Cell response to matrix mechanics: Focus on collagen

Biochimica et Biophysica Acta 1793 (2009) 893–902

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Biochimica et Biophysica Acta j o u r n a l h o m e p a g e : w w w. e l s e v i e r. c o m / l o c a t e / b b a m c r

Review

Cell response to matrix mechanics: Focus on collagen Anne L. Plant a,⁎, Kiran Bhadriraju b, Tighe A. Spurlin a, John T. Elliott a a b

National Institute of Standards and Technology, Biochemical Science Division, Gaithersburg, MD 20899, USA SAIC, Arlington, VA 22203, USA

a r t i c l e

i n f o

Article history: Received 11 January 2008 Accepted 27 October 2008 Available online 5 November 2008 Keywords: Extracellular matrix Collagen Type 1 collagen Collagen gel Collagen thin film Collagen fibril Vascular smooth muscle cell Integrin Mechanical cue Myosin light chain Cell spreading Actin cytoskeleton Proliferation

a b s t r a c t Many model systems and measurement tools have been engineered for observing and quantifying the effect of mechanics on cellular response. These have contributed greatly to our current knowledge of the molecular events by which mechanical cues affect cell biology. Cell responses to the mechanical properties of type 1 collagen gels are discussed, followed by a description of a model system of very thin, mechanically tunable collagen films that evoke similar responses from cells as do gel systems, but have additional advantages. Cell responses to thin films of collagen suggest that at least some of the mechanical cues that cells can respond to in their environment occur at the sub-micron scale. Mechanical properties of thin films of collagen can be tuned without altering integrin engagement, and in some cases without altering topology, making them useful in addressing questions regarding the roles of specific integrins in transducing or mitigating responses to mechanical cues. The temporal response of cells to differences in ECM may provide insight into mechanisms of signal transduction. © 2008 Published by Elsevier B.V.

1. Introduction: the extracellular matrix and mechanical cues Adhesion of cells to the extracellular matrix (ECM) has been shown to influence vital aspects of anchorage dependent cell behavior including survival, proliferation, and differentiation [1–6]. The ECM provides physical anchorage sites for cells in tissues, and also provides specific signaling cues to cells. The cues provided by the ECM are determined in part by the composition of the ECM and the identity of the cell-surface receptors, such as the integrins, that are engaged with ECM proteins. The composition of the ECM varies with the tissue type, developmental stage of cells, and with the state of health or disease of the tissue [7–9]. In addition, it has become increasingly apparent that the mechanical properties of the matrix can play a significant role in determining the cell program. Independent of receptor engagement per se, the cell's response to its mechanical environment has been shown to exert a strong influence on migration [10], gene expression [6,11–13], survival and proliferation [14], angiogenesis [15] and differentiation [16]. Change in the mechanical properties of the ECM is increasingly appreciated as a critical contributor to the etiology and progression of diseases such as vascular disease and cancer [17–22]. As evidence of the significance of ⁎ Corresponding author. Tel.: +301 975 3124; fax: +301 330 3447. E-mail address: [email protected] (A.L. Plant). 0167-4889/$ – see front matter © 2008 Published by Elsevier B.V. doi:10.1016/j.bbamcr.2008.10.012

matrix mechanics to cell physiology accumulates, an understanding of the mechanisms by which this parameter has such a marked effect on cell response continues to be sought. Cells ligate the ECM primarily through the family of integrin receptors, which consists of at least 24 different combinations of alpha and beta chains [23,24], and which determine specificity for ECM composition and initiate specific downstream signaling events. Integrins play key roles in development and differentiation [25–28], in homeostatic processes that are mediated by adhesion, migration and cell shape [29–32], and in assembly of extracellular matrix [33–35]. While integrin ligation with ECM is essential for signaling, it is not sufficient. Clustering of integrins in the plane of the membrane [36], and the development of mechanical tension [23,37], appear to be required for maturation of initial ligation complexes at the cytosolic side of the cell membrane, and subsequent intracellular signaling. Under mechanical tension, initial focal complexes formed in response to ECM ligation by integrins can mature into larger complexes such as focal adhesions, which can contain at least 50 different proteins and provide sites for anchoring bundles of actin microfilaments [37,38]. Functionally distinct types of receptor protein–matrix complexes have been reported in response to changes in physical properties of the matrix, independent of changes in the identity of integrins engaged [39]. The formation, size, composition, and intracellular location of the

