Micromechanical Design Criteria for Tissue Engineering Biomaterials

Chapter II.6.5  Micromechanical Design Criteria for Tissue ­Engineering Biomaterials

CHAPTER II.6.5 ­MICROMECHANICAL DESIGN CRITERIA FOR TISSUE ­ENGINEERING BIOMATERIALS Kaustabh Ghosh1, Charles K. Thodeti1, and Donald E. Ingber1–3 1Vascular

Biology Program, Departments of Pathology and Surgery, Children’s Hospital, Harvard Medical School, Boston, MA, USA 2Harvard School of Engineering and Applied Sciences 3Wyss Institute for Biologically Inspired Engineering at Harvard ­University, Boston, MA, USA

INTRODUCTION The field of tissue engineering emerged from work striving to construct artificial tissues and organs in vitro by integrating the principles of cell biology, polymer chemistry, and materials science (Langer and Vacanti, 1993). In this approach, cells isolated from whole organs (e.g., liver, kidney) or multi-potential stem cells are cultured within three-dimensional scaffolding materials, which can be designed in myriad forms to provide the necessary structural and chemical cues to induce cell growth and tissue differentiation. This cell–scaffold construct is then implanted in vivo with the aim of promoting new tissue formation following injury or disease. Skin was the first to be artificially constructed using natural scaffold materials in combination with patientderived skin cells (Bell et  al., 1981). Development of tissue-engineered skin is less challenging than constructing artificial internal organs, such as whole liver or kidney, because the relatively thin skin constructs can be implanted directly onto a pre-existing vascularized wound bed. In contrast, cells within larger three-­ dimensional cell–scaffold constructs from many solid organs (e.g., liver) die rapidly when implanted in  vivo without prior vascularization, due to insufficient delivery of oxygen and nutrients. Interestingly, despite their simplicity, artificial skin grafts have not made a significant impact on the market. This is, in part, due to their failure to maintain optimal cell viability; however, the major difficulty has arisen from the tough regulatory and cost challenges involved in the development of medical products containing living cells. These limitations highlight the importance of exploring alternative materials-based tissue engineering approaches that rely on development of more sophisticated biomimetic scaffolds to promote angiogenesis, and induce recruitment and reprogramming of endogenous multipotent progenitor cells (e.g., either resident within remaining tissues or bone marrow-derived cells that circulate in the blood) when implanted in  vivo. To do this effectively, however, we must fully-understand and recapitulate the dynamic and reciprocal interactions that occur within the cell–ECM ­microenvironment during normal tissue formation. Biomaterials used in tissue engineering were initially viewed as inert carriers for cell delivery. However,

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modern strategies for regenerative medicine now employ methods to manipulate their biochemical and biophysical properties, in order to best promote orderly tissue growth and restore normal function. Materials that more closely mimic the inductive environment of normal developing tissues have been fabricated using natural ECM proteins (e.g., collagen, fibrin, hyaluronan, etc.), as well as synthetic polymers (PLGA, PEG, PVA, etc.) (Lutolf and Hubbell, 2005). To enhance cell adhesion, these polymeric scaffolds have also been covalently derivatized with proteins or peptides found within natural ECMs (Ghosh et al., 2006). The RGD (arginine-glycineasparatic acid) tripeptide sequence is the most commonly used cell adhesion moiety (Hersel et al., 2003; Shu et al., 2004); however, other peptides such as the YIGSR and IKVAV peptides of laminin and VPGIG sequence of elastin have also been explored (Ranieri et  al., 1995; Panitch et al., 1999; Lin et al., 2006). Importantly, these ECM-mimicking biomaterials support key cell functions, including cell spreading, migration, growth, and differentiation in  vitro, as well as tissue repair in  vivo (Shin et al., 2003; Lutolf and Hubbell, 2005). To better mimic the fibrillar microarchitecture and anisotropy (alignment) of native ECMs, specialized fabrication techniques have been developed, such as electrospinning and three-dimensional weaving that produce fibrous biomaterial scaffolds with varying pore size and volume, ranging from macro- to nanoscale dimensions (Yang et  al., 2001; Muschler et  al., 2004; Tsang and Bhatia, 2007). These scaffolds may be rendered stable to hydrolytic or proteolytic degradation via inter- or intramolecular cross-linking, and cell recognition moieties can be chemically tethered to their polymer backbone to promote cell adhesion. To engineer mechanically active tissues that bear continuous cyclical hemodynamic and compressive loading (e.g., heart valve and cartilage, respectively), cell–biomaterial constructs are preconditioned prior to in vivo implantation by being cultured in specialized bioreactors that simulate the physical environment of native tissues (Waldman et al., 2003; McCulloch et  al., 2004; Fehrenbacher et  al., 2006; Isenberg et  al., 2006). This mechanical force regimen significantly alters ECM structure, as well as cellular growth and function. Despite efforts to mimic the properties of natural ECMs through various chemical and mechanical modifications, most tissue-engineered constructs fail to completely restore normal tissue form and function when implanted in  vivo. Our inability to produce ideal tissue substitutes may be due to the failure to address the role of micromechanical forces in guiding cell and tissue development. These microscale forces result from the tension (contractility) generated in the cell cytoskeleton, which is exerted on the cells’ adhesions to ECM and to neighboring cells. Cells sense these forces through transmembrane integrin and cadherin receptors that mediate cell–ECM anchorage and cell–cell adhesion, respectively (Chen et al., 2004). Mechanical stresses acting on these

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cell surface receptors are transmitted across the cell surface, and to internal cytoskeletal and nuclear scaffolds via specialized membrane adhesion complexes (e.g., focal adhesions, adherens junctions) (Wang et al., 2009). Mechanotransduction – the process by which mechanical signals are transduced into changes in cellular biochemistry and gene expression – can occur in these adhesion complexes or at many other locations along loadbearing structural elements in the cell where force induces molecular distortion (Ingber, 2006; Vogel and Sheetz, 2006). Importantly, these cell-generated forces influence various cell behaviors critical for developmental control, including growth, movement, differentiation, apoptosis, and stem cell fate switching (Chen et  al., 1997; Engler et al., 2006; Ghosh et al., 2007). In this chapter, we describe how micromechanical interactions between cells and ECM govern normal tissue morphogenesis and control cell fate switching. We also review various microscale design approaches to biomaterial fabrication that leverage this critical developmental paradigm. Finally, we explore how our growing knowledge of micromechanical control of cell function and biomaterials science can be integrated to develop improved biomimetic microenvironments for in situ tissue repair and regeneration.

