Functional Tissue Engineering: The Role of Biomechanics

Functional Tissue Engineering: The Role of Biomechanics David L. Butler Noyes-Giannestras Biomechanics Laboratories, Department of Aerospace Engineering and Engineering Mechanics, University of Cincinnati, Cincinnati, OH 45221-0070

Steven A. Goldstein Orthopædic Research Laboratories, Orthopædic Surgery, University of Michigan, Ann Arbor, MI 48109-0486

Farshid Guilak Orthopædic Research Laboratories, Department of Surgery, Duke University Medical Center, Durham, NC 27710

‘‘Tissue engineering’’ uses implanted cells, scaffolds, DNA, protein, and/or protein fragments to replace or repair injured or diseased tissues and organs. Despite its early success, tissue engineers have faced challenges in repairing or replacing tissues that serve a predominantly biomechanical function. An evolving discipline called ‘‘functional tissue engineering’’ (FTE) seeks to address these challenges. In this paper, the authors present principles of functional tissue engineering that should be addressed when engineering repairs and replacements for load-bearing structures. First, in vivo stress/strain histories need to be measured for a variety of activities. These in vivo data provide mechanical thresholds that tissue repairs/replacements will likely encounter after surgery. Second, the mechanical properties of the native tissues must be established for subfailure and failure conditions. These ‘‘baseline data’’ provide parameters within the expected thresholds for different in vivo activities and beyond these levels if safety factors are to be incorporated. Third, a subset of these mechanical properties must be selected and prioritized. This subset is important, given that the mechanical properties of the designs are not expected to completely duplicate the properties of the native tissues. Fourth, standards must be set when evaluating the repairs/replacements after surgery so as to determine, ‘‘how good is good enough?’’ Some aspects of the repair outcome may be inferior, but other mechanical characteristics of the repairs and replacements might be suitable. New and improved methods must also be developed for assessing the function of engineered tissues. Fifth, the effects of physical factors on cellular activity must be determined in engineered tissues. Knowing these signals may shorten the iterations required to replace a tissue successfully and direct cellular activity and phenotype toward a desired end goal. Finally, to effect a better repair outcome, cell-matrix implants may benefit from being mechanically stimulated using in vitro ‘‘bioreactors’’ prior to implantation. Increasing evidence suggests that mechanical stress, as well as other physical factors, may significantly increase the biosynthetic activity of cells in bioartificial matrices. Incorporating each of these principles of functional tissue engineering should result in safer and more efficacious repairs and replacements for the surgeon and patient. 关S0148-0731共00兲00206-5兴 Keywords: Biomechanics, Tissue Engineering, Gene Therapy, Biomaterials, Cellular Engineering

The goal of ‘‘tissue engineering’’ is to repair or replace tissues and organs by delivering implanted cells, scaffolds, DNA, proteins, and/or protein fragments at surgery. Tissue engineering merges aspects of engineering and biology, and many rapid achievements in this field have arisen in part from significant advances in cell and molecular biology 共e.g., the isolation and manipulation of cells, genes, and growth factors兲, biomaterials 共new and innovative delivery vehicles兲, and the integration of biology and materials to deliver viable cells in compatible support structures. Many of the tissues and organs to be replaced have an important biomechanical function. In fact, the biomechanical properties of these tissues are critical to their proper function in vivo. In order for tissue engineers to repair or replace these load-bearing structures effectively, they must address a number of significant questions. What are the thresholds of force, stress, and strain that normal tissues transmit or encounter? What are the mechanical properties of these tissues when subjected to expected in vivo stresses and strains, as well as under failure conditions? Which of these properties should a tissue engineer insist upon incorporating into the design? When evaluating the resulting tissue engineered

repairs, how good is good enough 共i.e., do tissue engineered repairs and replacements need to exactly duplicate the structure and function of the normal tissue or organ兲? When developing these implants in culture, how do physical factors such as mechanical stress regulate cell behavior in bioreactors as compared to signals experienced in vivo? And finally, can tissue engineers mechanically stimulate these implants before surgery to produce a better repair outcome? To address these questions and others, the United States National Committee on Biomechanics 共USNCB兲 formed a subcommittee1 in 1998. This committee adopted the concept of ‘‘Functional Tissue Engineering,’’ or FTE. The USNCB’s goals in advancing FTE were to: 共1兲 increase awareness among tissue engineers about the importance of restoring ‘‘function’’ when engineering tissue constructs; 共2兲 identify the critical structural and mechanical requirements needed for each tissue engineered construct; and 共3兲 encourage tissue engineers to incorporate these functional criteria in the design, manufacturing, and optimization of tissue engineered constructs. To address its first goal, the USNCB sponsored a panel session, ‘‘Functional Tissue Engineering—The Role of Biomechanics’’ at the recent 1999 Summer ASME Bioengineering Conference in Big Sky, Montana. Three speakers were invited to give their per-

Contributed by the Bioengineering Division for publication in the JOURNAL OF BIOMECHANICAL ENGINEERING. Manuscript received by the Bioengineering Division February 7, 2000; revised manuscript received July 24, 2000. Associate Technical Editor: R. Vanderby, Jr.

1 USNCB Subcommittee 共D. Butler, U. Cincinnati; S. Goldstein, U. Michigan; V. Mow, Columbia U.; G. Schmid-Schonbein, U. California, San Diego; L. Soslowsky, U. Pennsylvania; R. Spilker, Rensselaer Polytechnic Institute; S. Woo, U. Pittsburgh兲.

Introduction

570 Õ Vol. 122, DECEMBER 2000

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spectives on FTE for the musculoskeletal system. Drs. Farshid Guilak, Steven Goldstein, and David Butler evaluated aspects of FTE related to articular cartilage, bone, and tendon/ligament, respectively. For the purposes of this paper, common aspects of these presentations have been recast into ‘‘principles of functional tissue engineering’’ so as to emphasize commonalities in desired outcome. While the process used to repair these tissue classes successfully may differ, as do their expected mechanical environments in vivo, common features of all three deserve consideration. Musculoskeletal as well as other tissue systems 共e.g., cardiac muscle and blood vessel兲 can benefit from new and exciting developments in FTE. Obviously, the strategies employed to engineer these tissue products will be unique to the structure under consideration. What constitutes ‘‘success’’ will also vary among tissues. For example, tissues or systems that are designed to prolong life may tolerate a lower margin for error than those that are designed to improve the quality of life. The difficulty in performing a procedure, and the duration that a specific treatment lasts, may also factor into its perceived success. For example, therapies of replacement or regeneration of bone might be expected to last the lifetime of the individual, while replacement of cartilage may be considered successful if it delays total joint replacement for five to ten years. Some of these issues will be addressed in this paper.

