Advanced Tools for Tissue Engineering: Scaffolds

TISSUE ENGINEERING Volume 12, Number 12, 2006 # Mary Ann Liebert, Inc.

Advanced Tools for Tissue Engineering: Scaffolds, Bioreactors, and Signaling LISA E. FREED,1 FARSHID GUILAK,2 X. EDWARD GUO,3 MARTHA L. GRAY,1 ROBERT TRANQUILLO,4 JEFFREY W. HOLMES,3 MILICA RADISIC,5 MICHAEL V. SEFTON,5 DAVID KAPLAN,6 and GORDANA VUNJAK-NOVAKOVIC3

ABSTRACT This article contains the collective views expressed at the second session of the workshop ‘‘Tissue Engineering—The Next Generation,’’ which was devoted to the tools of tissue engineering: scaffolds, bioreactors, and molecular and physical signaling. Lisa E. Freed and Farshid Guilak discussed the integrated use of scaffolds and bioreactors as tools to accelerate and control tissue regeneration, in the context of engineering mechanically functional cartilage and cardiac muscle. Edward Guo focused on the opportunities that tissue engineering generates for studies of mechanobiology and on the need for tissue engineers to learn about mechanical forces during tissue and organ genesis. Martha L. Gray focused on the potential of biomedical imaging for noninvasive monitoring of engineered tissues and on the opportunities biomedical imaging can generate for the development of new markers. Robert Tranquillo reviewed the approach to tissue engineering of a spectrum of avascular habitually loaded tissues— blood vessels, heart valves, ligaments, tendons, cartilage, and skin. Jeffrey W. Holmes offered the perspective of a ‘‘reverse paradigm’’—the use of tissue constructs in quantitative studies of cell-matrix interactions, cell mechanics, matrix mechanics, and mechanobiology. Milica Radisic discussed biomimetic design of tissue-engineering systems, on the example of synchronously contractile cardiac muscle. Michael V. Sefton proposed a new, simple approach to the vascularization of engineered tissues. This session stressed the need for advanced scaffolds, bioreactors, and imaging technologies and offered many enlightening examples on how these advanced tools can be utilized for functional tissue engineering and basic research in medicine and biology.

INTRODUCTION

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are certainly at the core of all tissue-engineering strategies. Smart scaffolds and bioreactors are being developed to enable advanced regimens for delivery of multiple growth factors and IOREACTORS, SCAFFOLDS, AND SIGNALING

mechanical signals to cultured cells in order to direct their differentiation toward functional tissue outcomes. To monitor tissue development in real time and thereby enhance our ability for efficient design of experiments and optimization of culture parameters, nondestructive imaging technologies and imaging biomarkers are being increasingly

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Harvard-MIT Division of Health Sciences and Technology, Massachusetts Institute of Technology, Cambridge, Massachusetts. Duke University Medical Center, Durham, North Carolina. 3 Department of Biomedical Engineering, Columbia University, New York, New York. 4 Department of Biomedical Engineering, University of Minnesota, Minneapolis, Minnesota. 5 Institute of Biomaterials and Biomedical Engineering and Department of Chemical Engineering and Applied Chemistry, University of Toronto, Toronto, Canada. 6 Department of Biomedical Engineering, Tufts University, Medford, Massachusetts. 2

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utilized. A control-based approach to bioreactor operation would provide a rational basis for the structural and functional design of engineering tissues for eventual clinical application. The immediate availability of engineered tissues for use as model systems for biological and medical research offers advantages of getting more realistic answers to a wide range of biological questions, and serves as a feedback from tissue engineering to biology.

THE USE OF SCAFFOLDS AND BIOREACTORS AS TISSUE-ENGINEERING TOOLS Lisa E. Freed from Massachusetts Institute of Technology and Farshid Guilak from Duke University discussed the use of scaffolds and bioreactors as tools for studying, accelerating, and better controlling tissue regeneration. They argued that the next generation of functional tissue replacements may require additional exogenous influences to achieve many of the important requirements for long-term success. Two such needs—physiologic mechanical properties, and the ability to develop and remodel tissues in a manner that allows for restoration of physiological function—can potentially be studied and achieved through the use of ‘‘smart’’ scaffolds and bioreactors. Given the rapid evolution of the field of tissue engineering, it is important to consider the use of smart scaffolds and bioreactors as additional tools in light of the role of novel biomaterials, growth factors, genetic modifications, and other emerging technologies. The focus was on representative examples of engineered skeletal and cardiac tissue constructs in which scaffold architectures were used to improve construct biomechanical properties, and bioreactors were used to study and accelerate construct maturation and integration. The loss of function of these tissues with injury, disease, or aging accounts for a significant number of clinical disorders at a tremendous social and economic cost.1,2 Despite many early successes, there are few engineered tissue products available for clinical use, and significant challenges still remain in the successful long-term repair of biomechanically functional tissues. The precise reasons for graft failure in experimental animal studies and preclinical trials are not fully understood, but include a combination of biological and mechanical factors that can lead to the breakdown of repair tissues under physiologic loading conditions. The magnitudes of stresses and frequency of loading that tissues may be subjected to in vivo can be quite large, and few engineered tissue constructs possess the biomechanical properties to withstand such stresses at the time of implantation. For example, articular cartilage may experience stresses of up to 18 MPa for approximately 1 million loading cycles per year.3 Furthermore, the challenge is not as simple as matching a single mechanical parameter, such as modulus or strength; rather, most tissues possess complex viscoelastic, nonlinear, and anisotropic mechanical and

physicochemical properties that may vary with age, site, and other factors. Finally, a number of complex interactions must be considered, as the graft and surrounding host tissues are expected to grow and remodel in response to their changing environments postimplantation (e.g., see Badylak et al.4). Therefore, the long-term success of a cell-based tissueengineered construct depends on its ability to appropriately respond to biological and biomechanical signals. A new and evolving discipline termed ‘‘functional tissue engineering’’ has sought to address these challenges by developing guidelines for rationally investigating the role of biological and biomechanical factors in tissue engineering. A series of formal goals and principles for functional tissue engineering have been proposed in a generalized format5,6 as well as with specific reference to articular cartilage.7

The role of scaffolds Biomaterial scaffolds provide a critical means of controlling engineered tissue architecture and mechanical properties. Many studies have employed scaffolds in which macro- to microscale features are present in isotropic spatial orientation. To cite two examples, (i) textile fabrication technologies were used to generate a nonwoven mesh made of fibers with diameters similar to cells (10–20 mm) and very high void volumes (up to 97%) and (ii) solvent-casting and saltleaching technologies were used to generate foams made of interconnected pores with a wide range of diameters (100– 500 mm) and void volumes of up to 90%.8,9 However, biomechanical properties and low mechanical strength of isotropic porous scaffolds may not be suitable for the repair of anisotropic, load-bearing tissues in some clinical applications.10 In an attempt to recreate the complex, multiphasic mechanical properties of native articular cartilage, computercontrolled weaving technologies were used to recreate anisotropy, viscoelasticity, and tension-compression nonlinearity of the extracellular matrix (ECM).11 A novel microscale three-dimensional (3D) technique was used to weave polyglycolic acid (PGA) yarns into an orthotropic, porous textile (Fig. 1A), which was then infiltrated with a hydrogel. The scaffold displayed significant tensioncompression nonlinearity, with approximately three orders of magnitude difference in tensile and compressive moduli, an aggregate modulus of *0.2 MPa and a tensile modulus of *300 MPa. Significant anisotropy was observed in the failure stress, failure strain, tangent modulus, and energyto-failure (e.g., failure stress was *35% higher and tangent modulus *50% higher in the weft direction than the warp direction). These findings show that 3D woven scaffolds can be designed with anisotropic, nonlinear, and viscoelastic properties similar to those of articular cartilage, even in the absence of cells. In the present design, the 3D fabric provides tensile properties and the potential for anisotropy, and a hydrogel is used during cell seeding, to consolidate the structure and allow chondrocytes to retain their rounded

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FIG. 1. Scaffolds. (A) Woven anisotropic 3D woven fabric with multiple layers of orthogonally oriented fibers for articular cartilage repair: scanning electron microscopy (SEM), scale bar 1 mm; histology of a 6-week construct based on chondrocytes stained with safranin-O, scale bar 500 mm. (B) Knitted anisotropic elastomeric fabric for myocardial repair: SEM, scale bar 500 mm; histology of a freshly seeded construct based on cardiomyocytes immunostained for cardiac troponin-I, scale bar 500 mm. (Images A and B courtesy of F.T. Moutos, M. Moretti, L.E. Freed, and F. Guilak.) (C) Composite construct for osteochondral defect repair: 6-week (upper image) and 6-month (lower image) explants from knee joints of adult rabbits, stained with Alcian blue, scale bars 5.0 mm. Dashes outline the original defect. Color images available online at www.liebertpub.com/ten.