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receptor-containing protein complexes that are focal complexes, focal adhesions and fibrillar adhesions appear to be time-dependent as well as dependent on the mechanical forces exerted at these attachment points [18,38,40–44]. Despite identification of many of the molecular details of how force is transduced at focal adhesions and how tension influences downstream signaling, understanding of the process is still incomplete. Data that show that cellular response to matrix stiffness is different for different cell types suggest that adhesion receptors may be responsible for determining that variability [45]. On the other hand, data on the response of smooth muscle cells to matrix stiffness through stress fibers and membrane tension have led to the suggestion that mechanics may override matrix ligand pathways in directing cell responses [46]. At the moment it is unclear to what extent the cell program is influenced by force-induced signaling in concert with signaling induced by other parameters such as receptor engagement. Quantitative assessment of these environmental factors on signaling pathways is a challenge that well-designed model systems can address. Many model systems and measurement tools have been engineered for observing and quantifying the effect of mechanics on cellular response, and have contributed greatly to our current knowledge of the role of mechanics in cell biology. Engineered synthetic polymeric and other materials can allow precise and systematic control over the mechanical properties of the cell substrate, and have provided quantitative information about the forces that are sensed and exerted by cells. Such systems include polymers of tunable compliance that have been used to determine that cells are sensitive to a range of matrix compliance with respect to cell migration and spreading [45,47], and differentiation [16]. The use of soft lithography to limit cell spreading by patterning adhesive proteins onto areas of defined size has provided observations of the cross talk between receptor engagement, cytoskeleton organization and cell survival and proliferation [4,48]. The application of micronsized polymeric devices such as flexible pads and posts [49,50], and the use of optical laser trapping [51] and a hydrodynamic spinning disc device [52,53] have allowed quantitative evaluation of the adhesive and contractile forces that cells can exert. Experiments that have explicitly applied tension to receptor–ligand complexes have provided insight into the molecular scale response. These responses include strengthening of ligand–receptor adhesion [42], the addition of proteins to focal complexes [39,51], and the activation of downstream signaling molecules [54]. In this article, we briefly describe several types of cell responses to the mechanical properties of bulk collagen gels, followed by a description of a model system of very thin, mechanically tunable collagen films that evoke similar responses from cells as do bulk gel systems, but have certain additional advantages. Finally, some cellular responses that we have observed in this model system will be presented. 2. Model systems based on collagen for studying the influence of mechanical cues Model systems comprised entirely of naturally occurring ECM proteins have the advantage of providing physiologically relevant receptor engagement, although they are generally not as amenable to tunability of mechanical stiffness as are synthetic materials. The predominant ECM-derived experimental systems engaged in the study of mechanical influences include studies with matrigel [55], fibrin [56,57], and collagens. Collagens are the most prevalent proteins in the body, and so the response of cells to the mechanical properties collagen has been of great interest. The most prevalent collagen, type 1, which is the focus of this review, consists of 1.5 nm diameter and 300 nm long triple-helical monomeric units. These monomers organize into supramolecular fibrillar structures that can vary in