Principles of Functional Tissue Engineering What follows are principles of functional tissue engineering that can dramatically influence both the quality of the implants that tissue engineers design and the repair outcome after surgery. These principles are presented in an order that is designed to require fewer iterations to yield constructs that can effectively repair or replace diseased and injured structures. The principles are by no means complete, but are believed to be critically important to a successful outcome. The paper concludes with a brief description of opportunities that may be available for future research in functional tissue engineering. 1 In Vivo Stress andÕor in Vivo Strain Histories Need to Be Measured in Normal Tissues for a Variety of Activities. Knowing the mechanical ‘‘thresholds’’ that normal tissues encounter for different in vivo activities is critical to designing effective tissue repairs/replacements that can meet functional demands after surgery. While these measurements can be difficult to make, they establish the patterns of activity and the bounds of expected usage. Numerous investigators have measured in vivo forces and strains in ligaments and tendons. In vivo forces have been measured using buckle and E-type gages 关1–4兴, implantable force transducers 关5–11兴, and modified pressure transducers 共e.g., 关8,12兴兲. Investigators have shown that while peak forces can vary both within and between activities, these forces increase with the speed of the activity 关8,9兴. Tendons typically develop forces earlier than joint ligaments, primarily because muscles must first transmit their developed force to an in-series tendon to effect large enough motions before ligaments can develop loads 关12兴. When forces are expressed as percentages of their failure capacity, tendons typically develop much larger forces than ligaments, reaching 30 to 40 percent of ultimate strength, whereas ligaments develop forces that rarely exceed 10 to 12 percent of failure force 关5,9,13兴. These differences in percentage of failure force suggest that ligaments maintain a safety factor of 8 to 10 共except of course during trauma兲 whereas tendons have a factor of only 2.5 to 3. In situ loads measured in cadaveric knees support the small forces in the anterior cruciate ligament 关14,15兴. These forces can also vary within ligaments, as evidenced by the large variations in both in situ forces and in vivo strains in the anterior cruciate ligament for different activities 关14–17兴. For other tissues, however, there is a lack of information on the normal in vivo mechanical environment. Peak stresses in articular Journal of Biomechanical Engineering

cartilage loading against an endoprosthesis have been shown to exceed 18 MPa 关18兴, but stresses in a normal joint have been more difficult to measure. Experiments using pressure-sensitive films have shown that normal stresses may range from 5 to 10 MPa in vivo 关10兴, but may be significantly increased by the presence of an osteochondral defect on the joint surface 关19兴. Surprisingly little information has been reported on the deformation behavior of articular cartilage in vivo. One of the few reports of this nature utilized sequential planar radiographs to show that cartilage deforms no more than 15–20 percent under normal physiologic compression 关20兴. Because of the difficulties involved in measuring the in situ loads and deformations of cartilage under realistic in vivo conditions, many investigators have turned to theoretical models of the diarthrodial joint to predict these parameters 关21兴. This is an area that requires further study and will most likely benefit from both theoretical and experimental approaches. Analogous to cartilage, experimental difficulties have limited the information available on the normal in vivo stresses and strains engendered in bone tissue. For more than 100 years, it has been assumed that bone structure and morphology reflect the demands placed on them by normal physical activities. As a result, their organization and mechanical properties vary dramatically as a function of anatomic location and physiologic condition 共i.e., age, gender, disease state兲 关22,23兴. On the other hand, the observed highly regulated structure/function properties suggest that its inherent capacity for adaptation would result in a relatively narrow range of normal in vivo strains. Numerous studies designed to measure bone surface strains are supportive of this principle, with reported values ranging between approximately 400 and 3000 microstrain 共tension or compression兲, across many animal species 关22兴. More recently, relatively similar measures have been documented in human subjects 关24兴. It is important to note that these studies are focused on cortical bone, since its location and surface properties enable the use of strain gauge devices. The trabecular bone compartment presents substantial experimental challenges, due to its anisotropic, macroscopically porous architecture. In vivo strains can only be estimated by analytical or computational methods. Most investigators have approached this challenge by developing finite element models of the bone region of interest. Boundary conditions are typically determined by measuring joint or muscle insertion loads 共i.e., 关25兴兲 and applying them in computational models that have mesh geometries that are derived from high resolution images of the trabecular bone structure of interest 共see 关26兴兲. Interestingly, the range of predicted strains in trabecular bone are similar to those measured in cortical bone. 2 The Mechanical Properties of the Native Tissues Must Be Established for Subfailure and Failure Conditions. To design effective tissue-engineered implants, it is important to understand the subfailure and failure properties of the native tissue 关27兴. Subfailure properties can be measured within the bounds of expected loading established above. Subfailure properties can be determined from viscoelastic experiments such as static or cyclic creep or stress relaxation testing. However, the native tissue should also be tested up to failure, especially if tissue engineered implants are to be designed with safety factors like the native tissue. Such failure testing provides both structural and material properties of the tissue to be replaced. ‘‘Structural’’ properties allow comparison of tissues or constructs to a baseline functional level, and incorporate the role of important morphological parameters, such as tissue geometry or joint congruence. ‘‘Material’’ properties are valuable in that they may be determined in simplified loading configurations, in combination with physiologically relevant theoretical models, but may be used to describe the mechanical response of a material to any loading history. Due to their complex structure and composition, most biological tissues can be classified from a material standpoint as inhomogeneous, viscoelastic, nonlinear, and anisotropic materials 共Table 1兲. The fundamental basis for these behaviors is not fully underDECEMBER 2000, Vol. 122 Õ 571