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morphology and, presumably, their differentiated phenotype during in vitro culture (Fig. 1A). The large number of variables in this design (selection of fiber material, number of fibers up to several hundred, density, and packing in all directions) provides a wide range of possible scaffold mechanical properties. Other investigators have used solid free-form fabrication techniques to generate scaffolds with controlled architecture to improve the composition and biomechanical properties of engineered tissues for some orthopedic applications.10,12 To improve the biomechanical properties of engineered cardiac tissue, a knitted, elastomeric fabric (Fig. 1B) obtained from Fidia Advanced Biopolymers (HyalonectÒ, Abano Terme, Italy) was evaluated.13 Benzylated hyaluronan (Hyaff-11Ò) was extruded into microfilaments (15 mm diameter), made into a multifilament yarn (100 microfilaments per fiber), and processed into a fabric by circular weft knitting. Scaffolds freshly seeded with cardiomyocytes in a fibrin gel were comprised of densely populated cells interspersed between the multifilament yarns (Fig. 1B). In freshly seeded constructs, tensile stiffness was determined mainly by the fibrin component and was half as high as native heart tissue, whereas ultimate tensile strength, failure strain, and strain energy density were determined by the knitted fabric and were respectively 8-, 7-, and 30-fold higher than native heart tissue. At 1 week of in vitro culture, the constructs exhibited superior stress-strain behavior (within the range of native myocardium) but inferior contractility as compared to other studies in which the scaffold was made of more compliant materials such as collagen or poly(glycerolsebacate).14–16 Therefore, the mechanical design criteria for cardiac tissue-engineering scaffolds should also consider the very low values of contractile force (e.g., 250 mN17) that are currently measurable in cardiomyocyte-based constructs. To improve the integration and remodeling of engineered cartilage tissue, composites made of a cartilage construct and an osteoconductive support were implanted into large osteochondral defects in adult rabbit knee joints.18 Chondrocytes were cultured on PGA nonwoven mesh for 4 weeks and the resulting engineered cartilage was sutured to a CollagraftÒ support (Cohesion, Palo Alto, CA) and implanted. The composite implant was remodeled between the 6-week and 6-month time points (Fig. 2C) as follows. At 6 weeks, white blood cell infiltration at the base and lateral margins of the engineered cartilage construct was associated with mineralization, hypertrophic chondrocytes, and collagen type X, consistent with reformation of subchondral bone. At 6 months, the cartilage was of normal thickness and stained for collagen type II, and the subchondral bone was well mineralized. The Young’s modulus of the 6-month repair cartilage was approximately 0.80 MPa and comparable to the native articular cartilage of control rabbits, as assessed by indentation testing. Other investigators have suggested the use of a composite made of a soft cartilaginous component, to absorb compressive and shear loads, and a stiffer underlying component for osteochondral defect repair.19

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FIG. 2. Rotating bioreactor cultures. (A–D) Chondrogenesis on a nonwoven mesh: (A) 3-day, (B) 4-week, and (C) 8-week constructs, safranin-O stain, scale bars 1.0 mm (high magnification scale bars 200 mm). (D) Integration of a construct with native articular cartilage: 8-week composites, safranin-O stain, scale bar 500 mm (high magnification scale bar 50 mm). Black arrows indicate scaffold; white arrow indicates native cartilage. Color images available online at www.liebertpub.com/ten.

Complementary to the above approaches, scaffolds with micro-to-nanoscale features can be generated by using microfabrication and solid free-form fabrication technologies. For example, photopatterned hydrogels have been used to precisely control tissue architecture by spatially directed cell seeding and micropatterned membranes used to provide anisotropic guidance motifs for the orientation of cardiomyocytes and myofibrils.20,21 Recently, microfabricated scaffolds with large (100–1000 mm) rectangular pores with

defined aspect ratios (1:1, 2:1, 5:1, 10:1) were used to orient cells and collagen deposition in engineered tissue constructs.22

The role of bioreactors Bioreactors are laboratory tissue-culture devices that provide a controllable, mechanically active environment that can be used to study and potentially improve engineered

ADVANCED TOOLS FOR TISSUE ENGINEERING tissue structure, properties, and integration. For the cell seeding of 3D scaffolds, many tissue-engineering studies employ conventional methods (i.e., static seeding in petri dishes) and yield engineered constructs comprising a thin tissuelike layer at the base of the scaffold, due to gravitational settling of cells.23,24 In contrast, convective mixing (e.g., in spinner flask bioreactors) and convective flow (e.g., in perfused bioreactors) can improve initial cell seeding density and homogeneity, and thereby improve tissue architecture.24–26 For the cultivation of engineered tissues, bioreactors can improve construct size, cellularity, and molecular composition,24,27,28 and permit studies of (i) mass transport of oxygen15,26,29 and growth factors,30,31 (ii) hydrodynamic conditions,32–34 and (iii) physical stimuli, such as dynamic compression35–37 or cyclic stretch.38–41 For example, dynamic mechanical compression of engineered cartilage constructs increased ECM synthesis rates and improved biomechanical properties under some conditions, but not other conditions.35–37 Likewise, cyclic mechanical stretch of engineered cardiac constructs increased the alignment of myotubes and improved contractility and pharmacological responsiveness under some conditions,39–41 but not others.13 Properties of tissue-engineered cartilage, smooth muscle, blood vessels, and heart valves have also been improved by preconditioning grafts with pulsatile fluid flow (e.g., see Niklason et al.42) or hydrostatic pressure (e.g., see Toyoda et al.43). Bioreactors, and in particular rotating vessels, provide a favorable environment for studying construct maturation and integration in vitro.44– 47 Rotating vessels can be operated such that large 3D constructs or construct-explant composites are maintained freely suspended within the culture media and are exposed to efficient mixing in conjunction with predominately laminar flow conditions and low shear stresses (*1 dyn/cm2).48 Rotating bioreactors were used to study the development of engineered cartilage constructs.44,47,49 After 3 days of culture (Fig. 2A), cell-seeded scaffolds are highly porous and are comprised of attached cells and scattered islands of ECM. After 4 weeks (Fig. 2B) and 8 weeks (Fig. 2C) of culture, cells proliferated and deposited ECM such that the construct became cartilaginous over its entire cross-sectional area. We also carried out complementary studies in vitro, using rotating bioreactors, and in vivo, using subcutaneous implants in nude mice since the usual model of orthotopic implants in rabbit knee joints is complicated by high variability.46,47 Construct maturation and integration with articular cartilage (Fig. 2D) and trabecular bone were systematically studied, and significant individual and interactive effects of cell chondrogenic potential (primary or fifth passage chondrocytes), scaffold degradation rate (PGA or benzylated hyaluronan), and adjacent tissue architecture (articular cartilage or trabecular bone) on construct adhesive strength, compressive modulus, and biochemical composition were demonstrated. Freed and

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Guilak concluded that as the field of tissue engineering progresses, it is becoming more apparent that the development of functional tissue replacements may require additional exogenous influences to achieve many of the important requirements for long-term success. Smart scaffolds and bioreactors can be used to (i) provide more controllable tissue architecture such that constructs will exhibit physiologic biomechanical properties, and (ii) enhance construct development and integration such that constructs will be able to restore native tissue functionality. In addition to a defined architecture and biomechanical properties, scaffolds can be used to provide long-term physical and/or biochemical signals for the maintenance of cell phenotype, metabolism, and differentiation. Bioreactors can be used before implantation to accelerate tissue growth and differentiation, or to test the physiologic response of engineered constructs to the physical environment to which they may be exposed following implantation. Other rapidly evolving technologies also may have a significant impact on tissue engineering, and it is important to consider the use of smart scaffolds and bioreactors as additional tools in light of the role of novel growth factors, biomaterials, gene therapies, and other changing technologies.