MICROMECHANICAL CONTROL OF TISSUE FORM AND FUNCTION Tissues and organs arise during embryogenesis from coordinated self assembly of large populations of cells that organize themselves into precise three-dimensional spatial patterns. Multicellular self-organization is facilitated by molecular self assembly of cell-derived proteins (e.g., collagen, laminin, fibronectin) into fibrous multimolecular ECM scaffolds that function as physical templates for orderly cell attachment and renewal (Vracko, 1974; Leivo, 1983). Cells attach to this anchoring ECM scaffold through transmembrane integrin receptors (Ruoslahti, 1991; Hynes, 1992), and they exert cytoskeleton-generated traction forces on these cell–ECM adhesions, as well as on their attachments to neighboring cells (Ingber, 1991; Chicurel et al., 1998a; Galbraith and Sheetz, 1998; Geiger et  al., 2001; Ingber, 2003c). The resistance produced by cell–ECM and cell–cell linkages to cytoskeletal forces maintains the cell in a state of isometric tension or prestress that stabilizes cell and tissue form through a tensegrity mechanism (Ingber and Jamieson, 1985; Ingber, 1993, 2003b; Stamenovic and Ingber, 2009). The importance of ECM-dependent micromechanical forces in regulating tissue form is evident from observations of new epithelium formation in the early embryo and in specialized organs (e.g., kidney), which is always accompanied by concomitant formation of planar ECM scaffolds or “basement membranes,” that promote consistent cell orientation and generation of epithelial form (Leivo, 1983; Yurchenco and Ruben, 1988).

ECM is also centrally involved in the morphogenetic patterning that occurs during embryonic organogenesis. For example, formation of lobular epithelial glands in salivary gland and lung, as well as branching capillary networks, appear to require establishment of local differentials in ECM remodeling (Bernfield and Banerjee, 1978; Moore et  al., 2005). In general, regions of highest ECM turnover, which lead to basement membrane thinning, exhibit the highest rate of cell growth and tissue expansion, while cells in the neighboring areas (only several microns away) that experience increased ECM accumulation (due to low degradation relative to synthesis) remain quiescent. These microscale differentials in cell growth and ECM density drive the budding, folding, and branching of many epithelial and endothelial tissues (Ausprunk and Folkman, 1977; Bernfield and Banerjee, 1978). These observations have formed the basis of a micromechanical model of tissue development (Figure II.6.5.1), which postulates that morphogenetic events such as epithelial budding and branching are controlled by local micromechanical variations that lead to spatial differentials in the cellular force balance (Ingber and Jamieson, 1985; Huang and Ingber, 1999). In this model, local enzymatic degradation of the prestressed (tensed) basement membrane causes it to stretch and thin in regions of high turnover, with concomitant distortion of the adherent epithelial cells. These stretched cells become more responsive to soluble mitogens and undergo proliferation (Singhvi et  al., 1994; Chen et  al., 1997; Folkman and Moscona, 1978) compared to their neighboring counterparts that remain quiescent in regions experiencing basement membrane accumulation. Reiteration of this micromechanical control scheme that couples ECM stretching and cell growth could then produce the complex fractal-like tissue patterns that characterize development of virtually all organs (Huang and Ingber, 1999; Moore et  al., 2005). Stable tissue forms finally emerge when the stabilizing tissue prestress is restored by accumulating thickened ECMs and strengthened cell–cell adhesions that fully balance cell-generated tensional forces (Ingber and Jamieson, 1985; Hardin and Keller, 1988). This micromechanical model of tissue morphogenesis is supported by the finding that changing the cell–ECM force balance within embryonic mouse lung rudiments alters branching morphogenesis (Figure II.6.5.1) (Moore et  al., 2005). Specifically, suppressing cytoskeletal tension generation inhibits regional thinning of the basement membrane at the tips of expanding buds and interferes with epitheliogenesis, while increasing tension accelerates epithelial branching morphogenesis (Moore et  al., 2005). These treatments also had similar stimulatory and inhibitory effects on formation of neighboring capillary networks, as well as on the growth of the entire organ. Importantly, this force-dependent spatial patterning was not mediated by alterations in the overall rate of

Chapter II.6.5  Micromechanical Design Criteria for Tissue ­Engineering Biomaterials

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FIGURE II.6.5.1  Micromechanical control of tissue morphogenesis. (Left) Schematic diagrams of a micromechanical model of epithelial mor-

phogenesis (Huang, S. & Ingber, D. (1999). Nat Cell Biol) showing progressive development of a simple planar epithelium into complex branching patterns (top to bottom). Local increases in basement membrane degradation (green), coupled with traction generated by overlying adherent cells (white arrows), result in local thinning and stretching of the basement membrane, as well as the overlying adherent epithelium (red), whereas neighboring cells only microns away remain quiescent (white). Cell distortion enhances cell sensitivity to soluble growth factors, which results in increased cell proliferation in regions of highest ECM turnover. Spreading and growing cells simultaneously deposit new ECM, which causes local basement extension and new bud formation. Reiteration of this process over time and space leads to the development of complex tissue architecture, with characteristic fractal-like patterns. (Center) Photomicrographs of individual epithelial buds in developing mouse lung epithelium corresponding to the stages shown at left (Moore, K. et al., (2005). Dev Dyn). (Right) Theoretical mechanical strain distributions within the basement membrane of the developing mouse lung epithelium at corresponding times during branching morphogenesis. Increased spacing between the strain field lines indicates regions where the basement membrane thins and experiences greatest mechanical strain (distortion). These regions of increased strain correlate precisely with regions of epithelial expansion and new bud formation (Moore, K. et al. (2005). Dev Dyn). (Reproduced from Ghosh, K. & Ingber, D., (2007). Adv. Drug Deliver. Rev. 59: 1306–1318, with permission from Elsevier ©2007).

cell proliferation within the mesenchyme and epithelium, but rather by establishment of local growth differentials within the forming gland (Moore et al., 2005). Thus, although tissue development is driven by soluble

cytokines and changes in gene expression, it is also a highly mechanical process that is significantly influenced at the microscale by alterations in the mechanical force balance between cells and their surrounding adhesions to

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ECM and neighboring cells. The tissue architecture generated by this mechanical patterning mechanism also feeds back to alter gradients of soluble developmental cues (Nelson et al., 2006), thereby establishing a vital feedback mechanism that further integrates changes in tissue form and function.