Table 1 Functional properties of natural and engineered tissues

stood, and may differ among different tissues. Importantly, it remains to be determined which aspects of these mechanical properties are essential for the normal, healthy function of different tissues, as well as for successful tissue-engineered replacements. Articular cartilage, for example, normally exhibits little or no wear with millions of cycles of loading that may reach ten times body weight. Its unique mechanical and tribological properties, which are unparalleled in man-made bearings, have been attributed to the complex structure and composition of extracellular tissue matrix 关28兴. In response to an externally applied load, articular cartilage is subjected to a complex state of tensile, shear and compressive stresses. Because of the large water content of the extracellular matrix 共75–85 percent兲, mechanical loading also results in pressure gradients in the interstitial fluid. As the extracellular matrix is permeable to water and possesses a significant amount of fixed negative charge 共due to the presence of proteoglycans兲, pressure gradients cause movement and redistribution of the interstitial fluid. Fluid movement may also be accompanied by electrokinetic effects such as streaming potentials and currents as various ions are moved through the charged matrix 关29,30兴. It is now well accepted that the primary mechanism of viscoelasticity in cartilage results from frictional interactions between the solid and fluid phases, although there is evidence that the solid matrix exhibits intrinsic viscoelasticity. Cartilage also exhibits highly nonlinear mechanical properties such as strain-dependent moduli 关31兴, strain-dependent hydraulic permeability 关32兴, and a difference of nearly two orders of magnitude in tensile and compressive moduli 关33兴. These properties are also anisotropic, particularly in tension, and vary significantly with distance from the tissue surface and with site on the joint 关34兴. More complex but equally important mechanical behaviors include the presence of internal swelling pressures that give rise to inhomogeneous residual 572 Õ Vol. 122, DECEMBER 2000

stresses within normal articular cartilage 关35兴. Finally, cartilage possesses important geometric and material characteristics that endow it with unique frictional properties. This low coefficient of friction, coupled with fluid-dependent mechanisms of load support, allow for minimal tissue wear under a relatively harsh mechanical environment 关36兴. Investigators have also characterized subfailure and failure properties in tendons and ligaments. Creep testing reveals a large initial elastic response followed by a small additional viscoelastic deformation over time 关27兴. Corresponding stress relaxation experiments demonstrate a rather substantial reduction in stress over time to about 40 to 60 percent of initial values 关37–39兴. A uniaxial failure test can then be performed to force/stress levels that originally compromised the tissue both as a structure and as a collagenous material 关27,40–42兴. The linear stiffness, maximum force, and maximum deformation at failure, and the energies to maximum force and failure 共areas under the curve兲 provide descriptions of the structural capacity of the tissue. Linear modulus, maximum stress, and maximum strain, and strain energy densities to maximum stress and failure can also be computed to provide estimates of the quality of the tissue as a material. These parameters follow directly from the corresponding structural properties but factor out the size of the tissue by dividing forces by initial area and deformations by initial length. Typical values for these material properties can vary dramatically depending upon whether strains are measured in the midsubstance or between the tissue ends. Bone mechanical properties can best be characterized as a function of hierarchical organization. At each level of hierarchy, a continuum mechanics approach to measuring bone properties reveals significant anisotropic 共nearly orthotropic兲 behavior 关43,44兴. From a physiological, functional point of view, the integrity of a bone region is dominated by its gross geometry and apparent density. Experimental data suggest that its preyield material properties are substantially dependent on its mineral 共or mineral/matrix兲 content, and post-yield or failure behavior is more influenced by its glycoprotein matrix 关43,44兴. Since bone can adapt substantially to its physical demands, if given sufficient time and early protection from excessive loads, it may attain required properties through a self-regulating process. As a result, the most important design parameter for a tissue engineered bone construct may be the biophysical mechanisms that enable effective remodeling/ adaptation to physical demand. 3 A Subset of These Mechanical Properties Must Be Selected and Prioritized. The abundance of data now available in the literature from biomechanical testing does not, however, answer one important question facing the tissue engineer. Which, if any, of these parameters should be used when designing an engineered repair or replacement? Realistically, it will be difficult, if not impossible, to match all of the material properties and structure of native tissues with tissue engineered constructs. Thus, it would be advantageous to prioritize the multitude of complex properties that are sought 共Table 1兲. Should the tissue-engineered construct match the failure properties of the normal tissue? Should it mimic the viscoelastic behavior of normal tissue? Should its properties be identical to adjacent host tissues, even if they are potentially not ‘‘normal’’? Or should the construct simply be designed with peak stresses and strains that are some fixed percentage of ultimate values? These issues may be equally applicable with respect to structural parameters as well as material properties. Should the regenerated tissue contain all of the load-bearing constituents of the normal, uninjured structure? Is proper organization required to replicate normal tissue function? Does the tissue in question require the same geometry, organization, and microstructure? How does integration between the graft and host tissue influence success? Clearly the choice of these parameters will depend on the tissue in question, but may involve similar themes or methodologies. The answers to many of these questions cannot be answered Transactions of the ASME