size, mechanics and degree of supramolecular organization [58–61]. Type 1 collagen plays critical structural roles in the formation of cartilage and bone and provides the structural framework for most organs [61]. 2.1. Collagen gels Among the first model systems that demonstrated the significance of the mechanical features of the matrix environment to cell response used matrices of collagen fibrils to mimic wound healing [62]. Dense collagen gels (on the order of 2 mg/mL collagen) were either adhered to a culture dish or allowed to float [62–64]. Fibroblasts in the floating matrix, which provides relatively little resistance to cell traction, can contract the matrix in 3 dimensions. On the attached gels, as cells begin to contract, resistance against contraction apparently generates tension. In the first case, cells appear stellate in morphology and are quiescent; in the second case, cells resemble myofibroblasts and become oriented, bipolar, exhibit actin stress fibers and proliferate [64]. Differences in gene expression in fibroblasts in attached vs. floating gels include alpha 2 integrin subunit [65], matrix metalloproteinase 1 [66], tenascin-C [67], and type 1 collagen [68]. It was shown that the contraction of attached gels requires an intact cytoskeleton [62], and is dependent on engagement of alpha2 beta1 integrins [66,69,70]. Initial adhesive events result in activation of FAK [71]. Allowing cells to contract the matrix by floating the gels was observed to cause cells to be less responsive to serum or purified growth factors [72,73]. The observation that cell response to soluble factors such as growth factors can be determined by the degree of mechanical resistance provided by the extracellular matrix is likely the result of integrated signaling from mechanics and receptor engagement, which provides a rationale for the array of cell behaviors seen in tissues in vivo [15]. Adhered and floating collagen gels have also been used to elucidate the role of matrix compliance on morphogenesis and differentiation in the mammary gland [12]. Mammary epithelial cells will form appropriately organized structures on floating collagen gels but not on attached gels or stiff culture materials [74–76]. Cells cultured on floating gels, but not on cell culture dishes, produce caseins as expected from normal differentiated mammary tissue [6]. Type 1 collagen gels have also been employed extensively in the study of the etiology of vascular diseases. In hypertensive animals, arteries can become occluded by an ingrowth of smooth muscle cells into the lumen, accompanied by secretion of the extracellular matrix protein, tenascin C [77,78]. Vascular smooth muscle cells (vSMC) respond to fibrillar type 1 collagen gels in a manner similar to other cell types, characterized by engagement of alpha 1 beta 1 or alpha 2 beta 1 integrins [79–82], and poor spreading and proliferation [83–85]. Physiological proteolysis of collagen by matrix metalloproteinases (MMPs) has been implicated in the progression of vascular disease [78,86]. MMP activity can alter cellular response by altering integrin ligation, through the exposure of cryptic RGD binding sites in type 1 collagen, which are subsequently engaged by alpha v beta 3 integrins [87,88]. Heat denaturation of collagen also exposes RGD sites in collagen [88], and denatured collagen (gelatin) has been used to mimic the vascular disease state as well as for studying tumor induced angiogenisis [89] and for examining gene expression changes in cardiomyocytes [90]. The response of vSMC to denatured collagen is similar to that seen in the vascular disease state, namely cells proliferate and produce tenascin C [83]. The use of heat denatured collagen as a model system for vascular disease is somewhat ambiguous since in addition to exposing new integrin binding sites, the gelatin is a more rigid matrix than the fibrillar collagen gel. As will be described below, we have observed that simply increasing the stiffness of the collagen matrix, without altering the integrin recognition or binding of beta 1, can result in increased cell spreading, proliferation, and production of tenascin C [91,92].