MICROSCALE DESIGN OF BIOMIMETIC SCAFFOLDS FOR TISSUE RECONSTRUCTION Given the critical role of external physical cues and cellgenerated forces in regulating tissue form and function, it is paramount that these factors be considered in the design of rational tissue engineering approaches aimed at restoring the normal structure and functionality of damaged tissues and organs. Tissue engineering scaffolds have been developed using various types of polymeric materials (natural, synthetic, and semi-synthetic) and techniques (gas foaming, phase separation, etc.) (Figure II.6.5.2A) (Agrawal and Ray, 2001; Lutolf and Hubbell, 2005). The design challenges focused on in the past lay largely in achieving appropriate macroscopic features, such as bulk stiffness and suitable degradation

rate (to facilitate integration with host tissue), porosity (to promote nutrient exchange and tissue ingrowth), and adhesiveness (for cell adhesion and function). But more recent studies have shown that cell function within such bulk scaffolds can be more tightly regulated by introducing precise micro- or nanoscale structural features. For instance, microporous biopolymer scaffolds (pore size ~100–500 μm) induce cells to exhibit a flattened cell morphology similar to that seen when they are cultured on a two-dimensional tissue culture substrate, while nanofibrous scaffolds (Figure II.6.5.2B) (fiber diameter ~50–100 nm; pore size ~0.5–5 μm) cause these cells to assume a dendritic morphology similar to that seen in native tissues (Grinnell, 2003; Stevens and George, 2005; Ji et  al., 2006). Further, creating nanoscale features within an otherwise microporous scaffold alone can markedly enhance cell spreading, growth, and expression of tissue-specific ECM components (Pattison et al., 2005). Such apparent differences in cell morphology likely arise from the differential ability of these topographies to induce distinct membrane protrusions (e.g., lamellipodia formation on flatter surfaces while filopodia form on more fibrous ones) (Cukierman et al., 2001; Curtis et al., 2004).

FIGURE II.6.5.2  Microscale design of biomimetic scaffolds for tissue reconstruction. Increasing knowledge of the influence of microscale

ECM material properties on cell and tissue development has led to novel physical design strategies for tissue engineering biomaterials. (A) Initial attempts to engineer three-dimensional tissues used macroporous scaffolds. Their large surface area allows greater cell spreading, while the large pore size facilitates sufficient metabolic exchange (Eiselt, P. et al. (2000). Biomaterials 21: 1921–1927). (B) Electrospun scaffolds that better mimic the fibrous architecture of natural ECM are being increasingly used nowadays. Cells cultured on these scaffolds assume dendritic morphology similar to that exhibited by cells within native tissues (Ji, Y. et al. (2006). Biomaterials). (Reproduced with permission from Elsevier ©2006). (C) More sophisticated design strategies involving automated CAD/CAM programs with solid free-form fabrication approaches allow a bottom-up approach to scaffold construction, where the microscale material properties can be controlled in a precise spatiotemporal manner (Seitz, H. et al. (2005). J Biomed Mater Res). (D) Three-dimensional weaving technology has been recently used to produce fibrous scaffolds that mimic the anisotropy, porosity, and bulk mechanical properties of native ECM (Moutos, F. et al. (2007). Nat Mater). (E) Self-assembling molecules have been engineered that can spontaneously organize into complex three-dimensional structures in a manner reminiscent of the bottom-up molecular assembly of natural ECM materials (Hartgerink, J. et al. (2001). Science). (Reproduced with permission from AAAS ©2001.)

Chapter II.6.5  Micromechanical Design Criteria for Tissue ­Engineering Biomaterials Because local variations in ECM structure and elasticity can trigger differential cell responses required for tissue morphogenesis and development (as seen during lung development), it would be desirable to engineer biomaterial scaffolds that exhibit such structural and mechanical variations across small length scales. One way to address this challenge is to use solid free-form fabrication (SFF) techniques (e.g., three-dimensional microprinting, soft lithography, etc.) that enable the development of biomaterials with complex three-dimensional microscale features using computer-assisted design and manufacturing (CAD/CAM) platforms (Figure II.6.5.2C). In this bottom-up approach, thin sequential layers of polymer are deposited one-at-a-time, with each layer having a prescribed physical structure and composition (Giordano et  al., 1996; Seitz et  al., 2005). By depositing polymer solutions at precise predetermined locations, and by varying the type and concentration of polymers used, the architecture and mechanical properties (stiffness, degradation rate, etc.) of the scaffold can be tightly controlled with microscale resolution. Such an approach can produce anisotropic scaffolds that better recapitulate the physical complexity of loadbearing tissues such as cartilage, which is composed of distinct tiers of cell and ECM organization. If so, this may lead to significant improvement in both scaffold integration at the defect site and tissue function. Such biomaterials have, in fact, been fabricated using a novel microscale threedimensional weaving technology where the resultant porous scaffold closely resembles native cartilage, not only in terms of bulk (compressive, shear, and tensile) moduli, but also with respect to its physical anisotropy (tension-compression nonlinearity) (Figure II.6.5.2D) (Moutos et al., 2007). Specifically, small-pore scaffolds (pore dimensions 390 μm × 320 μm × 104 μm) produced by designing a biased woven structure that contained a higher fiber volume fraction in the weft direction than in the warp direction exhibited ~35% higher ultimate tensile stress when tested in the weft direction than in the warp direction, which is reminiscent of the mechanical anisotropy seen in native articular cartilage. Native ECM is produced through complex self assembly reactions involving various natural biopolymers, such as collagen, elastin, and hyaluronic acid, among others. Due to this systematic bottom-up approach, the ECM displays a remarkable level of hierarchical organization that spans from nano- to mesoscale. Artificially replicating this self-assembly process may produce scaffolds that better mimic both the material and biological properties of native ECM. To this end, novel polypeptide systems have been developed that spontaneously selfassemble in physiological solvents to produce nanoscale, branched fibrous networks analogous to those seen in native ECM (Ryadnov and Woolfson, 2003). Peptides have also been modified to contain alternating hydrophilic and hydrophobic groups that enable them to interact with their complementary moieties to form