MECHANICAL MODULATION IN FUNCTIONAL TISSUE ENGINEERING AND MECHANOBIOLOGY Edward Guo from Columbia University emphasized the importance of mechanical modulation in functional tissueengineering research and applications, and discussed interactions among functional tissue engineering, mechanobiology, and developmental biology. He argued that the desired physiological function of musculoskeletal, vascular, and cardiac tissues cannot be achieved without an adequate consideration of mechanical modulation, and that a combination of mechanical and biochemical factors, rather than any of the factors alone, will likely be the winning recipe. Tissue engineering calls for research in the field of mechanobiology, especially in coculture systems. In turn, tissue engineering generates opportunities for studies of mechanobiology. The last but not the least, it will be important for tissue engineers to learn about mechanical forces during tissue and organ genesis. Mechanical forces play important roles in organogenesis during embryonic development.50 For example, muscle forces are necessary for the development of skeleton from their cartilaginous analogs.51 Thus, it is not surprising that a physiologic level of mechanical force is needed for engineering tissues where mechanical functions are critical, such as musculoskeletal or cardiovascular system. In the past decades, especially since the clear articulation of the concept of ‘‘Functional Tissue Engineering,’’6 the field experienced an explosive growth and attracted a lot of academic and industrial attention. There are many successful

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stories of the development of engineered tissues when mechanical forces have been appropriately considered. In the musculoskeletal system, physiologic levels of mechanical loading increased the contents of proteoglycans and type II collagen, and most importantly improved mechanical properties of the engineered cartilage.36,37,47 In bone tissue engineering, mechanical loading was necessary to achieve an appropriate elastic modulus and microstructure organization.52–54 In vascular tissue engineering, pulsatile pressure was necessary to engineer patent vascular graft with the rupture strengths greater than 2000 mmHg, suture retention strengths of up to 90 g, and collagen contents of up to 50% of normal.42 In cardiac tissue engineering, electrical stimulation stimulated the functional assembly of cardiac myocytes cultured on collagen scaffolds into functional engineered myocardium.14,15 Mechanical forces increased the elasticity of tissue-engineering muscle constructs by two- to threefold.55 It becomes quite apparent from the progress in functional tissue engineering in the past decade that mechanical stimulation should be an integral part of any attempt to engineer a functional tissue where mechanical function is physiologically important. One should consider mechanical stimulation of engineered tissues equally important as is the supplementation of appropriate growth factors. Mechanical stimulation should become a necessary prerequisite of any future functional tissue-engineering studies. The incorporation of mechanical stimulation in functional tissue engineering naturally requires detailed knowledge of how cells and tissues respond to a particular mechanical stimulus, a subject of the exciting field of mechanobiology. Tremendous progress has been made in characterizing the mechanotransduction at the tissue, cell, or gene expression levels. However, it is far from clear how the mechanical loading is transduced to the individual cells (or which cells are the sensors), and what are the exact molecular and cellular responses. Therefore, functional tissue engineering presents the field of mechanobiology an urgent need and opportunity for comprehensive studies of a variety of cell and tissue types. Synergistic interaction between the fields of tissue engineering and mechanobiology should be pursued in the future. For example, sessions on mechanobiology and its application in tissue engineering should be planned in future tissue-engineering conferences. The advancement of tissue engineering also brings many challenges to mechanobiology. The successful development of engineering tissue constructs requires interaction and integration of multiple cell types in a coculture system. Indeed, a few mechanobiology studies have demonstrated that interaction and communication between different cell types are important for mechanotransduction. In a recent coculture study of live trabecular bone explants seeded with osteoblasts, osteocytes embedded in the mineralized bone tissue became more effective mechanical sensors.56 Many new scaffolding materials have been developed, and the interaction of cells with these new materials under me-

FREED ET AL. chanical loading will be important for the success of tissue engineering. Most in vitro mechanobiology studies involve cells adhered to simple surfaces, such as plastic or metal. It has been shown that the extracellular substrate can influence cell shape, cytoskeleton, elastic modulus,57 and mechanotransduction of cells.58 Research of tissue engineering with mechanical loading also provides many opportunities for studies of mechanobiology, both in vitro and in vivo. For example, cell-seeded scaffolds under mechanical loading can serve as outstanding models for mechanistic studies of mechanobiology. Thus, functional tissue engineering and mechanobiology should be considered as an integral part of modern biology and biotechnology. Edward Guo concluded that tissue engineers should learn from developmental biology what is the mechanical influence on morphogenesis of tissues or organs. Mechanobiology of embryonic development should be a part of the new initiatives of developmental biology and tissue engineering. As biomedical engineers, it is our mission to bring mechanical forces to the front seat of modern biology and functional tissue engineering. Mechanical forces are as important as growth factors or cytokines.

NONDESTRUCTIVE MONITORING OF NATIVE AND TISSUE-ENGINEERED CARTILAGE Martha L. Gray from Massachusetts Institute of Technology argued that a fundamental challenge in realizing the potential offered by tissue engineering is to reduce the time it takes for innovations and discoveries to reach the patient. The barriers to realizing the potential are not ‘‘simply’’ developing better, more effective tissue-engineered products and strategies. There are also significant barriers introduced because there are only few strategies and technologies available for evaluating the function and stability of tissueengineered products. This barrier has received much recent attention, especially by the FDA in its Critical Path Initiative.59 The FDA has recommended ‘‘an aggressive collaborative effort to create a new generation of performance standards and predictive tools.’’59 Biomedical imaging is cited as a specific opportunity, noting that ‘‘new imaging techniques will ultimately contribute important biomarkers and surrogate endpoints, but how soon these new tools will be available for use will depend on the effort invested in developing them specifically for this purpose.’’59 Imaging is a particularly attractive, enabling technology for biomarker development because of its capacity to provide nondestructive monitoring with spatial and temporal resolution. While biomarkers are widely acknowledged as being enabling, it should also be recognized that the continuum from discovery to utility of a biomarker could be quite involved and challenging. Critical elements of that continuum include the validation and qualification of a biomarker. Here we briefly describe these elements, using as a case example, a putative

ADVANCED TOOLS FOR TISSUE ENGINEERING imaging biomarker for native and tissue-engineered articular cartilage. Specifically, we will explore the validation and (ongoing) qualification of molecular imaging of glycosaminoglycan (GAG) by magnetic resonance (MR) imaging as a biomarker for the quality and likelihood of sustained quality of cartilage. Face validity assesses whether a measurement ‘‘on its face,’’ that is, qualitatively, seems like it is a reasonable measurement that would aid in analysis and decision making. Practically speaking, face validity is crucial to both the validation and qualification of a biomarker because it greatly aids dissemination if potential users of the measurement are likely to easily accept that the measurement might be useful. To illustrate with the specific case at hand, we ask whether, based on what is known, a measurement of GAG in cartilage would be generally viewed as relevant and useful. The short answer is yes, as evidenced by the fact that other measurements of GAG are widely used to assess cartilage quality. A longer answer is rooted in a deeper understanding of cartilage physiology: a key function of cartilage is to support mechanical loads. The load-bearing properties of cartilage are determined by the tissue composition and architecture. In turn, tissue composition and architecture are influenced by mechanical load. Of the many components of articular cartilage, GAG is responsible for much of the loadbearing capacity. Accordingly, histological evaluation of cartilage centers on the amount of GAG staining; assessment of tissue-engineered cartilage has, to date, consistently included a measure of GAG. These and other data support the case that GAG measures are viewed as relevant and useful, thus molecular imaging of GAG as a biomarker has face validity. Technical validity examines the extent to which the proposed measurement actually works. For the specific example of MR imaging of GAG, a variety of in vitro and in vivo validation studies have demonstrated that the MR approach is sensitive to and specific for GAG60–66 [also references within Refs. 60–66]. (Patho)physiological validity asks whether the measurement is sensitive to changes or differences that are physiologically or clinically meaningful, and whether differences in the measurement reflect important (patho)physiological differences. A demonstration of (patho)physiological validity is one of the early steps in qualifying a biomarker for widespread acceptance and use. The extent to which pathophysiological validity is demonstrated dictates the uses of the biomarker. (It need not be the case that a single biomarker be suitable for every imaginable circumstance.) For tissue-engineered cartilage, molecular imaging of GAG tracks very well the development of tissue as assessed histologically and biochemically, suggesting its utility in monitoring development over time (Fig. 3). Molecular imaging of GAG has also been used to monitor enzymatic- and cytokine-induced GAG loss, and subsequent recovery.67,68 In these examples, the ability of imaging to provide both spatial and temporal information nondestruc-