without further fundamental information on the biomechanical and biological properties of native and engineered tissues. Few studies have reported quantitative measures of the material properties of tissue-engineered cartilage. Of the few that have, focus has been placed on the compressive moduli and hydraulic permeability of cartilage 关45–48兴. While these properties are likely the most logical starting points, the relative importance of recreating the tissue’s compressive properties in comparison to tensile properties, failure properties, or frictional properties, for example, is not yet known. On a structural basis, many questions remain regarding the relative importance of recreating the native tissue and joint architecture. For example, most attempts at articular cartilage regeneration have sought complete integration between host and repair tissues 关49,50兴. Complete graft integration has been used as a ‘‘gold standard’’ of cartilage repair, yet the long-term implications of either complete or incomplete tissue integration are not fully understood. Additional factors such as the congruence of opposing cartilage surfaces in a joint may have important implications on the stress environment within the joint 关51兴. At this point, however, few tissue-engineering approaches are able to control the structure and geometry of newly formed cartilage precisely. An important consideration in such studies may also be the choice of an animal model and how representative the native tissue structure and properties are compared to the human 关52兴. Since all of these complex issues are unlikely to be addressed at once, it becomes important to prioritize their relative influence on the overall success of a given procedure. Choosing the most critical properties for a bone construct are somewhat different. From a long-term functional perspective, the most important property of the regenerating bone tissue will be its ability to remodel and adapt to habitual physical demands. Probably the most rapid way of achieving this goal is to design a construct with a high macroscopic porosity and surface-to-volume ratio and chemical properties that would promote osteoblast attachment and osteoclast resorption of its surfaces. However, this path of optimization would occur at the loss of initial mechanical integrity. Since the contruct could be augmented with many typical orthopaedic fracture fixation hardware devices, it provides designers an opportunity to select a balance between initial mechanical competence and rate of incorporation and replacement. Similar arguments arise when trying to prioritize structural and material properties for tendon and ligament replacement. Since these tissues carry primarily tensile forces, and only up to 10 percent 共ligament兲 to 40 percent of failure 共tendon兲, it would seem logical to try and replicate only the ‘‘toe’’ and early linear regions of the loading curve. Thus, toe limit stress and strain as well as tissue stiffness and modulus would seem to be the most important parameters to replicate in the tissue engineered replacement. However, these tissues are also repeatedly loaded under a combination of load and displacement control. Thus, cyclic creep and stress relaxation parameters are probably equally important characteristics to include in the tissue-engineered designs. The peak loads and displacements selected for these viscoelastic tests will change dramatically when new tissues are to be replaced. 4 Fourth, Standards Must Be Set When Evaluating the RepairsÕReplacements After Surgery So as to Determine, ‘‘How Good Is Good Enough’’? Assessment of the outcome of successful functional tissue engineering will require quantitative measures of graft properties, structure, and composition. Some aspects of the repair outcome may be inferior, but other mechanical factors of the repairs and replacements might be suitable. With an emphasis on the material properties and structure of tissue-engineered grafts, it will be necessary to quantify and report outcome measures directly related to the functional behavior of the tissues. Given the biomechanical nature of many tissueengineered products, there have been surprisingly few reports of the material or structural properties of engineered tissues. In articular cartilage, for example, several investigators have reported Journal of Biomechanical Engineering

either mechanical properties of grafts prior to implantation 关47兴 or at sacrifice 关45,46,53兴. In tendon, cell-based repairs using autogenous mesenchymal stem cells have resulted in 20 to 30 percent improvement in material properties when cells are not organized in culture 关54兴. When these cells are aligned in culture 关55,56兴, the structural and material properties of these repairs have been found to be 100 percent better than natural repairs of unfilled defects 关57兴. However, those conditions that optimize the repair outcome for tendon or ligament repair are still not known. An important direction for the field will be the development of new methodologies that will allow assessment of the material or structural properties of engineered tissues in a non-invasive or minimally invasive manner. For example, the use of biological markers of tissue metabolism 关58兴, in vivo 共arthroscopic兲 biomechanical probes 关59兴, magnetic resonance imaging 关60兴, CT, ultrasound, DEXA 关61兴, and other techniques to assess tissue function may prove to be critical in longitudinal studies of tissue engineered repair, particularly in the clinical setting. 5 What Physical Regulation Do Cells Experience in Vivo as They Interact With an Extracellular Matrix? Once implanted in the body, engineered constructs of cells and/or matrices will be subjected to a complex biomechanical environment, potentially consisting of time-varying changes in stresses, strains, fluid pressure, fluid flow, and cellular deformation 关29,30,62兴. It is now well accepted that these various physical factors have the capability to influence the biological activity of normal tissues 共reviewed in 关63兴兲, and therefore, may play an important role in the eventual success or failure of engineered grafts. In this regard, it would be important to better characterize the diverse array of physical signals that engineered cells may experience in vivo, as well as their biological response to such potential stimuli. This information may provide important insights into the long-term capabilities of engineered constructs to maintain the proper cellular phenotype and may shorten the iterations required to successfully replace a tissue. The cells used in the tissue-engineered construct or those that are recruited or migrate to the construct must be organized and stimulated to rapidly synthesize the desired extracellular matrix. The scaffold can play a significant role in influencing the behavior and ultimately the phenotype of the cells. Whether the cells are grown on the scaffold ex vivo 共e.g., 关54,57兴兲, or migrate to it, in vivo, their interaction with the scaffold material may be a critical factor in determining success. The chemical, mechanical, and architectural properties of the scaffold will affect the number, phenotype, and adherence properties of the cells. As a result, the mechanical influence on the cells will be related to the mechanical properties of the scaffold, the boundary loading conditions acting on the surgically placed construct, and the number and quality of focal adhesion contacts of the cell to the scaffold. In addition, the shape and morphology of the cells will be related to the cell/ scaffold interactions. All of these factors will contribute to the cell’s ability to respond to both mechanical and biologic signals, and subsequently to synthesize and express extracellular matrix 关64–67兴. 6 Finally, How Do Physical Factors Influence Cellular Activity in Bioreactors and Can Cell-Matrix Implants Be Mechanically Stimulated Before Surgery to Produce a Better Outcome? Mechanical stress is an important modulator of cell physiology, and there is significant evidence that physical factors may be used to improve or accelerate tissue regeneration and repair in vitro. For example, early studies showed that cyclic mechanical stretch of skeletal myofibers increased the alignment of myotubes that assembled into ‘‘organoids’’ in culture 关68兴. In other studies, mechanical stretch has been shown to increase cellular alignment, proliferation, and protein synthesis in many different cell types 关69,70兴. As the cells used in the tissue-engineered construct must be organized to synthesize the desired extracellular DECEMBER 2000, Vol. 122 Õ 573

matrix rapidly, control of cellular alignment a priori may provide important advantages in controlling matrix deposition, and presumably, a more rapid development of functional biomechanical properties. More recently, mechanical ‘‘bioreactors’’ have been used to increase matrix deposition in tissue-engineered cartilage by exposing chondrocytes to fluid flow 关71兴, simulated hypogravity 关48兴, and cyclic compression 关72,73兴. Recent studies have shown improved success of tissue-engineered systems such as blood vessels by preconditioning grafts with pulsatile fluid flow and pressure 关74兴. Cells even align when cell–collagen composites are contracted onto suture exposed to a static load 关55兴. Mesenchymal stem cells align within about 24 hours in culture. The cells align more rapidly when cell density is increased, although a threshold of cell density exists above which no change in the rate or extent of alignment occurs. The nuclei of these cells also become more spindle-shaped with increasing cell density 关55兴. However, little is known about how cyclic loading affects cell shape and alignment and mechanical loads influence cell proliferation, differentiation, and protein expression under these conditions.