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2.2. Thin films of type 1 collagen fibrils Gels of fibrillar collagen represent an important model system, but collagen gels can have disadvantages. They are fragile and can be difficult to handle; labeling agents such as fluorescent reagents can get trapped within the gel and contribute undesirable background signal; and it can be difficult to characterize gels with respect to physical properties, reproducibility, and homogeneity. In addition, collagen gels, which are typically at least several millimeters in thickness, can be a significant source of scattered light, making visualization of cells difficult. We have presented and characterized thin films of fibrillar collagen as an alternative model system that has many of the characteristics of collagen gels, but is robust and reproducible, enabling replicate experiments with a high degree of confidence. These films can be studied by atomic force microscopy (AFM), ellipsometry and other methods to quantify structural details and reproducibility [91,93–95]. AFM indicates that fibrils can extend to about 600 nm above the surface, and ellipsometry indicates that these fibrils are loosely packed [94]. Because they scatter less light than thick gels, visualization of cells by optical microscopy is greatly improved. This simplifies the use of automated microscopy, allowing examination of large numbers of cells to provide quantitative information on the variations in response across the cell population [96]. Robust protocols for fabrication of collagen films [94] and for staining cells for automated determination of cell spread area [97] on collagen films have been published. Thin fibrillar films of type 1 collagen are prepared by exposing a solution of collagen monomer, at neutral pH and appropriate ionic strength, to a hydrophobic surface, allowing adsorption of small collagen units, and subsequent addition and self-assembly of large collagen fibrils for several hours at 37 °C. The hydrophobic surface most often used up to now has been an alkanethiol monolayer on a glass slide coated with a semitransparent layer of gold [92–94]. Recently we have evaluated thin fibrillar films of collagen formed at untreated polystyrene and determined them to be statistically identical with respect to the resulting cell response (submitted for publication); these thin collagen films can be integrated into many protocols employing collagen gels.

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Atomic force microscopy (AFM) images of fibrillar collagen thin films are shown in Fig. 1, and provide insight into the topographical and mechanical environment that is available to cells. Collagen fibrils of approximately 200 nm in diameter appear to be randomly and loosely packed, not unlike their appearance in bulk gels [93]. The fibrils are very flexible, as indicated by the poor lateral resolution of topographical features in the height mode image taken in aqueous solution in Fig. 1A, due to small scale fibril displacement during scanning. Greater lateral resolution can be observed in AFM images collected in phase mode where the phase lag in the cantilever oscillation is monitored as it scans the sample (Fig. 1B). The difference in phase lag between the oscillations driving the fixed end of the cantilever and oscillations of its free end is very sensitive to subtle changes in adhesion and viscoelasticity, and provides qualitative information about material properties [98–100]. This can be seen in a graph (Fig. 1C) of the phase lag across the features underlying the white line in Fig. 1B. The greater phase lag associated with the visible fibrils compared to the areas between the fibrils suggests that the large fibrils are more compliant than the surrounding regions [101]. This conclusion is supported by measurements of contact stiffness from individual force curves [91] that also indicated that large fibrils are significantly more compliant than regions devoid of large fibrils. Higher resolution AFM imaging (Fig. 1D, E) of films formed at low concentrations of collagen allows examination of the features of the collagen that resides very close to the support. Images in Fig. 1C show the presence of very small (10–20 nm diameter) features [94,102] that likely correspond to the precursor units that appear early in fibril growth [97]. These small aggregates appear to fully cover the surface of the support and large fibrils appear to grow out of them [94]. They may provide nucleation sites at which large fibrils grow by addition of monomer from solution. Collagen fibril growth is thought to occur by accretion as well as by lateral and end-to-end fusion of collagen fibrils in the ECM [103,104]. It is significant that these very small structures appear to have nascent fibrillar structure, because it suggests that the collagen is in its native, i.e., non-denatured, conformation [94]. These data are the basis for the model of collagen films shown in the schematic in the last panel of Fig. 1F.

Fig. 1. Atomic force microscopy of collagen type 1 films. (A) Height (tapping mode) image and (B) phase image of collagen film under PBS; brighter features in the phase image indicate areas that are less stiff. (C) Quantification of the phase response under the line in B. (D, E) Higher resolution AFM images taken in air of samples prepared at lower densities of large fibrils, which allows the underlying bed of smaller fibrils to be sampled. These images provide evidence for nucleation of large collagen fibrils from very small collagen assemblies at the surface, and (E) for branching of collagen fibrils within the collagen network. (F) A schematic of the fibrillar film based on all the experimental data.

