The effects of mechanical forces induced by the surrounding environment (both soluble and insoluble factors) on cell behaviours are becoming an important topic in biological research. Cell adhesion to the ECM regulates cell proliferation and differentiation, and it is believed that tension has been shown to be one of the major mediators of both stimuli. However, most of our understanding of cell functions is based on the studies of cells cultured on stiff surfaces, such as glass coverslips and tissue culture dishes, which are often coated with a very thin layer of ECM. Such a thin ECM coating might not be relevant to the mechanical properties of the microenvironment for most in vivo tissues. A previous study has revealed that tissue stiffness runs from very stiff, such as Achilles' tendon (ca. 310 MPa), to very soft, such as mammary glands (ca. 160 Pa) . These tissue architectures serve as structural-based scaffolding and a source of inherent forces of mechanical stimulation for single cells. Cellular behaviours such as cell proliferation, differentiation and even apoptosis under stimulation by substrate stiffness are highly tuned . Aberrant regulation of in vivo tissue stiffness may result in severe and chronic pathological events, such as fibrosis and cancer [70-73]. Therefore, understanding cellular responses upon stimulation by mechanical inputs from the substratum or surrounding microenvironment may provide useful information for manipulating cellular behaviours. Several systems have been used to study the influence of substratum stiffness on cellular behaviours. A simple method that is typically used to change the stiffness of a substratum is protein-based ECM gel, such as collagen, fibrin and collagen mixed with fibrin, laminin and other ECM proteins [74-77]. Other materials such as polysaccharide-based alginate gel can also be manipulated to exhibit distinct compliance [78, 79]. By increasing the protein concentration, the stiffness of the gel can be increased. However, the major disadvantage of using natural gels is that changing the concentration of these natural polymers affects the mechanical stiffness and the ligand density, which may result in uncertain cellular responses upon cell plating on substrates of different stiffness levels. In addition to natural polymers, several groups have also developed synthetic polymers, such as PA and poly(ethylene glycol) (PEG) gels. These gels are chemically inert to cell adhesion unless the surface of the gel is pre-coated with ECM proteins, such as fibronectin or collagen. Thus, the stiffness of the gel can be manipulated by changing the cross-linking of the polymer without changing the material chemistry [43, 44]. Several studies have shown that matrix compliance does affect cellular functions. Fusion of myoblasts leads to the formation of polynuclear striated myotubes on collagen strips attached to glass or PA gels with various elasticities. Myotubes exhibit striations only on substrates of intermediate stiffness (ca. 8–10 kPa), but not on substrates of high (17 kPa) or low (ca. 1 kPa) stiffness (Fig. 2C and D) . Hepatocytes, as in the case of myotube formation, prefer slightly cross-linked Matrigel that is stiffer than basal Matrigel and can form aggregations and differentiation . Using such a tunable substrate system, it is demonstrated that elasticity of the matrix microenvironment can modulate MSC lineage commitment as well. hMSCs differentiate into neuronal-like cells on soft substrate that mimics the stiffness of brain tissues. On the substrate with intermediate stiffness similar to muscles, these cells differentiate into a myoblast lineage, while these cells plated on stiff substrate with a stiffness similar to bone differentiate into osteoblasts . In addition, matrix stiffness can modulate soluble factor-induced MSC differentiation. Park et al. have found that MSCs on a stiff substrate express smooth muscle cell (SMC) markers, such as α-actin and calponin, whereas MSCs express chondrogenic marker type II collagen and the adipogenic marker, lipoprotein lipase (LPL) on soft substrate. Treatment with transforming growth factor (TGF)-β increases SMC marker expression on stiff substrates, while TGF-β increases chondrogenic marker expression, but suppresses adipogenic marker expression on soft substrates . However, the major disadvantage of using synthetic gels is that changing the stiffness of the gel by modulating the cross-linker not only alters the mechanics of these gels but also the material properties, such as the surface porosity, geometry and ligand-binding properties. Therefore, microfabricated, micromolded elastomeric micropost arrays, which decouple the substrate stiffness from adhesive and surface properties to provide a wide range of substrate stiffness values, were reported by Fu et al. . These devices include the same micropost surface geometry, but differ in post-heights, which can generate substrate stiffness levels across a 1000-fold range. Such devices provide an ECM analogue with different stiffness levels to regulate stem cell commitment and also serve as a force detector to measure contractile forces preceding MSC differentiation at the single cell level . Together, these studies provide insights as to how substrate stiffness differentially regulates MSC lineage commitment and how mechanical stimulation cooperates with soluble factors to modulate MSC differentiation.
Figure 2. Mechanical stimulus–induced differentiation. (A) Cell shape drives mesenchymal stem cell (MSC) lineage commitment. Human (h)MSCs became bone only on large micropatterned islands, whereas adipogenesis occurred on small islands. (B) Quantitative results of MSC commitment on different-sized islands. Both A and B were reproduced from Ref. . (C) Myocytes cultured on collagen-coated polyacrylamide (PA) gels with various stiffness levels. Striated myotubes formed only on gels of intermediate stiffness. (D) Quantification results of optimal myotube formation on gels with different stiffness levels. Both C and D were reproduced from Ref. . (E) The elastic modulus of solid tissues. (F) The stiffness of the PA gel system can be modulated by changing the amount of the crosslinker. Cell adhesion to the gel can be controlled by covalent attachment of extracellular matrix (ECM) proteins (in this case, type 1 collagen). Human mesenchymal stem cells (hMSCs) seeded onto PA gels with different stiffness levels showed different morphologies. Cells were unspread with a branched morphology on soft substrate (0.1–1 kPa), had a bipolar morphology on intermediate stiffness (8–17 kPa) and had a polygonal morphology on stiff substrate (25–40 kPa) 96 hrs after seeding. (G) hMSCs differentiated into a neuronal lineage on soft substrate (0.1–1 kPa; as indicated by staining of βIII tubulin staining in cell branches); myogenic on intermediate stiffness (8–17 kPa; as indicated by MyoD staining of nuclei), and osteogenic on stiff substrate (as indicated by the punctuate CBFα1 staining of nuclei). E, F and G were reproduced from Ref.  (© 2004 Rockefeller University Press. Originally published in J. Cell Biol. 166:877–887).
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