Future Directions Clearly, the field of tissue engineering needs to establish functional criteria that will help those who seek to design and manufacture these repairs and replacements. Conferences and workshops must be organized to bring together experts from academia, industry, and government to agree on functional 共structural and mechanical兲 requirements for important load-bearing tissues. Expanded funding in this area will also be needed to help basic and applied researchers establish these criteria so that new and more innovative tissue engineered repairs and replacements can be provided to the surgeon and patient. Scale-up, packaging, storage, and handling properties are also critical. The implants must be capable of retaining their mechanical, structural, and biological integrity during large-scale production, packaging, and storage. If the tissue-engineering construct is too fragile or difficult to handle in the operating room, surgeons will likely not use the devices in patients. Understanding those conditions that preserve the character of the implants will be essential for the success of tissue-engineered products. With rapidly evolving new technologies being developed, the future of tissue engineering for tendon, ligament, cartilage, and bone repair should be quite bright. Growth factors, introduced during fabrication or delivery of the constructs, offer the promise of further improving repair quality. However, these factors must be temporally and spatially delivered appropriately if tissue engineers expect to stimulate early cell proliferation and subsequent matrix synthesis. This matrix must also be deposited and rapidly organized if the repair is to be capable of resisting the large in vivo forces of daily activities. The organization and orientation of these matrices might also be further enhanced through strategic mechanical stimulation of the constructs during the in vitro fabrication process or early post-surgical periods. Providing either controlled strain or load signals to the constructs will likely be beneficial by instructing and organizing the cells in a way in which they synthesize an appropriately aligned matrix, just as loads applied to the suture passing through the construct aligned the MSCs suspended in a collagen gel 关54,55兴. However, the load 共strain兲 levels that must be chosen and their frequencies must still be determined and are likely to be tissue and site specific. Novel matrices must also be identified with enough compliance so the cells can organize the construct but with enough stiffness to resist the expected in vivo loading regimes. These matrices may contain preformed organic or synthetic fibers that can initially sustain in vivo loading. These matrices should initially attract, anchor and protect the cells, but then degrade at controlled rates that prevent mechanical disruption of the repair and ensure biocompatibility. 574 Õ Vol. 122, DECEMBER 2000

Many of these treatments are now being investigated and offer real promise for the repair of problematic soft and hard tissue injuries.