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2.3. Cell response to thin films of type 1 collagen fibrils Comparative studies using A10 vSMC and NIH 3T3 fibroblasts have shown that cells appear to respond identically to thin films of fibrillar collagen as they do to thick gels with respect to cell spread area, cytoskeleton organization, tenascin-C expression, and proliferation rate [93,94]. Fig. 2A shows images of A10 vascular smooth muscle cells on native fibrillar collagen in the form of collagen gels or collagen films. Anti-beta 1 blocking experiments [93] have shown that cells interact with fibrillar collagen thin films through the beta 1 integrin receptor, which is the integrin ligated on non-denatured type 1 collagen gels. For comparison, A10 cells on gels or films of denatured collagen are also shown (Fig. 2B). Cells ligate denatured collagen through alph v beta 3 integrin [83,88] and this binding can be blocked with RGD peptide [93]. Compared to cell response on either gels or thin films of native fibrillar collagen, cells on denatured collagen assume a more spread phenotype, develop actin stress fibers, and secrete tenascin C (Fig. 2). We found that there is no significant difference in cell spreading on thin films compared to cell spreading on collagen gels (Fig. 2). Together, these data suggest that cell interaction with fibrillar collagen films is very similar to their interaction with collagen gels, and furthermore, that fibrillar collagen films can influence cell response in a manner that is highly similar, if not identical to collagen gels.

Because the fibrillar collagen films are very thin, the apparent recapitulation of the response seen on gels suggests something important about how cells respond to the mechanical environment. It suggests that, at least in part, cells interrogate and extract mechanical cues from their environment on a very local (submicron) scale. Because of the optical clarity they allow, the movement of collagen fibrils by cells on thin fibrillar films can be readily observed by light microscopy. Fig. 3A shows sequential images taken minutes apart and indicate the position of a fibril that is clearly moved by the cell. Cells have been observed to pull fibrils up and bury themselves underneath them as shown in Z-axis sections in Fig. 3B. The adsorption of small collagen aggregates to the hydrophobic surface and the subsequent self-assembly of fibrils to form fibrillar films of collagen apparently results in a stable, yet flexible matrix. 2.4. Tuning the mechanical properties of collagen films Thin films of collagen provide an experimental system for systematically modifying matrix mechanical properties for studying the effects of mechanics independently from integrin receptor ligation. There are a number of ways that mechanical properties of collagen can be altered in the absence of changes in integrin recognition.

Fig. 2. Vascular smooth muscle (A10) cells on (A) native fibrillar type 1 collagen gels or films as indicated, under phase contrast, stained for f-actin with phalloidin, immunostained to detect the expression of tenascin-C, as indicated. Quantification of the range of areas to which cells spread on the two substrates is shown in the panel on right; solid line is for cells on collagen films, dotted line is for cells on collagen gels. (B) Comparison the same cells on films or gels of denatured collagen. Treatment is as indicated. For cell area quantification, solid line is for cells on denatured collagen films, dotted line is for cells on denatured collagen gels.

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Fig. 3. Mobility of collagen fibrils and effect of dehydration of fibrils on cell response. Differential Interference Contrast micrographs of A10 vSMC on fully hydrated collagen thin films showing manipulation of fibrils by cells. (A) Sequential images showing fibril movement due to cell action. The arrow, which remains in the same location in all images, is included for reference. The interval between frames is 6 min. Scale = 10 μm. (B) A series of images taken in the Z-axis showing a cell after 48 h on a fully hydrated collagen thin film showing that fibrils have accumulated on the dorsal surface of the cell. Scale = 20 μm.