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tively offers the opportunity for studies that are not feasible (or even doable) with destructive measurements. Turning to in vivo applications, molecular imaging of GAG has been used to examine a cross section of subjects who had received autologous chondrocyte transplantation (ACT), one of the tissue-engineering strategies for cartilage repair. In that study, it was straightforward to observe that, as expected, soon after ACT, there was little GAG in the region of the transplant. In this cohort, subjects who had had their transplant for more than 18 months showed considerable GAG within the implant (in many cases, GAG that was comparable to the surrounding native articular cartilage).69 Randomized longitudinal study is required to use these kinds of data to evaluate the efficacy of ACT. What these data do suggest, however, is (a) that the measurement is sensitive to the range of GAG expected to be important clinically and (b) these kinds of measurement offer the ability to assess efficacy in a way that is not possible using the ‘‘gold standard’’ of randomized biopsies. Martha L. Gray concluded that the delivery of engineered products to clinical use will require methods for assessing their development and efficacy. Nondestructive imaging technologies have the potential to enable studies (both in vitro and in vivo) that are not possible or practical with other technologies. Furthermore, they offer the potential to accelerate the development, evaluation, and delivery of tissueengineered products to patients. To realize that potential, research attention is needed to develop specific imaging biomarkers. Such biomarkers must be associated with the problem (disease) being addressed, must change as the problem (disease) changes, and must be impacted by the innovation (i.e., the tissue-engineered product). Some biomarkers may prove to be surrogate markers, that is, to be predictive of outcome. While a surrogate marker is undeniably valuable, a biomarker does not need to reach that standard to have considerable positive impact. Successful imaging (and other) biomarkers should be ones that are easier, faster, cheaper, more robust, and more reliable than existing markers; should be developed for a specific purpose; and should have demonstrated validity. With the codevelopment of tissue-engineered products and associated biomarkers, we have the potential to accelerate the development of these products to advance human health.

ENGINEERING STRUCTURAL TISSUES IN VITRO Robert Tranquillo from the University of Minnesota focused on the in vitro fabrication of structural tissues, whose primary function is mechanical and which do not have a major vascular network component—blood vessels, cardiovascular valves, ligaments, tendons, cartilage, and skin. The approach he described eliminates the need to address the major hurdle of high oxygen and nutrient demand intrinsic to highly metabolic tissues, and motivates the need to

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FIG. 3. Example using magnetic resonance imaging (MRI) of glycosaminoglycan (GAG) as a means of monitoring the development of tissue-engineered constructs. These constructs were composed of chondrocytes isolated from immature bovine articular cartilage within a polyglycolic acid (PGA) scaffold, and cultured for 6 weeks in static conditions. (A) Accumulation of GAG within the tissueengineered constructs and the loss of GAG from explants were observed using nondestructive GAG monitoring. (B) GAG as measured by MRI corresponds well with GAG measured biochemically. (C) GAG images by MRI correspond well with GAG staining by safranin-O histology (S.N.O. Williams, B. Obradovic, D. Burstein, M.L. Gray, R. Langer, L.E. Freed, and G. Vunjak-Novakovic, unpublished). Color images available online at www.liebertpub.com/ten.

emulate the structure-function property characteristic of structural tissues. A logical consequence is that a tissueengineered construct should serve as a functional remodeling template, so that while providing function during the remodeling, the artificial tissue also provides a template for the alignment of the tissue that grows during remodeling. One paradigm for a functional remodeling template is based on controlled cell remodeling of biopolymers (see Fig. 4), wherein a suspension of cells in a biopolymerforming solution is injected into a mold of appropriate geometry. The first phase of remodeling following cell entrapment in the biopolymer gel, which is a highly hydrated network of self-assembled protein fibrils, is primarily a structural remodeling (termed ‘‘compaction induced alignment’’), wherein the cells compact the fibril network by cell traction forces, causing exudation of the interstitial liquid. Mechanical constraints imposed on the compaction result in fibril alignment, which in turn results in cell alignment via a contact guidance response. The second phase is primarily a compositional remodeling, wherein the aligned biopolymer fibrils are enzymatically degraded and replaced by cellproduced ECM fibrils/fibers; this occurs over weeks as opposed to days for the first phase. While the first phase occurs robustly with either type I collagen or fibrin gels, the second

phase only occurs to a degree of significance with fibrin gel, as tissue cells in collagen gel exhibit little proliferation and ECM production. As suggested in Figure 4 and discussed below, the cell-produced ECM fibrils/fibers can be aligned with the degrading fibrin fibrils. Some cell types require the inhibition of fibrinolysis to ensure that the time scale of fibrin degradation matches that of ECM production. A more detailed summary of this line of research is available,70 as well as a review of studies of collagen and fibrin gel properties and cell behavior in these gels.71 This paradigm can be realized in a construct intended to mimic the medial layer of an artery, wherein constrained compaction of fibrin gel tubes by entrapped smooth muscle cells around a nonadhesive mandrel over 3–5 weeks of incubation yields a circumferentially aligned tube comprised of cell-produced ECM and residual fibrin fibrils (complete medium is supplemented with transforming growth factor-b1 and insulin to promote collagen and elastic fiber production as well as with the fibrinolytic inhibitor aminocaproic acid).70 This paradigm can also be realized in more complicated geometries, such as the aortic valve, by designing an appropriate mold that results in circumferential alignment in the root and commissural alignment in the leaflets following compaction induced alignment.72

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nonadhesive mandrel

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compacted construct

gelled collagen/fibrin with entrapped SMCs

remodeled construct

nonadhesive mandrel leads to circumferential alignment of SMCs and collagen/fibrin fibrils

alignment of SMCs and collagen/fibrin fibrils leads to alignment of SMC-produced ECM

FIG. 4. Controlled cell remodeling of biopolymers illustrated for a tubular construct termed ‘‘media equivalent.’’ A suspension of smooth muscle cells (SMCs) in a solution of monomeric type I collagen or fibrinogen and thrombin is injected into the tubular cavity of a mold possessing a nonadhesive mandrel (e.g., TeflonÒ). Following rapid fibrillogenesis under physiological conditions, the cells are entrapped in a network of native protein fibrils that are randomly oriented. Following gel compaction that is driven by cell traction forces on the fibrils, the fibrils and consequently the cells (via contact guidance) become circumferentially aligned in a structural remodeling termed ‘‘compaction induced alignment.’’ Particularly in the case of fibrin, compositional remodeling then occurs over a longer time course, wherein the cells degrade the fibrils and produce ECM that exhibits the same circumferential alignment. Reprinted with permission from Isenberg et al.70 Color images available online at www.liebertpub.com/ten.

It is evident that this paradigm is dependent on a large number of fabrication variables that affect remodeling, including biopolymer type and structural characteristics (e.g., density, pore size, fiber diameter, entanglement/ branching frequency, network and fibril stiffness, intra- and inter-fiber cross-linking), cell type and source (including species, age, passage number, and culture conditions), medium composition (serum and/or supplementing factors, as they affect cell proliferation, ECM production, and ECM degradation), nutrient availability (diffusional limitations determined by construct thickness and composition and by cell loading), autocrine factors (related to construct thickness and cell density), and mechanical stress state of the construct (static vs. cyclic loading, mechanical stretching vs. hydrodynamic shear). Many of these variables are of importance to other approaches based on fabricating cell-polymer constructs in vitro even though compaction-induced alignment does not occur because the polymers are too rigid to become compacted or are not fibrillar and so cannot be aligned. A central question (relevant to any approach) is does the exquisite design of the target native tissue need to be reproduced in order to achieve adequate function? If true, the

objective is equivalent to engineering native tissue, which is not realistic. The success of tissue engineering indeed rests on the assumption that adequate function can be achieved via compositions that are different from and structures that are much simpler than the native tissues, and which are realizable via manipulation of cells and biomaterials in vitro. The development and analysis of tissue mechanical models to identify solutions consistent with adequate function would aid in identification of realizable objectives. For example, bioprosthetic heart valves with tissues comprised of bovine pericardium yield adequate function despite lacking the trilayered structure of native leaflets. The eventual failure of bioprosthetic valves may be due to their inability to remodel significantly because of the chemical fixation used to minimize antigenicity; in a tissue-engineered construct, postimplantation remodeling, that is, growth (if appropriate) and long-term self-maintenance, would presumably occur and overcome this failure mechanism. Another central question is how does the cell integrate mechanical forces/deformations and chemical stimulation (via soluble or substratum bound ligands) into the relevant responses, such as traction, migration, proliferation, ECM degradation, ECM production, and gene expression/