References 关1兴 Biewener, A. A., Blickhan, R., Perry, A. K., Heglund, N. C., and Taylor, C. R., 1988, ‘‘Muscle Forces During Locomotion in Kangaroo Rats: Force Platform and Tendon Buckle Measurements Compared,’’ J. Exp. Biol., 137, pp. 191– 205. 关2兴 Lewis, J. L., Lew, W. D., and Schmidt, J., 1982, ‘‘A Note on the Application and Evaluation of the Buckle Transducer for Knee Ligament Force Measurement,’’ ASME J. Biomech. Eng., 104, pp. 125–128. 关3兴 Komi, P. V., 1990, ‘‘Relevance of In Vivo Force Measurements to Human Biomechanics,’’ J. Biomech., 23, pp. 23–34. 关4兴 Komi, P. V., Fukashiro, S., and Jarvinen, M., 1992, ‘‘Biomechanical Loading of Achilles Tendon During Normal Locomotion,’’ Clin. Sports Med., 11, pp. 521–531. 关5兴 Cummings, J. F., Holden, J. P., Grood, E. S., Wroble, R. R., Butler, D. L., and Schafer, J., 1992, ‘‘In-Vivo Measurement of Patellar Tendon Forces and Joint Position in the Goat Model,’’ Trans. 37th Orthopaedic Res. Soc., 15, p. 601. 关6兴 Glos, D. L., Butler, D. L., Grood, E. S., and Levy, M. S., 1993, ‘‘In Vitro Evaluation of an Implantable Force Transducer 共IFT兲 in a Patellar Tendon Model,’’ ASME J. Biomech. Eng., 115, pp. 335–343. 关7兴 Herzog, W., Leonard, T. R., Renaud, J. M., Wallace, J., Chaki, G., and Bornemisza, S., 1992, ‘‘Force-Length Properties and Functional Demands of Cat Gastrocnemius, Soleus, and Plantaris Muscles,’’ J. Biomech., 25, pp. 1329– 1335. 关8兴 Holden, J. P., Grood, E. S., Korvick, D. L., Cummings, J. F., Butler, D. L., and Bylski-Austrow, D. I., 1994, ‘‘In Vivo Forces in the Anterior Cruciate Ligament: Direct Measurements During Walking and Trotting in a Quadruped,’’ J. Biomech., 27, pp. 517–526. 关9兴 Korvick, D. L., Cummings, J. F., Grood, E. S., Holden, J. P., Feder, S. M., and Butler, D. L., 1996, ‘‘The Use of an Implantable Force Transducer to Measure Patellar Tendon Forces in Goats,’’ J. Biomech., 29, pp. 557–561. 关10兴 Ronsky, J. L., Herzog, W., Brown, T. D., Pedersen, D. R., Grood, E. S., and Butler, D. L., 1995, ‘‘In Vivo Quantification of the Cat Patellofemoral Joint Contact Stresses and Areas,’’ J. Biomech., 28, pp. 977–983. 关11兴 Xu, W. S., Butler, D. L., Stouffer, D. C., Grood, E. S., and Glos, D. L., 1992, ‘‘Theoretical Analysis of an Implantable Force Transducer for Tendon and Ligament Structures,’’ ASME J. Biomech. Eng., 114, pp. 170–177. 关12兴 Korvick, D. L., Holden, J. P., Grood, E. S., Cummings, J. F., and Rupert, M. P., 1992, ‘‘Relationships Between Patellar Tendon, Anterior Cruciate Ligament and Vertical Ground Reaction Force During Gait: Preliminary Studies in a Quadruped,’’ Advances in Bioengineering, ASME BED-Vol. 22, pp. 99– 102. 关13兴 Malaviya, P., Butler, D. L., Korvick, D. L., and Proch, F. S., 1998, ‘‘In Vivo Tendon Forces: Do They Correlate with Activity Level and Remain Bounded? Evidence in a Rabbit Flexor Tendon Model,’’ J. Biomech., 31, pp. 1043–1049. 关14兴 Livesay, G. A., Rudy, T. W., Woo, S. L., Runco, T. J., Sakane, M., Li, G., and Fu, F. H., 1997, ‘‘Evaluation of the Effect of Joint Constraints on the In Situ Force Distribution in the Anterior Cruciate Ligament,’’ J. Orthop. Res., 15, pp. 278–284. 关15兴 Sakane, M., Fox, R. J., Woo, S. L., Livesay, G. A., Li, G., and Fu, F. H., 1997, ‘‘In Situ Forces in the Anterior Cruciate Ligament and Its Bundles in Response to Anterior Tibial Loads,’’ J. Orthop. Res., 15, pp. 285–293. 关16兴 Beynnon, B. D., Ryder, S. H., Konradsen, L., Johnson, R. J., Johnson, K., and Renstrom, P. A., 1999, ‘‘The Effect of Anterior Cruciate Ligament Trauma and Bracing on Knee Proprioception,’’ Am. J. Sports Med., 27, pp. 150–155. 关17兴 Fleming, B. C., Beynnon, B. D., Renstrom, P. A., Johnson, R. J., Nichols, C. E., Peura, G. D., and Uh, B. S., 1999, ‘‘The Strain Behavior of the Anterior Cruciate Ligament During Stair Climbing: An In Vivo Study,’’ Arthroscopy, 15, pp. 185–191. 关18兴 Hodge, W. A., Carlson, K. L., Fijan, R. S., Burgess, R. G., Riley, P. O., Harris, W. H., and Mann, R. W., 1989, ‘‘Contact Pressures From an Instrumented Hip Endoprosthesis,’’ J. Bone Jt. Surg., 71A, pp. 1378–1386. 关19兴 Nelson, B. H., Anderson, D. D., Brand, R. A., and Brown, T. D., 1988, ‘‘Effect of Osteochondral Defects on Articular Cartilage. Contact Pressures Studied in Dog Knees,’’ Acta Orthop. Scand., 59, pp. 574–579. 关20兴 Armstrong, C. G., Bahrani, A. S., and Gardner, D. L., 1979, ‘‘In Vitro Measurement of Articular Cartilage Deformations in the Intact Human Hip Joint Under Load,’’ J. Bone Jt. Surg., Am. Vol., 61, pp. 744–755. 关21兴 Donzelli, P. S., Spilker, R. L., Ateshian, G. A., and Mow, V. C., 1999, ‘‘Contact Analysis of Biphasic Transversely Isotropic Cartilage Layers and Correlations With Tissue Failure,’’ J. Biomech., 32, pp. 1037–1047. 关22兴 Martin, R. B., and Burr, D. B., 1989, Structure, Function and Adaptation of Compact Bone, Raven Press, NY, pp. 143–185. 关23兴 Goldstein, S. A., Hollister, S. J., Kuhn, J. L., and Kikuchi, N., 1990, ‘‘The Mechanical and Remodeling Properties of Trabecular Bone,’’ Biomechanics of Diarthrodial Joints, V. C. Mow, A. Ratcliffe, and S. L.-Y. Woo, eds., Springer Verlag, NY, pp. 61–81. 关24兴 Burr, D. B., Milgrom, C., Fyhrie, D., Forwood, M., Nyska, M., Finestone, A., Hoshaw, S., Saiag, E., and Simkin, A., 1996, ‘‘In Vivo Measurement of Human Tibial Strains During Vigorous Activity,’’ Bone, 18, pp. 405–410.

