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This protocol describes a cell/hydrogel molding method for precise and reproducible biomimetic fabrication of three-dimensional (3D) muscle tissue architectures in vitro. Using a high aspect ratio soft lithography technique, we fabricate polydimethylsiloxane (PDMS) molds containing arrays of mesoscopic posts with defined size, elongation and spacing. On cell/hydrogel molding, these posts serve to enhance the diffusion of nutrients to cells by introducing elliptical pores in the cell-laden hydrogels and to guide local 3D cell alignment by governing the spatial pattern of mechanical tension. Instead of ultraviolet or chemical cross-linking, this method utilizes natural hydrogel polymerization and topographically constrained cell-mediated gel compaction to create the desired 3D tissue structures. We apply this method to fabricate several square centimeter large, few hundred micron-thick bioartificial muscle tissues composed of viable, dense, uniformly aligned and highly differentiated cardiac or skeletal muscle fibers. The protocol takes 4–5 d to fabricate PDMS molds followed by 2 weeks of cell culture.
Tissue engineering technology classically combines living cells and biomaterials with the primary goal of creating functional tissue substitutes for the repair of diseased and damaged organs1. This technology is also utilized to generate biomimetic three-dimensional (3D) tissue culture systems for fundamental studies of cell–matrix interactions2, tissue morphogenesis3,4 and structure–function relationships5. A 3D tissue culture environment allows individual cells to assume a shape and exhibit matrix adhesions that are more in vivo-like than those formed in a two-dimensional (2D) environment2. Furthermore, compared with conventional 2D cultures in a large extracellular bath, dense 3D cultures with confined extracellular space can amplify local autocrine and paracrine actions of cells6 and, in addition, directly affect the function of electrically active tissues, such as nerve and muscle7,8.
The physiological and pathophysiological functions of various tissues in the body are critically dependent on the particular spatial arrangement of cells as well as the extracellular matrix in 3D (ref. 9). For example, native skeletal muscle contains long parallel muscle bundles composed of densely packed and highly aligned myofibers. This anisotropic tissue structure directly determines the contractile force and passive mechanical stiffness of the muscle10. Similarly, efficient pumping of the heart relies on the spatially and temporally coordinated electromechanical activity that is uniquely governed by the specialized intercellular connections11 and complex 3D alignment of cardiac fibers and sheets12,13. To be able to study these intricate 3D structure–function relationships in vitro or to restore native function in vivo, bioartificial muscle tissues need to contain a relatively large volume of densely packed, differentiated muscle fibers that can be locally aligned in desired directions using well-controlled and reproducible methodologies.
Aligning one or two muscle cell layers can be readily achieved by a number of techniques that use parallel microgrooves14,15, electrospun micro- and nano-fibers16,17, micropatterned protein lines18–20 or mechanical stretch16,21. In contrast, creating uniform 3D cell alignment of many muscle cell layers over a large area is considerably harder to achieve. One potential approach is to manipulate a polymer scaffold microstructure to provide topographical cues for orienting 3D cell growth. We, for example, utilized sucrose leaching techniques to fabricate poly(lactic-co-glycolic) acid scaffolds with oriented pores that guided the 3D alignment of cardiac cells over few square cm area5. Others used laser-drilled poly(glycerol sebacate) membranes22 or collagen matrices with oriented pores23 to achieve similar results. Despite the relatively simple fabrication processes, the use of anisotropic polymer scaffolds for the formation of thick and aligned 3D bioartificial muscle tissues is suboptimal because (1) the polymer phase represents an obstacle to dense and spatially continuous muscle cell growth, (2) no methods exist to precisely vary the local cell alignment and tissue thickness by controlling the polymer scaffold structure, (3) the rigidity of the polymer scaffold may prevent or dampen the macroscopic contractions of bioartificial muscle and (4) eventual biodegradation of the polymer may adversely affect the established cellular alignment.
In comparison with polymer scaffolds, naturally derived hydrogels (e.g., collagen, fibrin) have several properties that make them well-suited for the engineering of functional muscle tissues, including: (1) rapid polymerization that enables spatially uniform cell seeding, (2) significant compaction that yields high cell density, (3) abundant cell attachment sites that facilitate cell spreading and (4) high mechanical compliance that permits macroscopic tissue contractions and allows the application of tensile forces to align cells in 3D (refs. 24,25). Thus far, muscle cells in hydrogel systems have been aligned by constraining cell growth to one direction using a bundle or ring geometry24,26–29. These and similar methods for fabricating pseudo-1D muscle geometries are unsuitable for the generation of larger and more complex muscle tissue structures in which cell alignment needs to be varied in a controllable manner over different spatial scales.