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differentiation? The cell can be considered an extremely complicated system, and the answer to this question is tantamount to definition of the open-loop system. In a fascinating and well-accepted but poorly understood feedback process, cell phenotype, which dictates the nature of a cell’s future ECM remodeling ability, is determined by the nature of its past ECM remodeling as it determines the current local ECM composition and stress-strain state. In addition to these ECMassociated feedback signals, an evolving cell concentration will generate feedback through autocrine and paracrine factors. Defining the associated closed-loop system, a collection of these feedback processes, involves knowledge of the cell’s set point, which is the steady-state (homeostatic) condition that the cell attempts to attain and maintain. This will require exhaustive and carefully controlled studies that still exceed current capabilities, since precise measurements of a multitude of molecules and localized forces/deformations within and around cells are involved. These studies will also necessarily be high-throughput and require the development of small constructs in microfluidics/microelectromechanical system devices (MEMS) that are valid models for large constructs needed for actual applications. There are challenges and opportunities more peculiar to the paradigm of construct fabrication illustrated in Figure 4. Since cell traction drives fibrillar network compaction and ultimately alignment, one key question in the compaction and alignment phase is what determines the degree and directionality of cell traction forces? It has been concluded that cell traction varies during network compaction, although the controlling variables were not defined, and mechanical models must assume this in order to predict the observed compaction attaining a plateau. Such models assume that the principal direction of cell traction is along the cell polarity axis, as suggested by polarized light micrographs, but this has not been comprehensively studied as for the more tractable case of cells on a 2D elastic substratum. A related question is how does cell traction result in alignment of the fibrils in the network when compaction is mechanically constrained? While the strain-based model we proposed yields accurate predictions of macroscopic alignment, microstructural models that account for fibril properties (e.g., entanglements, branching, interfibrillar cross-links) will allow the more complex alignment patterns that exist on the order of the cell dimension to be predicted. Presumably the cells cannot detect anisotropic fibril orientation directly, but are detecting the associated anisotropy of adhesion sites on the fibrils, anisotropic viscoelasticity of the fibrils, and/or anisotropy of network porosity via interactions of extending and retracting pseudopods with the local fibrils.73 Experiments to manipulate just one of these anisotropies without affecting the others are difficult to conceive and conduct, so the predominant mechanism that could be exploited in the design of the fibril network remains elusive. Another major question is how do the aligned fibrin fibrils cause the cell-produced ECM fibrils to exhibit the

FREED ET AL. same alignment? One possibility is a direct influence by providing a physical template for assembly of the nascent ECM fibrils. Another possibility is an indirect influence via the cell alignment associated with the contact guidance response (i.e., new ECM fibrils grow in direct association with the cell surface), and since a cell is typically elongated and aligned with fibrin fibrils and moving bi-directionally with respect to them, the ECM fibrils thereby become aligned with the fibrin fibrils. In this mechanism, cell migration may therefore play a key role. It must be recognized that the mechanisms for collagen fibril and elastic fiber production, the two key ECM proteins in determining mechanical properties of the growing tissue, and therefore their organization and alignment may fundamentally differ. The critical role of elastic fibers in the elasticity of most tissues underscores the need to improve the generally inadequate elastogenesis that occurs in tissue-engineered constructs. Robert Tranquillo concluded that elucidating and exploiting these mechanisms is clearly important for construct fabrication using multiscale tissue mechanical models that relate detailed microstructure and composition to macroscopic mechanical properties. Ultimately, constructs could be designed by solving the related inverse problem, first declaring the target mechanical properties, and then using such tissue mechanical models to evolve backward in time in concert with the closed-loop system model indicated above, which together would define the optimal bioreactor operation, to the appropriate starting formulation of the construct given the choices of cell type and polymer used. The ability to noninvasively monitor all of the relevant tissue construct variables (mechanical properties, cellularity, gene expression, ECM composition, concentrations of soluble ligands, nutrients, etc.) with high spatial resolution would be critical to this approach. While a control-based approach to bioreactor operation for optimal tissue growth represents a long-term vision, it would provide a rationale basis for the engineering of structural tissues.

ENGINEERED TISSUES AS TOOLS FOR STUDYING MECHANOBIOLOGY Jeffrey W. Holmes from Columbia University proposed that the ability to design and construct engineered tissues tailored for use as model systems to investigate specific biologic questions is an important but under-appreciated application of tissue engineering. He provided a conceptual model for the creative use of engineered tissues to answer biologic questions by considering the fibroblast-populated collagen gel, an extremely simple engineered tissue described 25 years ago. Cell-populated collagen gels have been employed to ask basic questions about cell-matrix interactions, cell and matrix mechanics, and mechanobiology, taking advantage of the ability to control composition, geometry, and microstructure in this system.

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Fibroblast-populated collagen gel system In 1979, Bell et al.74 described the fibroblast-populated collagen gel system in a publication titled ‘‘Production of a tissue-like structure by contraction of collagen lattices by human fibroblasts of different proliferative potential in vitro.’’ Their original description of this rudimentary tissue anticipates so accurately both its subsequent application and the present argument that the abstract is reproduced here in its entirety (emphasis added): ‘‘Fibroblasts can condense a hydrated collagen lattice to a tissue-like structure 1/28th the area of the starting gel in 24 hr. The rate of the process can be regulated by varying the protein content of the lattice, the cell number, or the concentration of an inhibitor such as colcemid. Fibroblasts of high population doubling level propagated in vitro, which have left the cell cycle, can carry out the contraction at least as efficiently as cycling cells. The potential uses of the system as an immunologically tolerated ‘‘tissue’’ for wound healing and as a model for studying fibroblast function are discussed.’’

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cast the gels in a convenient rectangular geometry, suspended between embedded grips, and measure force generation along one axis of the rectangle with a force transducer.79–81 Another strategy was to track deformation of the collagen by the embedded cells and use measured collagen material properties to estimate the associated forces.82 Working over much longer time scales, Barocas and Tranquillo took an analogous approach, using a biphasic model of collagen gel compaction to estimate the associated cell forces.78,83 More recently, Elson and coworkers have cyclically loaded fibroblast-populated collagen gels and used serum activation and cytoskeletal disruption to separate the active cell component from the passive matrix properties.84 In these studies, collagen gels were used to control both geometry (employing rectangles with embedded grips, spheres, hemispheres, cylinders, and rings) and composition (varying collagen density and fibroblast-to-collagen ratios) of the engineered tissue analogs.

Matrix mechanics Cell-matrix interactions One of the earliest uses of fibroblast-populated collagen gels was to understand the elements required for fibroblasts to attach to collagen and compact the gels. These studies (see Grinnell75 for a brief review) revealed that attachment and compaction require attachment through a2b1 integrins, an intact cytoskeleton, myosin light chain kinase activity, and serum or specific growth factors. This system was promoted as an in vitro model of wound healing, and the impact of different cytokines and growth factors present in healing wounds was explored, as was the question of whether fibroblasts from wounds at different stages of healing demonstrate different activity in compacting gels.76 While these early studies did not take full advantage of the ability to control gel composition, geometry, and structure, they did rely on one important tissue-like feature of these simple engineered tissues—the fact that cells are embedded in a biologically appropriate 3D matrix rather than plated on a 2D surface. The studies on the interaction between cellular and matrix cues in determining cell alignment have taken advantage of the ability to control geometry in fibroblast-populated gels. For example, comparing gels of different shapes allowed Costa et al. to test and reject the hypothesis that fibroblasts align along the direction of greatest local tension.77 Comparing experimental to model results over a range of different shapes allowed Barocas and Tranquillo to establish that fibroblast compaction of the matrix by the cells leads to matrix alignment, which then provides orientation cues to the cells.78 In both cases, hypothesis testing was greatly facilitated by the ability to cast gels with desired geometry.