Transactions of the ASME

关25兴 An, K. N., Chao, E. Y. S., and Kaufman, K. R., 1991, ‘‘Analysis of Muscle and Joint Loads,’’ Basic Orthopaedic Biomechanics, V. C. Mow and W. C. Hayes, eds., Raven Press, New York, pp. 1–50. 关26兴 Huiskes, R., and Hollister, S. J., 1990, ‘‘From Structure to Process, From Organ to Cell: Recent Developments of FE-Analysis in Orthopaedic Biomechanics,’’ ASME J. Biomech. Eng., 115, pp. 520–527. 关27兴 Rasmussen, T. J., Feder, S. M., Butler, D. L., and Noyes, F. R., 1994, ‘‘The Effects of 4Mrad Gamma Irradiation Sterilization on the Initial Structural Properties of ACL and PCL Patellar Tendon Allografts,’’ J. Arthroscopic Related Surgery, 10, pp. 188–197. 关28兴 Mow, V. C., Ratcliffe, A., and Poole, A. R., 1992, ‘‘Cartilage and Diarthrodial Joints as Paradigms for Hierarchical Materials and Structures,’’ Biomaterials, 13, pp. 67–97. 关29兴 Lai, W. M., Hou, J. S., and Mow, V. C., 1991, ‘‘A Triphasic Theory for the Swelling and Deformation Behaviors of Articular Cartilage,’’ ASME J. Biomech. Eng., 113, pp. 245–258. 关30兴 Frank, E. H., and Grodzinsky, A. J., 1987, ‘‘Cartilage Electromechanics—II. A Continuum Model of Cartilage Electrokinetics and Correlation With Experiments,’’ J. Biomech., 20, pp. 629–639. 关31兴 Kwan, M. K., Lai, W. M., and Mow, V. C., 1990, ‘‘A Finite Deformation Theory for Cartilage and Other Soft Hydrated Connective Tissues—I. Equilibrium Results,’’ J. Biomech., 23, pp. 145–155. 关32兴 Lai, W. M., Mow, V. C., and Roth, V., 1981, ‘‘Effects of Nonlinear StrainDependent Permeability and Rate of Compression on the Stress Behavior of Articular Cartilage,’’ ASME J. Biomech. Eng., 103, pp. 61–66. 关33兴 Ateshian, G. A., and Soltz, M. A., 1999, ‘‘A Biphasic Conewise Linear Elasticity Model for Modeling Tension-Compression Nonlinearity in Articular Cartilage,’’ in: Cartilaginous Tissue Mechanics, ASME BED-Vol. 42, pp. 69– 70. 关34兴 Akizuki, S., Mow, V. C., Muller, F., Pita, J. C., Howell, D. S., and Manicourt, D. H., 1986, ‘‘Tensile Properties of Human Knee Joint Cartilage: I. Influence of Ionic Conditions, Weight Bearing, and Fibrillation on the Tensile Modulus,’’ J. Orthop. Res., 4, pp. 379–392. 关35兴 Setton, L. A., Tohyama, H., and Mow, V. C., 1998, ‘‘Swelling and Curling Behaviors of Articular Cartilage,’’ ASME J. Biomech. Eng., 120, pp. 355– 361. 关36兴 Mow, V. C., and Ateshian, G. A., 1997, ‘‘Lubrication and Wear of Diarthrodial Joints,’’ in: Basic Orthopaedic Biomechanics, V. C. Mow and W. C. Hayes, eds., Philadelphia, Lippincott-Raven, pp. 275–315. 关37兴 Lam, T. C., Frank, C. B., and Shrive, N. G., 1993, ‘‘Changes in the Cyclic and Static Relaxations of the Rabbit Medial Collateral Ligament Complex During Maturation,’’ J. Biomech., 26, pp. 9–17. 关38兴 Chimich, D., Shrive, N., Frank, C. B., Marchuk, L., and Bray, R., 1992, ‘‘Water Content Alters Viscoelastic Behavior of the Normal Adolescent Rabbit Medial Collateral Ligament,’’ J. Biomech., 25, pp. 831–837. 关39兴 Kwan, M. K., Lin, T. H., and Woo, S. L., 1993, ‘‘On the Viscoelastic Properties of the Anteromedial Bundle of the Anterior Cruciate Ligament,’’ J. Biomech., 26, pp. 447–452. 关40兴 Butler, D. L., Noyes, F. R., and Grood, E. S., 1978, ‘‘Measurement of the Biomechanical Properties of Ligaments,’’ in: CRC Handbook on Engineering and Biology, G. Bahriuk and A. Burstein, eds., Section B, Vol. 1, pp. 279– 314. 关41兴 Butler, D. L., Grood, E. S., Noyes, F. R., Zernicke, R. F., and Brackett, K., 1984, ‘‘Effects of Structure and Strain Measurement Technique on the Material Properties of Young Human Tendons and Fascia,’’ J. Biomech., 17, pp. 579–596. 关42兴 Butler, D. L., Kay, M. D., and Stouffer, D. C., 1986, ‘‘Comparison of Material Properties in Fascicle–Bone Units From Human Patellar Tendon and Knee Ligaments,’’ J. Biomech., 19, pp. 425–432. 关43兴 Cowin, S. C., 1989, Bone Mechanics, CRC Press, Boca Raton, FL, pp. 97– 158. 关44兴 Hoffler, C. E., McCreadie, B. R., Smith, E. A., and Goldstein, S. A., 2000, ‘‘A Hierarchical Approach to Exploring Bone Mechanical Properties,’’ in: Mechanical Testing of Bone and the Bone-Implant Interface, Y. H. An, and R. A. Draughn, eds., CRC Press, Boca Raton, FL, pp. 133–150. 关45兴 Kwan, M. K., Coutts, R. D., Woo, S. L., and Field, F. P., 1989, ‘‘Morphological and Biomechanical Evaluations of Neocartilage From the Repair of FullThickness Articular Cartilage Defects Using Rib Perichondrium Autografts: A Long-Term Study,’’ J. Biomech., 22, pp. 921–930. 关46兴 Mow, V. C., Ratcliffe, A., Rosenwasser, M. P., and Buckwalter, J. A., 1991, ‘‘Experimental Studies on Repair of Large Osteochondral Defects at a High Weight Bearing Area of the Knee Joint: A Tissue Engineering Study,’’ ASME J. Biomech. Eng., 113, pp. 198–207. 关47兴 Vunjak-Novakovic, G., Martin, I., Obradovic, B., Treppo, S., Grodzinsky, A. J., Langer, R., and Freed, L. E., 1999, ‘‘Bioreactor Cultivation Conditions Modulate the Composition and Mechanical Properties of Tissue-Engineered Cartilage,’’ J. Orthop. Res., 17, pp. 130–138. 关48兴 Freed, L. E., Langer, R., Martin, I., Pellis, N. R., and Vunjak-Novakovic, G., 1997, ‘‘Tissue Engineering of Cartilage in Space,’’ Proc. Natl. Acad. Sci. U.S.A., 94, pp. 13885–13890.