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β-sheet membranes, which further self-assemble to form higher-order structures (Zhang, 2003). Hydrogel scaffolds developed using such techniques can successfully encapsulate cells, as well as enhance cell viability and expression of tissue-specific ECM components (Kisiday et al., 2002). In another novel, and potentially valuable, application of this ­self-assembly process, amphiphilic molecules have been precisely engineered such that each nanofiber presents a template for hydroxyapatite mineralization, as well as osteoblast adhesion, which together enhance bone regeneration (Hartgerink et al., 2001) (Figure II.6.5.2E). An interesting feature of this technique is the ability of nanofibers to direct alignment of hydroxyapatite crystallization, which mimics the collagen fiberinduced hydroxyapatite alignment seen in native bone. The apparent effects of scaffold structure on cell form and function may be due to differences in membrane receptor clustering and binding that result from alterations in the physical context in which adhesive cues are presented to cells. Varying the nanoscale (~50 nm) distribution of the integrin-binding RGD tripeptide on an otherwise non-adhesive two-dimensional surface reveals that cell spreading and motility on lower ligand density surfaces (~103–104 RGD molecules/μm2) occur best when RGD is presented in clusters of nine ligands per cluster (Maheshwari et al., 2000). Notably, this influence of RGD clustering on cell function is markedly weakened when RGD is presented either at saturated levels (105 molecules/μm2) or as individual peptides (Maheshwari et  al., 2000). This functional difference correlates with the ability of RGD clusters to induce actin stress fiber formation (Maheshwari et  al., 2000), confirming that adhesive ligands influence cell behavior by promoting optimal traction-mediated cell spreading and simultaneously activating integrin-dependent biochemical signaling inside the cell (Ingber, 2003c). Achieving this precise nanoscale control over adhesive signaling is desirable in biomaterial scaffolds, as that will allow the creation of well-defined three-dimensional cellular niches. To this end, poly(ethylene) glycol scaffolds containing separate RGD and PHSRN peptide sequences spaced precisely 4 nm apart have been developed that promote osteoblast growth and metabolism, while inhibiting ECM deposition, whereas different RGD-PHSRN spacings are less effective (Benoit and Anseth, 2005).

CELL AND ECM MECHANICS AS KEY REGULATORS OF TISSUE DEVELOPMENT The mechanism of control of tissue development and related tissue engineering approaches described above are based on the concept that local changes in ECM structure and mechanics result in physical distortion of adjacent adherent cells. Cell distortion alters the level of cell prestress and deforms the cytoskeleton, which changes the cell’s ability to respond to soluble growth factors and morphogens. To directly test this concept that cell shape

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distortion is critical for functional control, we applied microfabrication techniques first developed to make microchips for the computer industry to microengineer adhesive ECM islands with controlled size, shape, and position on the micrometer scale ­(Whitesides et al., 2001). Using this approach, cell shape can be controlled independently of the nature and density of immobilized ECM molecules or the presence or absence of soluble growth factors. Specifically, we created single cell-sized adhesive islands coated with a saturating density of ECM molecules (e.g., fibronectin) that were surrounded by non-adhesive regions. Cells spread by adhering to ­immobilized ECM ligands and exerting traction forces on these adhesions. Thus, these microengineered adhesive islands provided a simple method to control cell shape distortion, because cell spreading was limited only to the area of the ECM island (Singhvi et al., 1994; Chen et al., 1997). When single capillary endothelial cells or primary hepatocytes were plated on each island, they spread and flattened, eventually taking on the precise size and shape (e.g., circle, square, hexagon, etc.) of the adhesive island. Most importantly, in the presence of optimal soluble mitogens, cells that distorted (spread) exhibited the highest growth rates (Singhvi et al., 1994; Chen et  al., 1997), whereas round endothelial cells adherent to the same ECM coating underwent apoptosis (Figure II.6.5.3). Interestingly, cells that were cultured on intermediate size islands exhibited enhanced cytodifferentiation (secretion of blood proteins by liver epithelial cells) as well as histodifferentiation (e.g., tube formation by capillary cells) (Singhvi et al., 1994; Dike et al., 1999).

In addition, when cells cultured on square-shaped ECM islands were stimulated with motility factors (e.g., PDGF, FGF), they preferentially extended lamellipodia at their corners, which are also the sites of maximal cytoskeletal tension and focal adhesion density (Parker et al., 2002). Subsequent studies revealed that this form of cell shape-dependent developmental control results from cytoskeletal distortion and alterations in the level of cytoskeletal tension within the cell. For example, the Rho family GTPase, RhoA, mediates cell shape-dependent growth regulation by altering the balance of activities between its downstream effector proteins mDia1 and ROCK that are involved primarily in cytoskeletal remodeling (Mammoto et al., 2004). Dissipation of cytoskeletal tension using pharmacological or genetic methods inhibits cell distortion-dependent cell cycle progression, stimulates apoptosis, and suppresses lamellipodia ­formation (Parker et al., 2002; Numaguchi et al., 2003). The mechanical properties of the ECM can also regulate cell shape, as well as these same behaviors including cell growth, motility, and differentiation, and again these effects are mediated at least in part through modulation of cytoskeletal tension. For example, cells plated on rigid ECM substrates generate higher levels of tension and proliferation, whereas softer substrates reverse this phenotype (Wang et al., 2000; Ghosh et al., 2007). ECM rigidity can also influence tissue-specific differentiation (Engler et al., 2004, 2006), as well as guide directional cell motility (durotaxis) by regulating cell-substrate mechanical interactions that lead to spatial differentials in cell traction forces and focal adhesion dynamics