Cell force generation and cell mechanics Collagen gels have also been adapted to measure force generation by the embedded fibroblasts. One strategy was to

Interest in using collagen gels as a foundation for tissue engineering has led to a number of studies of the mechanical properties of these gels. Two of these studies have taken particular advantage of the ability to vary gel microstructure in a series of test specimens to develop a better understanding of the structural basis for gel mechanical properties. By varying collagen concentration and pH during gel polymerization, Roeder et al. were able to vary the collagen fibril diameter, length, and density and determine the impact of these microstructural features on mechanical properties as reflected in uniaxial tensile tests.85 While the authors’ intent was to provide information useful in the design of scaffolds for engineered tissues, this study also sheds light on the origin of mechanical properties in collagenous tissues with different microstructure. More recently, by manipulating boundary conditions on fibroblast-populated collagen gels, the contractile forces exerted by the cells on the matrix were harnessed to produce structurally and mechanically anisotropic engineered tissues.86 In these studies, Holmes and associates utilized the ability to cast fibroblast-populated gels in virtually any geometry, first by creating specimens with aspect ratios appropriate for planar biaxial mechanical testing and second by coupling the gels to a simple, low-cost static loading system for manipulating boundary conditions during development in the incubator.87 The primary interest was to employ engineered tissues as a tool for developing mathematical descriptions of tissue material properties that are based on measured tissue structure (structural constitutive models), varying structural features such as fiber distribution and levels of cross-linking that are thought to be critical determinants of mechanical anisotropy in tendon, skin, and myocardial scar tissue, and establishing how to best model each important feature. However, the process by which asymmetric boundary conditions direct the development of

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structural anisotropy in these gels may also be a reasonable analog for important developmental processes.

Fibroblast mechanobiology Fibroblast-populated gels have particular appeal for mechanobiology studies because the baseline expression of many genes of interest is much closer to in vivo levels for cells embedded in a 3D matrix than for cells in planar culture.88 Lambert et al. compared collagen and metalloproteinase expression in free-floating and restrained gels containing dermal fibroblasts to assess the potential impact of mechanical stress on healing skin wounds, and found that restrained gels showed increased collagen and decreased collagenase expression, suggesting a role for mechanical stress in regulating scar formation.88 A collagen gel-based system for investigating dose-response relationships between the load and fibroblast biology87 was recently applied to investigate the impact of mechanical load on aspects of dermal fibroblast biology that are important to wound healing.89 These simple studies only hint at what could be accomplished in the area of mechanobiology given modern methods for design of scaffolds, cell-scaffold interfaces, and application of various types of mechanical stimuli outlined elsewhere in this special issue.

Engineered tissues as tools for studying mechanics and mechanobiology Generalizing from the examples presented above, there is a strong case to be made for using engineered tissues as model systems for studying cell mechanics, matrix mechanics, cellmatrix interactions, and mechanobiology. Engineered tissues provide a balance between the benefits of cell culture (control of the hormonal, chemical, and mechanical environment, ease of visualization) and the benefits of traditional organ culture (more physiologic 3D cell-cell and cell-matrix interactions) in addressing biologic questions. In addition, engineered tissues offer unprecedented control over specimen composition, organization, and geometry, allowing the researcher to design the specimens that best facilitate experimentation, modeling, and interpretation for a particular question of interest. Jeffrey W. Holmes suggested that as the field of tissue engineering moves forward, increased use of engineered tissues as model systems for research offers both strategic and scientific advantages. Strategically, the main advantage is that current techniques can provide engineered tissues appropriate to research on a wide range of biologic questions, allowing the field to promote its current capabilities while continuing to pursue technically more ambitious goals such as total replacement of various tissues and organs. Scientifically, studies employing engineered tissues as model systems would promote collaboration between engineers, biologists, and biomedical scientists on hypothesis-driven basic research in areas such as mechanotransduction and mechanobiology.

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BIOMIMETIC REQUIREMENTS FOR CARDIAC TISSUE ENGINEERING Milica Radisic from the University of Toronto discussed the first and second generation of engineered cardiac tissues. She explained that the motivation for cardiac tissue engineering comes from the great clinical need, with nearly 8 million people in the United States who have suffered from myocardial infarction (MI), and 800,000 new cases occurring each year.90 MI results in the substantial death of cardiomyocytes in the infarct zone followed by a vigorous inflammatory response and removal of dead cells by marrow-derived macrophages. Over the subsequent weeks to months, fibroblasts and endothelial cells proliferate forming granulation tissue and ultimately a dense collagenous scar. Formation of the scar tissue severely reduces contractile function of the myocardium and leads to a pathological remodeling process that includes ventricle wall thinning, dilatation, and ultimately heart failure, which affects more than 500,000 patients in the United States each year.91 Conventional therapies are limited by the inability of myocardium to regenerate after injury92 and the shortage of organs available for transplantation. Patients with large transmural akinetic scars often benefit from the Dor procedure (endoventricular circular patch plasty).93,94 In this procedure the scar tissue is excised and the ventricle is closed using a circular Dacron (polyethylene terephtalate) patch lined with endocardium. Cell injection95,96 and tissue engineering were proposed as alternative treatment options. Both approaches appear viable, and the appropriate regeneration strategy will likely depend on time postinfarction and the size of the affected area. Application of cells and growth factors within hours and days after MI has a potential of directing the wound repair process so that the minimum amount of scar tissue is formed, the contractile function is maintained in the border zone, and pathological remodeling is attenuated. Tissueengineering strategies will work in the acute phase as well, but may be more necessary after scar has formed for replacement of noncontractile areas. In addition, tissue engineering may yield functional replacement tissue for repair of congenital malformations.

Cardiac tissue engineering: the first generation Tissue engineering generally involves the presence of reparative cells, the use of scaffolds (designed to provide a structural and logistic template for tissue development and biodegradation at a controlled rate), and bioreactors (designed to control cellular microenvironment, facilitate mass transport to the cells, and provide the necessary biochemical and physical regulatory signals). Three-dimensional cardiac tissue constructs that express structural and physiological features characteristic of native cardiac muscle have been engineered using fetal or neonatal rat cardiac myocytes on collagen fibers,97 fibrous PGA scaffolds,34,98–101 and porous

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collagen scaffolds.26,102,103 In these early studies, cells were seeded onto scaffolds and cultivated in dishes,34,99,102,104 spinner flasks,34,98,99 or in rotating vessels.34,97,99 In all of these systems, oxygen dissolved in medium was transported to the cells by molecular diffusion. Whereas human cardiac muscle is *2 cm thick, diffusion alone was able to support only four to seven cell layers (up 100 mm thick) in these constructs.34,98,99,105 Oxygen gradients measured in statically grown cardiac constructs correlated strongly with the gradients of cell viability and density.106 Another limitation of conventional culture systems was the lack of appropriate physical stimuli that yield poorly differentiated nonaligned cardiomyocytes. The exception was a mechanical stimulation system pioneered by Eschenhagen and colleagues.39,41,105–107 Application of cyclic unidirectional stretch to the constructs based on neonatal rat cardiomyocytes and collagen gel yielded contractile engineered heart tissue consisting of differentiated cardiomyocytes aligned in parallel. However, diffusional oxygen supply limited the size of compact cardiac strands to approximately 100 mm.41

Cardiac tissue engineering: the next generation In an attempt to address some of the identified limitations, a ‘‘biomimetic’’ approach was developed to design culture systems such that they mimic some aspects of the native myocardial environment (Table 1). Since cardiac myocytes have limited ability to proliferate, cells need to be seeded at high densities while maintaining their viability.26 This can Table 1.

be achieved via a two-step approach, in which a suspension of cells in Matrigel was inoculated onto collagen sponges followed by the alternating flow direction medium perfusion. Cultivation of cardiac constructs in the presence of convective-diffusive oxygen transport in perfusion bioreactors maintained aerobic cell metabolism, viability, and uniform distribution of cells expressing cardiac markers.15 To investigate the effect of multiple cell types on the properties of engineered cardiac tissue, cardiac fibroblasts and cardiac myocytes were cultivated synchronously, separately, or serially (pretreatment of scaffolds with fibroblasts followed by the addition of myocytes) (Radisic, unpublished). Pretreatment remarkably improved contractile response as well as biochemical and morphological properties of engineered cardiac tissue, most likely due to the deposition of ECM (e.g., collagen) and soluble factors that enhanced functional tissue assembly after addition of myocytes. In order to mimic capillary structure, cardiac fibroblasts and myocytes were cocultured on a scaffold with a parallel channel array that was perfused with culture medium supplemented with synthetic oxygen carrier [perfluorocarbon (PFC) emulsion]. The presence of PFC emulsion resulted in significantly higher cell density and improved contractile properties compared to the constructs cultivated in the culture medium alone, presumably by increasing total oxygen content.16 Mathematical model of oxygen transport indicated that the release of oxygen from PFC particles was not rate limiting and that the improvements in the oxygen transport were due to the increase in effective diffusivity of the culture medium supplemented with PFC emulsion and increase in

