA cross-linking protocols and increased another three-fold by the EDAC cross-linking protocol. ANOVA revealed the highly significant effect of cross-link treatment on modulus (p<0.0001). Post-hoc testing showed that each group was different from the others except for the UV and GTA groups.

The stiffness of DHT scaffolds evaluated after six weeks of storage in sterile PBS did not change significantly. The stiffness of the GTA scaffolds, however, increased from 369 ±56 to 663 ±64 Pa.

Increasing cross-linking density, as indicated by the inverse of the swelling ratio, correlated with an increase in compressive stiffness (R2_{=0.80, Figure 2.3). The} correlation increased to R2_{=0.98 when the compressive stiffness of the GTA scaffold} after 6 weeks in storage was used.

2.3.1.3.GAG Content of Scaffolds

All scaffolds were fabricated from similar batches of collagen-GAG slurry

containing 8.8% GAG/DW. The GAG content of the rinsed, unseeded scaffolds was less than that of the original slurry, indicating that not all of the added chondroitin sulfate became tightly bound to the collagen fibrils during the mixing and freeze-drying process. Samples cross-linked by UV irradiation or EDAC had significantly more GAG than the samples cross-linked by DHT or GTA (Figure 2.4; p<0.01). The EDAC scaffolds had

110 ± I ptg GAG/disk (3.4 ± 0.06% GAG/DW), compared to only 33.7 ± 3.6 [tg

GAG/disk (1.2 ± 0.09% GAG/DW) in the DHT cross-linked scaffolds. These values

span the range measured for adult canine articular cartilage (1.5-3% GAG/DW [Lee,

1999]). The GAG content of the unseeded scaffolds decreased slightly from the day 1 to

4.0%-

3.5%-3.0%

2.5%

2.0%

1.5%

0

1.0%-0.5%

0.0%

DHT

UV

GTA

EDAC

Figure 2.4. GAG content of unseeded scaffolds. GAG content normalized to dry weight of

cross-linked scaffolds was higher for UV and EDAC cross-cross-linked scaffolds. Mean ± SEM; n=4-8; GAG

content was significantly different for all comparisons except the pairing marked by t.

day 7 readings, indicating that not all of the unbound GAG was rinsed from the scaffolds in the initial PBS rinse (see also GAG analysis of seeded scaffolds, section 2.3.3). There was no noticeable change in GAG content of the unseeded scaffolds beyond the first week.

2.3.2. Cell-Seeded Assays

2.3.2.1. Cell-mediated Contraction

The diameters of the unseeded scaffolds changed less than 5% between the first and 2 8th day in culture. In contrast, all cell-seeded constructs underwent a minimum of 30% reduction in diameter by the end of the four-week culture period (Figure 2.5). The

patterns and the magnitudes of contraction were different for the different cross-linking protocols. The average diameter of the seeded DHT and UV scaffolds decreased to 60% of their original value during the first week, and by the end of four weeks these scaffolds had contracted to 40% and 45% of their original dimensions, respectively. In contrast, the seeded GTA and EDAC scaffolds did not contract during the first week in culture. Thereafter, the seeded GTA scaffolds contracted at an approximately constant rate to a