FIGURE II.6.5.3  ECM-mediated geometric control of cell shape regulates key developmental programs. Shown here is a schematic summary

of experiments where microfabricated fibronectin islands (solid green) of desired shape and size were used to control cell shape and spreading. Under the influence of the same growth factors (light green dots) and ECM composition, round cells die (apoptosis) and spread cells proliferate, while those partially spread become quiescence (Huang, S. & Ingber, D. (1999). Nat Cell Biol; Chen, C. et al. (1997). Science). Note that cells plated on the small 5 μm dots in the center are in contact with approximately the same absolute amount of fibronectin as on the small 30 μm islands, yet they proliferate because they are allowed to extend over the ECM dots. The first four panels in the bottom row show fluorescence micrographs of phalloidin-stained actin “stress fibers.” Allowing endothelial cells to extend on 10 μm-wide linear ECM arrays (right panel) stimulates tube formation, a manifestation of endothelial cell differentiation, as verified using confocal fluorescence microscopy of cytoplasmstained cells (Dike, L. et al. (1999). In Vitro Cell Dev Biol Anim). (Reproduced from Huang, S. & Ingber, D. (2006). Breast Disease, 27–54, with permission from IOS Press ©2006, 2007.)

Chapter II.6.5  Micromechanical Design Criteria for Tissue ­Engineering Biomaterials (Pelham and Wang, 1997; Lo et al., 2000). Importantly, matrix stiffness alone can control three-dimensional tissue morphogenesis, as normal mammary gland development commonly seen in compliant three-dimensional gels is disrupted when the gel stiffness is raised from 170 Pa to 1200 Pa (Paszek et al., 2005). Likewise, varying ECM elasticity from 700 Pa to 900 Pa can significantly alter three-dimensional endothelial capillary formation in vivo, with the intermediate stiffness (800 Pa) promoting the highest VEGF receptor expression and capillary ingrowth (Mammoto et al., 2009). Thus, it is critical to take into account the importance of mechanical features of the tissue microenvironment in future studies focused on engineering artificial inductive scaffolds.

BIOMATERIALS FOR STEM CELL DEVELOPMENT AND TISSUE REGENERATION Engineered biomaterials that aim to promote complete regeneration of diseased or injured tissues and organs must recapitulate the developmental programs that drive endogenous tissue formation and function. This goal can be accomplished by developing biomimetic scaffolds that leverage the regenerative potential of stem and multipotent progenitor cells – the primitive cells that retain an ability to differentiate into mature cells of various tissue types. Since cell and ECM mechanics alone can regulate the development and function of mature tissue cells, it is conceivable that stem cell differentiation and functionality similarly depend on its physical microenvironment. Indeed, stem cells implanted in the absence of proper mechanostructural cues form teratomas (Wakitani et al., 2003; Hentze et al., 2007), which result from stem cells growing and differentiating into multiple different lineages in the absence of higher-order pattern controls. Thus, a major challenge in regenerative medicine will be to develop biomaterials that can provide appropriate physical cues required to precisely drive stem cell differentiation along a specific lineage, and stimulate regeneration in damaged tissues and organs. In this regard, it is important to note that mechanical cues such as ECM topography and stiffness alone have been shown to direct mesenchymal stem cell (MSC) lineage specification. When human MSCs are grown on single cell-sized microfabricated ECM islands and exposed to a cocktail of soluble differentiation-inducing factors, adipogenic differentiation is favored on smaller (~1000 μm2) ECM islands, while osteogenesis is induced in cells spread on larger (~10,000 μm2) islands (McBeath et al., 2004). Notably, this shape-dependent lineage specification is mediated by cytoskeletal organization and tensile prestress as disruption of Rho activity and actin assembly favors adipogenic differentiation while inhibiting osteogenesis; moreover, overexpressing Rho has an opposite effect. Even more striking is the finding that this Rhomediated lineage switching occurs independently of the

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appropriate differentiation-inducing medium (McBeath et al., 2004), suggesting that cell shape and tension alone can govern which cues are necessary and sufficient for stem cell commitment. The nanotopography of biomaterial surfaces can similarly regulate stem cell differentiation. When human MSCs are cultured on titanium oxide (TiO2) nanotubes, they adhere strongly but fail to differentiate when the nanotubes are small in size (~30 nm diameter), whereas they differentiate into bone on larger nanotubes (~70– 100 nm diameter) that promote cytoskeletal stressdependent changes in cell form and function (Oh et al., 2009). MSCs grown on ECM-coated flexible substrates can also be directed towards different lineages by tuning the substrate stiffness to match the elasticity of the whole living tissue. For example, neurogenesis is induced on softer substrates (0.1–1 kPa) that mimic brain’s elasticity; stiffer substrates (8–17 kPa) that mimic muscle mechanical properties promote myogenic differentiation; and bone induction is observed on the stiffest substrates (25–40 kPa) that matched bone elasticity (Engler et al., 2006). Again, this elasticity-driven lineage switching is dependent on cytoskeletal tension, as disrupting nonmuscle myosin II activity abolishes this effect (Engler et  al., 2006) in a manner reminiscent of Rho-mediated cytoskeletal control of shape-dependent lineage specification (McBeath et  al., 2004). Neural stem cell (NSC) differentiation can similarly be controlled in an ECM stiffness-dependent manner; compliant substrates (~100– 500 Pa) favor neuron formation, whereas stiffer hydrogels drive glial cell formation (Saha et al., 2008). Thus, biomaterials developed for tissue engineering and regeneration must be tailored to provide the correct mechanostructural cues that promote desired tissue functionality while suppressing undesirable responses. Embryonic stem (ES) cells, the pluripotent cells that can give rise to derivatives of all three embryonic germ layers (i.e., endoderm, mesoderm, and ectoderm), possess greater regenerative potential when compared with MSCs and, as such, they are being increasingly explored for cell-based therapies. Notably, human ES (hES) cell self-renewal and differentiation are remarkably sensitive to the physical dimensions of the culture microenvironment. Cell colonies generated from single hES cells initially cultured on micropatterned ECM islands of varying sizes revealed that smaller (200 µm diameter) hES cell colonies preferentially undergo endodermal differentiation, while larger (1200 µm) colonies differentiate into mesoderm (Bauwens et al., 2008; Lee et al., 2009). However, high mesodermal and cardiac induction is observed when the smaller endodermal-biased hES cell colonies are grown into embryoid bodies of large sizes, suggesting that the size of both the cell colony and embryoid body regulate hES cell fate determination (Bauwens et  al., 2008). Taken together, these findings suggest that mechanical interactions at the cell–cell and cell–ECM interface guide tissue development not only by