BIOMIMETIC APPROACH TO CARDIAC TISSUE ENGINEERING: MAIN FACTORS OF THE IN VIVO MYOCARDIAL ENVIRONMENT AND THEIR IN VITRO COUNTER PARTS In vivo 8

In vitro 3

Cells

High density (*10 cells/cm ) Multiple cell types (cardiac myocytes, fibroblasts, endothelial cells)

High density (*0.3
108 cells/cm3) Multiple cell types (myocytes, fibroblasts, endothelial cells)

Geometry

Capillary network (*7 mm diameter, *20 mm spacing)

Parallel channel array (100–300 mm diameter, 100–300 mm spacing)

Mass transport

Convection: Blood flow *500 mm/s Diffusion in tissue space

Convection: Medium flow at velocities *50–500 mm/s Diffusion in construct space

Oxygen carrier

Hemoglobin (arterial blood) O2 dissolved in plasma 130 mM O2 as oxyhemoglobin 8,500 mM O2 total 8630 mM

Physical signals

Excitation (by electrical signals from pacer cells); synchronous contractions and electro-mechanical cell coupling (functional gap junctions; contractile apparatus).

PFC emulsion (at 160 mmHg and 378C) O2 dissolved in aqueous phase 220 mM O2 in PFC particles 230 mM O2 total 450 mM Electrical stimulation (by pacing) of cultured constructs; synchronous contractions and electro-mechanical cell coupling

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the convective term proportional to the fraction of PFC particles.106 To improve cell morphology and functional tissue assembly, cardiac constructs were cultivated with electrical stimulation of contraction in a physiologically relevant regime.14 Orderly excitation-contraction coupling enabled formation of tissue with elongated cells aligned in parallel and with organized ultrastructure remarkably similar to the one present in the native heart. The parameters of interest have been tested in individual systems enabling medium perfusion or electrical stimulation. In order to achieve a truly biomimetic system, it is necessary to integrate these individual factors into a single advanced bioreactor capable of simultaneous application of medium flow (for oxygen supply as well as mechanical stimulation via pulsatile regime) and electrical stimulation. The use of such an advanced bioreactor and channeled scaffolds16 would provide an environment for coculture of major cardiac cell populations to engineer vascularized myocardium. These multiparametric bioreactors can be made as macro-systems [for engineering of clinically sized (*2 cm), differentiated, and vascularized myocardium] or micro-systems (for highthroughput drug testing).

Cell source The major limitation in engineering cardiac tissues (either for cardiac patch or for drug testing) is the lack of an appropriate human cell source. Adult cardiac myocytes are terminally differentiated and have no ability to proliferate;92 thus, they cannot be utilized as a source of autologous cells for tissue engineering. Cardiac myocytes can be obtained at potentially unlimited quantities from embryonic stem cells.108,109 However, nuclear transfer is required to make them autologous and the presence of undifferentiated cells may lead to teratomas upon implantation.95 Recent studies report the cultivation110,111 and implantation of cardiac grafts based on mouse embryonic stem cells, and the use of human mesenchymal stem cells (hMSCs) from bone marrow112 or liposuction aspirates.113 Transplantation of cell sheets based on adipose tissuederived MSCs improved contractile function in a rat model of coronary artery ligation.114 A small fraction of MSCs differentiating into cardiomyocytes and newly formed blood vessels might contribute to the improvement in function. Recent emerging work suggests that the heart may contain resident progenitor cells. This is an exciting possibility, as resident progenitor cells may be an ideal source of autologous cardiomyocytes. However, it appears that there is more than one heart cell sub-population that fits the description of a cardiac progenitor. C-kitþ cells isolated from adult rats’ hearts and expanded under limited dilution gave rise to cardiomyocytes, smooth muscle cells, and endothelial cells when injected into ischemic myocardium.115 Oh et al.116 reported Sca-1 as a marker of resident cardiac progenitors, and expression of cardiac markers upon treatment with 5-azacytidine. LIM homeodomain islet 1 (isl1þ)

was also identified as a marker of resident cardiac progenitor cells.117 The isl1 þ cells from the hearts of mice were propagated in culture and they differentiated into functional cardiac myocytes when in contact with terminally differentiated cardiomyocytes. It remains to be determined if the progenitors can be isolated from adult human biopsies and if sufficient numbers of cardiomyocytes (>108 cells/patient/ patch) can be generated in vitro.

In vivo studies While significant progress has been made in constructing in vitro cultivation systems14,41 and biomaterial scaffolds,13,111,118 few studies have focused on implantation of cell-based cardiac patches onto viable or injured myocardium. In a pioneering study, Li et al.102 implanted a construct based on neonatal rat cardiomyocytes and collagen sponges onto the surface of the cryoinjured myocardium of Lewis rats reporting vascularization and cell survival after 5 weeks in vivo. Attenuation of pathological remodeling (i.e., prevention of ventricle dilatation and maintenance of contractile function) was observed in a study by Leor et al.,104 where cardiac constructs based on neonatal rat cardiomyocytes and porous alginate scaffolds were implanted onto myocardium of Sprague-Dawley rats that underwent permanent main coronary artery inclusion. Zimmermann et al.119 placed cardiac tissue rings cultivated in the presence of mechanical stimulation onto uninjured hearts of Fisher rats for 14 days. In an extension of cell sheet techniques,120 serial stacking of cell sheets around femoral artery and vein branches leads to a formation of *1-mm-thick vascularized cardiac tissue suitable for transplantation.121 One of the key requirements for the success of the implantation strategies is electrical coupling between the graft and host myocardium. Recent studies with the engineered heart tissue122 and cell sheets121 suggested such coupling and the possibility of functional integration between the graft and host tissue, without arrhythmias. Milica Radisic concluded that although these studies demonstrated feasibility of implantation of contractile cardiac patches, more studies are necessary to correlate the in vitro parameters to the in vivo outcomes. Questions yet to be answered through in vivo studies include (i) How long should the constructs be cultivated before implantation? (ii) How differentiated should the cardiomyocytes be? and (iii) What are the optimal cellular composition and the optimal biomaterial scaffold?

STRATEGIES FOR THE VASCULARIZATION OF 3D CONSTRUCTS Michael V. Sefton from the University of Toronto emphasized that the difficulty of supplying cells within the construct with nutrients is the major limitation for engineering large tissues. He reminded that diffusion is sufficient

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FIG. 5. Vascularization strategies. (A) Growth factors such as vascular endothelial growth factor (and platelet-derived growth factor) can be delivered by controlled release delivery systems, such as microspheres2 or cell-containing microcapsules10. (B) Microchannel networks are fabricated into silicon chips; multiple layers are used to generate arborized networks with capacity for perfusion19. (C) Materials have been discovered which induce blood vessels to grow without exogenous growth factors. (D) Endothelial cells are seeded within tissue constructs to create a vascular network21. Figure (A) adapted from Ref. 2 and from authors own work. Figure (C) courtesy of Rimon Therapeutics, Inc. Figure (D) courtesy of A. McGuigan. Figure copyright: Sabiston and Spencer Surgery of the Chest I, 7th ed., eds. P.J. del Nido, MD, and S.J. Swanson, MD, Chapter 51, pp. 817–831 (2005). Color images available online at www.liebertpub.com/ten.

for only about 100-mm-thick layers of cells, and that low cell densities can extend this limit, but at the cost of making constructs less useful. Thin or essentially 2D constructs are feasible without an internal blood/nutrient supply. However, it is hard to combine cells at high densities (*109 cells/cm3) into large tissues without some sort of prevascularization. Thus, a capillary network (and a lymphatic network) needs to be ‘‘engineered’’ as part of the creation of a larger structure. Approaches to create vascularized constructs are illustrated in Figure 5. The most studied approach is growth factor delivery to promote angiogenesis, the body’s natural process for generating new capillaries from the endothelial cells of preexisting blood vessels, with the goal of creating a localized, functional increase in vascularization. The targeted delivery of one or more of the relevant angiogenic growth factors [primarily vascular endothelial growth factor (VEGF)] is derived from studies that explore these agents for the clinical treatment of ischemic diseases. An endothelial cell-specific mitogen, VEGF is a central molecule in the stimulation of the proliferation, migration, and tube formation of endothelial cells in vitro, as well as a pronounced angiogenic response in vivo.