However, approaches that utilize photolithographic patterning of hydrogels enable a relatively fast layer-by-layer assembly of cells into free-standing 3D structures with controllable geometry and size30. These methods mainly utilize synthetic31 or modified natural hydrogels32, both of which possess limited ability to support muscle cell spreading and growth. In addition, the use of photosensitive cross-linkers and ultraviolet radiation for hydrogel polymerization may adversely affect the viability, proliferation and differentiation of embedded cells33,34. As muscle cells significantly compact natural hydrogels and undergo spontaneous contractions in culture, a hydrogel-based methodology for the fabrication of bioartificial muscle tissues should provide robust mechanical support for cell growth and function, while allowing cell-mediated hydrogel remodeling to occur without a loss in 3D cell alignment.
Here we present details of a hydrogel molding method developed in our laboratory35 (Fig. 1a) to fabricate relatively large and thick muscle tissue constructs in which 3D tissue geometry and local muscle fiber alignments can be precisely varied. Specifically, a high aspect ratio (height/width of 5–10) soft lithography technique has been optimized to create polydimethylsiloxane (PDMS) tissue molds containing an array of elongated, mesoscopic-sized posts with typical dimensions of 1.2 mm × 0.2 mm × 1.5 mm (length × width × height). These posts are one to two orders of magnitude taller (up to 2.5 mm) than the photolithographic features typically used in microfluidic36 and other biological applications37. A mixture of muscle cells and fibrin gel is then cultured for two weeks in the PDMS molds to create a muscle sheet with elliptical pores formed around the posts. Through cell-mediated gel compaction and anchoring at the ends of the posts, a strain field is formed within the hydrogel that guides local 3D cell alignments along the pore boundaries. The pores serve to increase the diffusion of oxygen and nutrients to embedded cells, allowing the formation of 100- to 400-μm-thick, dense and viable muscle tissues. The PDMS molds are reusable, thereby allowing the reproducible fabrication of a large number of identical 3D tissue replicas.
As we have previously shown, this approach enables 3D alignment of muscle cells over a relatively large area (>1 cm2) and the design of complex tissue geometries, such as abrupt changes in muscle fiber orientation35. The natural enzymatic action of thrombin on fibrinogen38 during hydrogel polymerization has no adverse effects on embedded cells. The abundant cell adhesion sites present in fibrin gel facilitate the interaction between the cells and the extracellular matrix (ECM), and, subsequently, the structural and biochemical remodeling of the ECM by the cells. The facilitated ECM remodeling, along with the topographical and mechanical cues provided by the PDMS mold, guide the proper 3D assembly and integration of cells into a functional, aligned bioartificial muscle tissue. Long-term spontaneous contractions of differentiated muscle fibers are adequately supported by the relatively high compliance of fibrin gel39. In contrast, collagen gel-based tissues made using this method disintegrate during long-term culture due to vigorous tissue contractions35. In the case of noncontractile tissues, such as engineered tendon nets made by molding smooth muscle cells and collagen gel around 3 mm × 3 mm square posts, the resulting tissue structure can be stable for 6–8 weeks40.
Importantly, the computer-aided design of photomasks used to fabricate the template wafers enables precise variations in tissue mold parameters, such as overall mold dimensions, post height, length, width, orientation and spacing, which in turn determine the resulting tissue size, thickness, porosity, local and overall myofiber alignment as well as the dimensions of the formed muscle bundles. In addition, the resulting tissue properties are also determined in a highly complex manner using the characteristics of cell–gel mixture. Therefore, precise fabrication of a desired tissue structure currently requires that appropriate dimensions of the tissue mold are determined empirically. Accurate control of tissue structures based solely on the mold dimensions will necessitate the development of advanced computer models that robustly predict the process of cell-mediated gel compaction using realistic descriptions of cell proliferation, differentiation, cell–cell and cell–matrix interactions41.
As this method provides precise control of the engineered 3D tissue structure, it could be applied to construct planar in vitro analogues of various muscle architectures found in vivo. In particular, the computer-aided design of tissue molds using structural data from native muscle, as obtained using a variety of modern imaging modalities (e.g., diffusion tensor magnetic resonance imaging42, CT scanning43), would allow the design of more realistic muscle tissue structures. By incorporating different cell types and applying genetic and pharmacological manipulations, this methodology would allow systematic studies of the individual and combined roles of 3D structure, cellular composition and gene and protein expression in muscle development, physiology and pathology. Our ongoing study is focused on applying this method to create functional, differentiated skeletal and cardiac muscle tissues starting from stem cell-derived myogenic progenitor cells. Stacking multiple stem cell-derived muscle tissues in vitro31 or during implantation44 to form a thicker and stronger bioartificial muscle patch may eventually enable the application of this methodology to the burgeoning field of regenerative medicine.