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facilitating morphogenetic shape changes (as seen during lung branching), but also through precise spatiotemporal control of stem cell fate commitment. To be effective, tissue and organ regeneration strategies must simultaneously promote the formation of a robust vascular network to maintain optimal oxygen delivery and metabolic exchange. But, despite years of research, achieving this single functionality still poses a huge challenge for tissue engineers. Past in vitro studies have shown that heterotypic interactions between endothelial cells and mesenchymal precursor cells, mediated by specific growth factors such as TGF-β and PDGF, cause significant enhancement in vessel stabilization through mesenchymal differentiation into mural cells (pericytes or smooth muscle cells) that are ultimately recruited to the perivascular niche (Hirschi et al., 1998; Ding et al., 2004). Indeed, when such co-cultures are grown in threedimensional ECM gels and implanted in animals, longlasting and robust vessels form that remain stable up to one year (Koike et al., 2004). hES cell-derived endothelial cells have also been co-cultured with mouse myoblasts and embryonic fibroblasts within three-dimensional porous polymeric scaffolds to obtain prevascularized tissue constructs in vitro, with improved vascular network and blood perfusion in vivo (­Levenberg et al., 2005). However, given the complexity of culturing multiple cell types in a controlled and reproducible manner, alternate approaches have been developed that use biomaterial scaffolds to release single or multiple angiogenic growth factors (e.g., VEGF, PDGF) in a spatiotemporally controlled manner to induce new vessel formation. These materials stimulate rapid and robust capillary ingrowth and maturation, as well as increased blood perfusion and improved tissue functionality (Richardson et  al., 2001; Nillesen et al., 2007). More recent studies show that in addition to soluble cues (e.g., VEGF) scaffold elasticity can exert a significant effect on three-dimensional capillary formation by endothelial cells in  vivo via transcriptional control of VEGF-receptor expression. In this study, ECMs with intermediate stiffness (800 Pa) promoted the highest VEGF-receptor expression and capillary ingrowth, compared with both softer (700 Pa) and stiffer (900 Pa) scaffolds in vivo (Mammoto et al., 2009). In addition to vascular density, the three-dimensional spatial patterning of the newly formed endothelial capillaries can also be controlled by providing appropriate mechanical cues. For instance, application of exogenous cyclic stretch to endothelial cell-seeded polymeric scaffolds results in three-dimensional sprouting in a direction perpendicular to the stretch axis (Matsumoto et al., 2007). Thus, formation of a robust vasculature within engineered tissues will benefit from the use of instructive biomimetic scaffolds that can effectively integrate and present both soluble and physical cues in a precisely controlled manner. In addition to their pivotal role in cell-based regenerative approaches, biomaterials are also crucial for targeted

delivery of therapeutic drugs and genes that, in turn, has direct implications for regenerative medicine. The physical properties of biomaterials are as important for optimizing drug and gene delivery, as they are for control of tissue development (Mitragotri and Lahann, 2009). Cellular uptake of injected microparticles, for instance, is greatly influenced by particle shape, where rod-like particles with an intermediate aspect ratio of 3 are preferably internalized relative to their spherical counterparts (Decuzzi and Ferrari, 2008; Gratton et al., 2008). Subsequent theoretical modeling and experimental work have revealed that particle internalization is greatly favored with increasing contact angle between the particle and the cell surface (which is high for cylindrical particles and low for spherical ones), which influences the development of a complex actin structure within the cell required for particle internalization (Champion and Mitragotri, 2006). Oblate particles are also expected to adhere more strongly to the vascular endothelium than spherical particles of the same volume (Decuzzi and Ferrari, 2006; Muro et al., 2008), a feature that has important implications for drug delivery to blood vessels. In addition, the degree of internalization depends on particle elasticity, as softer particles are less likely to be phagocytosed by macrophages, and thus have a longer lifetime in the circulation (Beningo and Wang, 2002; Geng et al., 2007). The efficiency of gene delivery to cells similarly depends on the elasticity of cell adhesive hydrogels, and the uptake and expression of plasmid DNA were shown to be higher on stiffer substrates (Kong et al., 2005).

MOLECULAR MECHANISMS OF CELLULAR MECHANOTRANSDUCTION It is now clear that cells can actively sense and respond to physical cues from the ECM by modulating their level of contractility, which further alters ECM mechanics, thereby establishing a feedback loop that ultimately controls cell fate and function. This dynamic reciprocity of cell–ECM mechanical interactions is mediated by transmembrane integrin receptors that transmit mechanical forces across the cell surface and facilitate mechanochemical transduction events that control cell function and govern cell fate (Alenghat and Ingber, 2002). Integrins refer to a family of heterodimeric transmembrane proteins that bind various ECM molecules on the external surface of the cell (Hynes, 2002) and interact with actin-binding proteins (e.g., vinculin, paxillin, and zyxin) within focal adhesion anchoring complexes inside the cell (Zamir and Geiger, 2001; Ingber, 2003c) ­(Figure II.6.5.4A). Focal adhesions play a central role in mechanochemical transduction, because in addition to physically coupling the ECM to actin cytoskeleton and bearing high mechanical loads focused through integrins, they also contain many signaling molecules that alter their activity when mechanically stressed (Chicurel et al., 1998b; Alenghat and Ingber, 2002).