Growth factor delivery Mooney and colleagues have developed systems enabling controlled release of multiple angiogenic growth factors

delivery123 as well as seeding of endothelial cells into polymer scaffolds that release angiogenic factors.124,125 Sustained delivery of bioactive VEGF translated into a significant increase in blood vessel ingrowth in mice, and the vessels appeared to integrate with the host vasculature. Alternatively Hubbell and coworkers have developed a range of polymers to encourage endothelial cell migration using adhesion molecules or growth factors.126–130 The nascent vessels formed through angiogenesis must also undergo maturation, or vessel regression will occur. Remodeling and pruning of the vessels occur to produce the final, optimized vessel network pattern, which is dictated by both insoluble (ECM) and soluble cues. Maturation relies heavily on the spatial and temporal pattern and concentration of signaling cues presented to the endothelial cells. To form mature vessels therefore requires precise spatiotemporal regulation of a large number of stimulatory and inhibitory molecules. Dual delivery systems131 are a step toward this goal, and vessels formed using the delivery of both VEGF and platelet-derived growth factor (PDGF) appeared more mature than those created with the delivery of VEGF alone. However, VEGF and PDGF are but two growth factors, and issues associated with vessel maturity and the need for the controlled presentation of multiple factors at multiple time scales may limit this vascularization strategy. Furthermore it is not clear how any underlying pathology, which necessitates the need for a tissue

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replacement in the first place, may alter vessel development and maturation. The delivery of VEGF from microencapsulated cells would likely enhance the local vascularization around implanted capsules (a prototype tissue construct) and result in improved oxygen delivery and cell survival. Using microencapsulated L929 cells transfected to secrete human recombinant VEGF165 and implanted in mice using a subcutaneous matrigel plug model, VEGF delivery from encapsulated cells was found to increase the vascularization over a period of 21 days.132 MTT [3-(4,5-dimethylthiazol-2yl)-2,5-diphenytetrazolium bromide] analysis of explanted capsules suggested increased viability of encapsulated cells. It is becoming apparent that transfected cell strategies may be inadequate for providing high doses of growth factors needed initially for a robust angiogenic response, and a combination of protein and gene/cell delivery systems may be necessary.

Endothelial cell seeding Another vascularization strategy involves seeding both the endothelial cells and the tissue specific cells into the scaffold prior to implantation. Using this strategy, Atala and coworkers have shown human smooth muscle and endothelial cells seeded on biodegradable scaffolds form vascularized corpus cavernosum muscle when implanted in vivo,133 and Auger and coworkers have shown improved vascularization of artificial skin grafts in vivo.134 Koike et al.135 have implanted similar microvessel networks, formed in vitro from human umbilical vein endothelial cells (HUVECs) and mysenchymal precursor cells in a fibronectin-collagen type I gel. Networks formed from HUVECs alone showed minimal blood perfusion, while those generated from both HUVECs and mesenchymal cells integrated with the host vasculature and remained stable and functional for 1 year in vivo. The permeability of these stable vessels, however, was higher than that of normal quiescent vessels, and vessel regression could be observed in some regions of the network. Prevascularized skeletal muscle was created by coculturing skeletal muscle cells with endothelial cells and fibroblasts in a poly(lactic-glycolic) acid scaffold136 and contained blood vessels some of which ‘‘connected’’ to the host vasculature upon implantation. Vacanti and coworkers have pioneered an alternative endothelial cell-seeding approach in which microvessel capillary networks are created in microfabricated channels in vitro allowing more directed vessel formation than that is possible by simply adding endothelial cells to the construct.137,138 By seeding both hepatocytes and endothelial cells on micropatterned silicon substrates, microvessel networks embedded within hepatocyte sheets were created,139 and when implanted into omentum formed tissue consisting of both hepatocytes and bile duct-like structures by 2 weeks.140 The requirement of template wafers for large sheets could prove problematic for scale up of this approach. It is also unclear how multiple cell types can be incorporated

in a controlled manner, and blood perfusion through these engineered channels has not yet been demonstrated. Sefton and colleagues have been adapting an endothelialseeding strategy in a modular approach to create scalable vascularized tissue constructs. Endothelial cells were seeded onto sub-millimeter-sized collagen gel cylindrical modules that contained a second cell type (e.g., HepG2 or smooth muscle cells). These modules were packed into a larger tube, thereby creating interconnected channels lined with endothelial cells. These channels permitted the perfusion of whole blood, creating a means of producing uniform, scaleable tissue constructs with an internal vascular supply.141 The modules prepared using an automated cutter were 2 mm long and 0.7 mm in diameter when cut and then shrunk (after HUVEC seeding) to *0.6 mm long  0.4 mm in diameter. This modular design strategy relies critically on the endothelial cell layer on the module surface behaving in an antithrombogenic manner. The presence of HUVECs prolonged clot formation in whole blood-module mixtures and enabled blood perfusion of an assembled modular construct with no increase in platelet loss compared to background levels.

Angiogenic biomaterials Biomaterials can be designed to induce vascularization, even without the use of growth factors. Brauker et al. screened a variety of biomaterials and discovered that neovascularization occurred in membranes whose pore sizes allowed penetration by host cells (0.8–8 mm pore size). In studies comparing 0.02 and 5 mm pore size PTFE membranes, 80–100 times more vascular structures were seen at the tissue interface of the larger pore membranes.142 Boswell and Williams have also studied the vascularization of expanded PTFE (ePTFE) implants and found that removing the air trapped within the interstices of grafts with 60 mm pores reduced fibrous capsule formation and increased blood vessel formation in subcutaneous ePTFE implants.143 Similarly, Sharkawy et al. found that the tissue surrounding porous polyvinyl alcohol (PVA) implants was much more vascular than tissue surrounding either nonporous implants or normal subcutaneous tissue; the optimum size was also approximately 60 mm.144 Sefton and colleagues have also investigated the use of angiogenesis-inducing biomaterials. The polyacrylate caused vascularization in vivo, without any exogenous growth factors, presumably due to the presence of trace amounts of methacrylic acid (MAA) in the polyacrylate. A copolymer with higher MAA contents [45 M % MAA and 55% methyl methacrylate (MMA)] was tested in the form of 200 mm diameter cross-linked beads. In rats, when these beads were spread over the wound bed before full-thickness skin grafting, extensive new blood vessel growth was observed, the muscle layer in rat skin did not degenerate, and the skin graft remained viable, looking much like native tissue.145 Control

ADVANCED TOOLS FOR TISSUE ENGINEERING poly(methyl methacrylate) beads lead to necrosis of the muscle layer and the entire graft. Michael V. Sefton concluded that the critical importance of vascularization for tissue engineering has been recognized and multiple strategies are being explored. All are in early stages of exploration, and it is not yet clear which approach will work best and for which situation. The lesson from exploring therapeutic angiogenesis is that translation from the concept to animal model and from animal model to clinic is long and complex. To make tissue engineering successful, much effort will be required, and there is considerable opportunity for tissue engineers to make the transition from in vitro construct fabrication to addressing the consequences of in vivo biology.

ACKNOWLEDGMENTS Drs. Freed and Guilak thank Frank Moutos, Jan Boublik, Enrico Tognana, and Robert Padera for their contributions to the studies they described; Robert Langer for advice; Suzanne Kangiser for help in manuscript preparation; and NASA (NNJ04HC72G) and NIH (P41 EB002520, AR49294, AG15768, AR50245, and AR48182) for funding. Dr. Guo thanks NIH/NIAMS for funding (AR048287). Dr. Holmes thanks NIH/NHLBI (HL075639). Dr. Sefton thanks Alison McGuigan and Jennifer Vallbacka for help with the preparation of his part of the manuscript.

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