A protocol based on the standard UV photolithography techniques has been developed to create well-defined, high aspect ratio photoresist features. Multiple layers of SU-8 photoresist are spin-coated on a master silicon wafer to a final thickness of up to 2.5 mm. After extensive soft-baking to remove residual solvent, the template is patterned by selective exposure to ultraviolet light through a transparency photomask (Fig. 1b). The resulting photoresist features typically have a width of no less than 0.2 mm and a height-to-width aspect ratio between 5:1 and 10:1 (Fig. 1c). Creating taller and higher aspect ratio features has not been possible using this protocol. The use of X-ray lithography45 is recommended if this is to be attempted.
PDMS tissue molds replicating the high aspect ratio photoresist features of the patterned master wafer are created using a double-casting method46. A negative replica PDMS template (Fig. 1d) is first molded off the master wafer and then silanized to make it nonadhesive to further application of PDMS. The silanized template is then used to mold a final set of PDMS structures (Fig. 1e) that can be directly utilized as tissue molds. One advantage of using the double-casting method is that one master wafer can be used to create several negative replica PDMS templates, each of which can be used to generate a large number of identical tissue molds, which in turn can be reused multiple times for the tissue culture. The relatively fragile master wafer is thus protected from being damaged by frequent use. As PDMS elastomer can be cast with sub-0.1-μm fidelity47, all PDMS tissue molds have dimensions virtually identical to those of the original master wafer.
PDMS tissue molds are rendered hydrophilic by oxygen plasma treatment and coated with 0.2% (wt/vol) solution of pluronic F-127 to prevent hydrogel adhesion. AVelcro frame is pinned within the tissue mold (Fig. 1f) to anchor the hydrogel. The frame provides mechanical tension during gel compaction as well as facilitates handling and transfer of the resulting tissue constructs. A mixture of fibrin gel, Matrigel and muscle cells is then polymerized within the PDMS molds at 37 °C (Fig. 1g). The hydrogel-containing PDMS molds are immersed in culture medium (Fig. 1h) and cultured in static conditions for 2 weeks to allow the formation of viable, dense and aligned bioartificial muscle tissues.
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For the initial cell seeding density of 5 × 106 NRVMs or 10 × 106 NRSKMs per ml of hydrogel, cell-mediated gel compaction reaches steady state within 8-13 d of culture (Fig. 3). After two weeks of culture, a 3D network of interconnected muscle tissue bundles is formed with the dimensions and spatial distribution that are directly controlled through the precise fabrication of the PDMS molds (Fig. 4). Specifically, this method allows for the independent control of the muscle construct thickness, overall porosity, bundle diameter, and global and local cell alignment in a reproducible manner, by controlling the thickness of the PDMS mold, the length-to-width ratio of the hexagonal post and the initial porosity before gel compaction35. To quantitatively characterize the overall degree and spatial distribution of cell alignment, we have constructed maps of local cell alignment angles within the tissue constructs (Fig. 5). The average deviation of all the alignment angles from the mean angle has been used as a quantitative measure of the degree of cell alignment. The presence of elongated pores in the obtained muscle constructs significantly increases the degree of overall cell alignment compared with the nonporous cell per hydrogel sheet made in molds of identical dimensions but without posts (Fig. 5b).
Immunohistological analysis reveals that the resulting 3D muscle bundles are composed of densely packed, evenly distributed and uniformly aligned muscle cells (Fig. 6). Specifically, aligned NRSKMs in tissue bundles (Fig. 6a,b) fuse into multinucleated myogenin positive myotubes, develop distinct cross-striations and spontaneously twitch after the onset of differentiation (see Supplementary Video 1). Similarly, aligned NRVMs in tissue bundles (Fig. 6c,d) show distinct cross-striations, interconnect through connexin-43 gap junctions, spontaneously contract in a synchronous manner (see Supplementary Video 2) and develop directional differences (anisotropy) in the conduction velocity of electrical propagation (data not shown).
We thank Ava Krol and Lisa Satterwhite for their assistance with cell isolation. This study is supported by a national science scholarship from the Singapore Agency for Science, Technology and Research (A*STAR) to B.L.; an American Heart Association predoctoral fellowship (No. 0715178U) to N.B. (Nima Badie); and NIH grants HL080469 from the National Heart, Lung, and Blood Institute and AR055226 from the National Institute of Arthritis and Musculoskeletal and Skin Diseases to N.B. (Nenad Bursac).
Author Contributions W.B. and B.L. contributed equally to this work. W.B., B.L. and N.B. (Nenad Bursac) jointly developed the protocol; B.L. optimized the fabrication of PDMS tissue molds; W.B. carried out the cell–hydrogel molding experiments and structural assessment of the resulting tissue constructs; N.B. (Nima Badie) developed the program for the analysis of cell alignment. All authors contributed the preparation of the paper.
Note: Supplementary information is available via the HTML version of this article.