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FIGURE II.6.5.4  (A) Integrin-associated focal adhesions mediate cellular mechanotransduction. Forces applied to cells via their basal ECM adhesions

are transmitted across cell surface integrin receptors to their cytoplasmic domains down to the underlying focal adhesion scaffolds that link to the cytoskeleton. Forces such as fluid shear stresses, which act on the apical cell surface and produce generalized distortion can also result in increased tension on the cell’s basal focal adhesions that are stiffened relative to the remainder of the cell and cytoskeleton. Internally-generated tension (contractility) and forces transmitted via intercellular junctions similarly reach focal adhesions through the cytoskeleton. Forces concentrated within the focal adhesion can stimulate clustering of dimeric (α,β) integrin-receptors, and induce recruitment of focal adhesion proteins (e.g., Vinculin (Vin), Paxillin (Pax), Talin (Tal)) that connect directly to actin microfilaments and indirectly to microtubules and intermediate filaments. Forces applied to this specialized focal adhesion complex activate various integrin-associated signaling cascades, including focal adhesion kinase (FAK), extracellular signal-regulated protein kinase (ERK), Rho, mDia1, heterotrimeric G-proteins, and protein kinase A (PKA), among others, which ultimately regulate gene expression and thereby cell behavior (Ingber, D. (2003c). J Cell Sci). (B) Tensegrity model as the architectural basis of cellular mechanotransduction. (Top) A schematic diagram of the tensegrity-based complementary force balance between tensed microfilaments, compressed microtubules, and transmembrane integrin-receptors (gray oval dimer) in living cells (intermediate filaments are not shown for simplicity). Black forms indicate regulatory proteins and enzymes that are physically immobilized on load-bearing cytoskeletal filaments; red oval represents a transmembrane protein that does not link to the internal cytoskeletal lattice. (Bottom) When force is applied to integrins, thermodynamic and kinetic parameters change locally for cytoskeletonassociated molecules that physically experience the mechanical load; when force is applied to non-adhesion receptors that do not link to the cytoskeleton, stress dissipates locally at the cell surface, and the biochemical response is muted. In this schematic, new tubulin monomers add onto the end of a microtubule (yellow symbols) when tension is applied to integrins, and the microtubule is decompressed as a result of a change in the critical concentration of tubulin. The blue form indicates a molecule that is physically distorted by stress transferred from integrins to the cytoskeleton and, as a result, changes its kinetics (increases its rate constant for chemical conversion of substrate 1 into product 2). In this manner, both cytoskeletal structure (architecture) and prestress (tension) in the cytoskeleton may modulate the cellular response to mechanical stress. (Ingber, D. (2003c). J Cell Sci; Stamenovic, D. & Ingber, D. E. (2009). Soft Matter 5: 1137–1145.)

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Integrin-mediated force transfers occurs bidirectionally. In addition to sensing external mechanical forces, cells also exert cytoskeleton-generated contractile forces on these same integrin-ECM linkages ­(Ingber, 1997; Alenghat and Ingber, 2002; Ingber, 2006; Lele et  al., 2006; Parker and Ingber, 2007). Moreover, since ECM is relatively rigid compared to the cell, differences in the ability of the ECM to physically resist ­cell-generated traction forces directly modify cytoskeletal prestress in a manner that is consistent with the “cellular tensegrity” hypothesis (Ingber et  al., 1985; Ingber, 2006; Stamenovic and Ingber, 2009). In this model of cell structure, mechanical stability of the cell and cytoskeletal network results from inward-directed tensional forces borne by actin microfilaments and intermediate filaments that are balanced by microtubule struts and cell surface adhesions to ECM and neighboring cells (Stamenovic and Coughlin, 1999; Wang et  al., 2001; Ingber, 2003b) (Figure II.6.5.4B). In addition to ­stabilizing cell shape and structure, the cytoskeletal tension-mediated cell–ECM force balance regulates cell sensitivity to external mechanical and chemical cues, thereby exerting a significant impact on overall cell function (Ingber, 2003b, 2006; Ghosh et  al., 2008; Mammoto et  al., 2009). Thus, although mechanical forces applied to integrins initiate a local response at the site of focal adhesion (Chicurel et  al., 1998b; ­Geiger et  al., 2001), the cell integrates this intracellular transduction response with other external cues in a tensegrity-dependent manner to elicit a global behavioral response (Ingber, 2003a).

IMPLICATIONS FOR FUTURE MATERIALS DESIGN FOR IN SITU TISSUE ENGINEERING The availability of artificially-engineered skin and cartilage as commercial products serves as a testament to the significant progress that has been made in the field of tissue engineering over the past two decades. However, it must be noted that the success in artificially engineering these tissues is, to a large extent, due to the relative simplicity of their structure and composition, and the low vascularity of the cartilage matrix and skin epidermis. Yet, producing these “living” tissue constructs in a largescale and reproducible manner has proved challenging, primarily owing to high cost and stringent quality control measures. Thus, to achieve the ultimate goal of regenerating more complex tissues and whole organs, we will need to develop novel in situ engineering approaches that can address these limitations by harnessing ­endogenous tissue and stem cells. One way to accomplish this goal would be to design strategies for “targeted” delivery of biomaterials to specific injury or diseased sites in the body (Figure II.6.5.5). These injectable delivery systems could be used to target therapeutic agents (drugs, genes) or soluble factors