Journal of Biomechanical Engineering

关49兴 Hunziker, E. B., and Rosenberg, L. C., 1996, ‘‘Repair of Partial-Thickness Defects in Articular Cartilage: Cell Recruitment From the Synovial Membrane,’’ J. Bone Jt. Surg., 78, pp. 721–733. 关50兴 Ahsan, T., and Sah, R. L., 1999, ‘‘Biomechanics of Integrative Cartilage Repair,’’ Osteoarthritis Cartilage, 7, pp. 29–40. 关51兴 Ateshian, G. A., Rosenwasser, M. P., and Mow, V. C., 1992, ‘‘Curvature Characteristics and Congruence of the Thumb Carpometacarpal Joint: Differences Between Female and Male Joints,’’ J. Biomech., 25, pp. 591–607. 关52兴 Hunziker, E. B., 1999, ‘‘Biologic Repair of Articular Cartilage. Defect Models in Experimental Animals and Matrix Requirements,’’ Clin. Orthop., 367, pp. S135–146. 关53兴 Wakitani, S., Goto, T., Pineda, S. J., Young, R. G., Mansour, J. M., Caplan, A. I., and Goldberg, V. M., 1994, ‘‘Mesenchymal Cell-Based Repair of Large, Full-Thickness Defects of Articular Cartilage,’’ J. Bone Jt. Surg., 76, pp. 579– 592. 关54兴 Awad, H. A., Butler, D. L., Boivin, G. P., Smith, F., Malaviya, P., Huibregtse, B., Caplan, A. I., 1999, ‘‘Autologous Mesenchymal Stem Cell-Mediated Repair of Tendon,’’ Tissue Eng., 5, pp. 267–277. 关55兴 Awad, H. A. et al., 2000, ‘‘In Vitro Characterization of Mesenchymal Stem Cell-Seeded Collagen Scaffolds for Tendon Repair: Effects of Initial Seeding Density on Contraction Kinetics,’’ J. Biomed. Mater. Res., in press. 关56兴 Butler, D. L., and Awad, H. A., 1999, ‘‘Perspectives on Cell and Collagen Composites for Tendon Repair,’’ CORR, 367S, pp. S324–S332. 关57兴 Young, R., Butler, D., Weber, W., Caplan, A., Gordon, S., and Fink, D., 1998, ‘‘Use of Mesenchymal Stem Cells in a Collagen Matrix for Achilles Tendon Repair,’’ J. Orthop. Res., 16, pp. 406–413. 关58兴 Lohmander, L. S., and Felson, D. T., 1997, ‘‘Defining the Role of Molecular Markers to Monitor Disease, Intervention, and Cartilage Breakdown in Osteoarthritis,’’ J. Rheumatol., 24, pp. 782–785. 关59兴 Lyyra, T., Jurvelin, J., Pitkanen, P., Vaatainen, U., and Kiviranta, I., 1995, ‘‘Indentation Instrument for the Measurement of Cartilage Stiffness Under Arthroscopic Control,’’ Med. Eng. Phys., 17, pp. 395–399. 关60兴 Karvonen, R. L., Negendank, W. G., Fraser, S. M., Mayes, M. D., An, T., and Fernandez Madrid, F., 1990, ‘‘Articular Cartilage Defects of the Knee: Correlation Between Magnetic Resonance Imaging and Gross Pathology,’’ Ann. Rheum. Dis., 49, pp. 672–675. 关61兴 Miller, P. D., and Bonnick, S. L., 1999, ‘‘Clinical Application of Bone Densitometry,’’ in: Primer on the Metabolic Bone Diseases and Disorders of Mineral Metabolism, 4th ed., M. J. Favus, ed., Lippincott Williams & Wilkins, Philadelphia, pp. 152–159. 关62兴 Guilak, F., Ratcliffe, A., and Mow, V. C., 1995, ‘‘Chondrocyte Deformation and Local Tissue Strain in Articular Cartilage: A Confocal Microscopy Study,’’ J. Orthop. Res., 13, pp. 410–421. 关63兴 Guilak, F., Sah, R. L., and Setton, L. A., 1997, ‘‘Physical Regulation of Cartilage Metabolism,’’ in: Basic Orthopaedic Biomechanics, V. C. Mow and W. C. Hayes, eds., Philadelphia, Lippincott Raven, pp. 179–207. 关64兴 Goldstein, S. A., Patil, P. V., and Moalli, M. R., 1999, ‘‘Perspectives on Tissue Engineering in Bone,’’ CORR, 367S, pp. S419–S423. 关65兴 Guilak, F., and Mow, V. C., 1992, ‘‘Determination of the Mechanical Response of the Chondrocyte In Situ Using Finite Element Modeling and Confocal Microscopy,’’ Advances in Bioengineering, ASME BED-Vol. 20, pp. 21–23. 关66兴 Caplan, A. I., and Bruder, S. P., 1997, ‘‘Cell and Molecular Engineering of Bone Regeneration,’’ in: Principles of Tissue Engineering, R. Lanza, R. Langer, and W. Chick, eds., R. G. Landes Company, Chap. 37, pp. 603–618. 关67兴 Ishaug-Riley, S. L., Crane-Kruger, G. M., Yaszemski, M. J., and Mikos, A. G., 1998, ‘‘Three-Dimensional Culture of Rat Calvarial Osteoblasts in Porous Biodegradable Polymers,’’ Biomaterials, 19, pp. 1405–1412. 关68兴 Vandenburgh, H. H., 1982, ‘‘Dynamic Mechanical Orientation of Skeletal Myofibers In Vitro,’’ Dev. Biol., 93, pp. 438–443. 关69兴 Buckley, M. J., Banes, A. J., Levin, L. G., Sumpio, B. E., Sato, M., Jordan, R., Gilbert, J., Link, G. W., and Tran Son Tay, R., 1988, ‘‘Osteoblasts Increase Their Rate of Division and Align in Response to Cyclic, Mechanical Tension in Vitro,’’ Bone Miner., 4, pp. 225–236. 关70兴 Sumpio, B. E., Banes, A. J., Link, W. G., and Johnson, Jr., G., 1988, ‘‘Enhanced Collagen Production by Smooth Muscle Cells During Repetitive Mechanical Stretching,’’ Arch. Surg., 123, pp. 1233–1236. 关71兴 Wu, F., Dunkelman, N., Peterson, A., Davisson, T., De La Torre, R., and Jain, D., 1999, ‘‘Bioreactor Development for Tissue-Engineered Cartilage,’’ Ann. N.Y. Acad. Sci., 875, pp. 405–411. 关72兴 Buschmann, M. D., Gluzband, Y. A., Grodzinsky, A. J., and Hunziker, E. B., 1995, ‘‘Mechanical Compression Modulates Matrix Biosynthesis in Chondrocyte/Agarose Culture,’’ J. Cell. Sci., 108, pp. 1497–1508. 关73兴 Mauck, R. L., Soltz, M. A., Wang, C. B., Wong, D. D., Chao, P. G., Valhmu, W. B., Hung, C. T., and Ateshian, G. A., 2000, ‘‘Functional Tissue Engineering of Articular Cartilage Through Dynamic Loading of Chondrocyte-Seeded Agarose Gels,’’ ASME J. Biomech. Eng., 122, pp. 252–260. 关74兴 Niklason, L. E., Gao, J., Abbott, W. M., Hirschi, K. K., Houser, S., Marini, R., and Langer, R., 1999, ‘‘Functional Arteries Grown in Vitro,’’ Science, 284, pp. 489–493.

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