to these critical locations, where they can then recruit endogenous tissue cells or bone marrow-derived stem cells required for tissue repair and regeneration. Peptides that selectively home to the vasculature of specific organs have been identified (Rajotte et al., 1998), and biomaterials derivatized with such targeting moieties could be used for tissue-specific therapies. These injectable, homing biomaterials will need to be programmed such that, upon reaching their target site, they self-assemble into stable, higher-order structures that integrate into the host tissue, and provide the correct adhesive and morphogenetic cues essential for orderly tissue ingrowth and development. Various approaches that facilitate self-assembly of nanoscale materials into supramolecular structures have already been developed. For instance, ampiphilic peptides have been engineered that can self-assemble into a three-dimensional hydrogel in situ and promote local tissue repair (Zhang, 2003). ­Self-assembled, tensegritybased three-dimensional nanostructures have also been built using DNA that can potentially be used as templates for in situ cell and tissue engineering (Douglas et al., 2009). Regardless of the chemical strategies, it will be advantageous to tune the microscale chemical and mechanical properties of three-dimensional scaffolds in situ and in real-time, in order to dynamically regulate the cellular microenvironment, as seen during different stages of endogenous tissue development. This concept has recently been explored using photosensitive PEGDAbased hydrogels, where three-dimensional chemical and mechanical micropatterns were created in situ by exposing cell-seeded hydrogels to light in a spatiotemporally defined manner (Hahn et al., 2006; Kloxin et al., 2009). Such variations in local scaffold properties produce distinct microenvironmental niches that permit precise spatiotemporal control over MSC differentiation (Kloxin et al., 2009). It may be possible to use similar approaches to pattern three-dimensional scaffolds in a way that can simultaneously establish differentials in cell growth and apoptosis, as seen during tissue morphogenesis. Such “smart” biomaterials offer great promise as scaffolds that can induce tissue and organ regeneration through recapitulation of the complex morphogenetic and differentiation events that underlie endogenous developmental processes.

CONCLUSION Today, biomaterials are an indispensable tool in our efforts to repair and regenerate injured or diseased tissues and organs. They are no longer viewed as inert carriers for cell delivery; instead, biomaterials are now designed to actively interact with cells and promote tissue development, as well as restore its function. As cells adhere and spread on a substrate, they exert traction forces at their adhesion points that are balanced by the resistance (stiffness or elasticity) of the substrate. This

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FIGURE II.6.5.5  Biomaterial design strategies for in situ tissue engineering. The schematic diagram depicts an example of how smarter approaches that leverage targeting nanomaterials and endogenous stem cells might be used to treat a diseased organ; myocardial infarction is used as an example. The first challenge will be to identify moieties (e.g., peptides, aptamers) that bind to site-specific ligands with high affinity. Nanomaterials derivatized with these targeting moieties (top) can be injected intravenously (middle left) and delivered selectively to sites of tissue damage (bottom left). These materials will need to be programmed to self-assemble into a three-dimensional scaffold that integrates into the host tissue (bottom center), but only when they reach to their target site. Once this occurs, they will be designed to release potent soluble bioactive factors that can increase expansion of critical stem cell populations within bone marrow (middle right), mobilize them into the circulation and recruit them to the infarct site (bottom left). Either the newly formed scaffold materials will need to release different factors over time that stimulate proliferation and differentiation of recruited stem cells and endothelial progenitor cells, or other materials with these inducing properties may be injected and targeted to these same sites at later times. In this manner, formation of vascular networks that can provide continued supplies of oxygen and nutrients will be stimulated in parallel with development of functional tissue structures (e.g., muscle bundles, nerves, connective tissue) (bottom right). This type of in situ tissue engineering approach may potentially lead to development of more effective and more cost-effective therapeutics for tissue and organ regeneration. mechanical force balance drives cytoskeleton-dependent changes in cell shape, which can switch cells between different fates (e.g., growth versus differentiation) in both mature tissue cells and in various types of stem cells. Analysis of embryonic development reveals that micromechanical interactions between cells and the ECM are also critical for tissue morphogenesis, thus suggesting the existence of an overarching micromechanical control point for developmental control. However, the mechanical and structural features of developing tissues vary over space and time on the micrometer scale. Thus, biomaterials that allow precise spatiotemporal regulation of adhesive and micromechanical cues might have tremendous regenerative potential if they can better recapitulate the dynamic cellular microenvironment characteristic of endogenous developmental processes. The ultimate challenge, however, will be to promote tissue repair and regeneration in situ at the local site of tissue defect. This goal could be accomplished by developing injectable biomimetic nanomaterials that specifically home to injury or diseased sites, where they self-assemble into higher-order structures that provide

the correct developmental and morphogenetic cues required for orderly tissue renewal.

ACKNOWLEDGEMENTS This work was supported by grants from DoD and NIH. K.G., C.K.T., and D.E.I. are recipients of NIH U54 Interdisciplinary Research Training Grant, American Heart Association Scientist Development Award and DoD Breast Cancer Innovator Award, respectively. The authors thank Kristin Johnson for her assistance with the illustration.

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CHAPTER II.6.6  BIOREACTORS FOR TISSUE ­ENGINEERING Nina Tandon, Elisa Cimetta, Sarindr Bhumiratana, Amandine Godier-Furnemont, Robert Maidhof, and Gordana Vunjak-Novakovic Department of Biomedical Engineering, Columbia University, New York, USA

INTRODUCTION This chapter is a review of the principles of bioreactor design for tissue engineering. We describe the design and operation of tissue engineering bioreactors developed to direct the differentiation and functional assembly of cells

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cultured on three-dimensional biomaterial scaffolds. We first discuss the general design requirements for tissue engineering bioreactors, with a focus on mass transport considerations associated with environmental control, and biophysical signals necessary to modulate cell differentiation and the formation of engineered tissues. Next, we discuss the specifics of bioreactor design and operation using six examples of distinctly different tissue engineering systems: (1) cartilage tissue engineering with mechanical loading; (2) tissue engineering of anatomically shaped human bone; (3) cardiac tissue engineering with mechanical stretch; (4) cardiac tissue engineering with electrical stimulation and medium perfusion; (5) tissue engineering of heart valves with mechanical stimulation and perfusion; and (6) tissue engineering of blood vessels with pulsatile