2.4.1. Altering the density of large collagen fibrils One way to modify mechanical signaling while still engaging alpha 1 beta 1 or alpha 2 beta 1 integrins is to vary the density of collagen fibrils. The density of large fibrils can be tuned by altering the

concentration of collagen in the solution during film formation as shown by AFM images of collagen films in Fig. 4. At low solution concentrations (ca. 25 μg/mL), the film is almost devoid of large fibrils and instead the surface is dominated by very small collagen

Fig. 4. A10 vSMC spreading and proliferation in response to mechanical properties of collagen. (A) Cell spread area was determined by segmenting cells stained with Texas Red. At least 100 cells were counted for each of 3 replicates. The white bar indicates the mean area of cells on collagen gels; the black bar indicates the mean area of cells on denatured collagen. The gray bars indicate mean areas of cell on thin films of fully hydrated fibrillar collagen at the concentrations of solution collagen indicated. A complete fibrillar film was formed at 750 and 250 μg/mL collagen; few to no fibrils were visible when collagen concentration was 25 μg/mL. The striped bar indicates mean area of cells on collagen films formed at 250 μg/mL and allowed to dehydrate. (B) Cells were plated on all surfaces at 400 cells/cm2. At the times indicated, replicate samples were removed from the incubator and cells were trypsinized and counted. Open square = collagen gel; closed square = collagen thin film; open diamond = dehydrated collagen film; closed diamond = denatured collagen. (C) AFM images of collagen films formed at the various concentrations of solution collagen; scale bar = 5 μm. (D) Phase contrast images of fibrillar collagen films that are fully hydrated or dehydrated. The films appear identical, but cells spread more on the dehydrated films.

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structures, which cells also ligate through beta 1 integrin [93]. Cells respond to this change in collagen fibril density by increased spreading as the density of the flexible fibrils is decreased. When thin films are prepared with a relatively high fibril density of fibrils (by preparing from collagen solutions at concentrations of 750 and 250 μg/mL), cells spread poorly. Under these conditions, cell spreading is similar to their spreading on collagen gels. At the lowest density of large collagen fibrils (formed from solutions of collagen at concentrations of 25 μg/mL), cell spreading was significantly greater (Fig. 4A). In fact, in the absence of large collagen fibrils, these cells spread to an area equivalent to the area they spread on denatured collagen, where their engagement with the matrix is through the alpha v beta 3 integrin. A number of earlier studies have also concluded that cells respond differently to collagen assembled into large fibrils compared to monomeric or unpolymerized collagen even when cells ligate both forms of collagen through alpha 1 beta 1 and alpha 2 beta 1 integrins [13,105,106]. Many protocols have been reported in the literature for preparing collagen coatings, including allowing collagen to absorb from low concentration or acidic solutions onto glass or tissue culture plastic or other synthetic polymeric materials. In most published reports, the structure of collagen was not examined, and is variously reported as monomeric, non-polymerized, or denatured, which may not be entirely accurate. (A recent systematic study has shown that collagen does not form fibrils at hydrophilic supports even at high concentrations [95]). It has been shown that primary vSMC [85,107] on fibrillar collagen gels, but not on collagen adsorbed onto dishes from an acid solution are arrested in the G1 phase of the cell cycle, and exhibit a decrease in a number of activities including secretion of ECM proteins like tenascin C, production of alpha actinin, levels of protein synthesis associated proteins, and a number of signaling molecules. Melanoma cells are also blocked from cell cycle progression when cells are on, or in, collagen gels [106] compared to collagen adsorbed to tissue culture dishes from acid solution. Fibroblasts on fibrillar collagen gels show reduced spreading and actin cytoskeleton organization compared to cells on monomeric collagen, even though cells mediate interaction with both collagen substrates through alpha 1 beta1 and alpha 2 beta1 integrin [105]. Primary hepatocytes show similar responses [13], although some differences in the effectiveness of anti-alpha 1 antibody to block integrin-dependent adhesion on polymerized gels vs. nonpolymerized collagen were reported. Some of this difference could potentially be due to the activity of other collagen receptors including the discoidin domain receptor (DDR2), which binds fibrillar collagen but not nonfibrillar collagen [108]. 2.4.2. Stiffening collagen fibrils by dehydration The distinctly different responses from cells to large fibrils of type 1 collagen compared to collagen that is in an unpolymerized form suggest that the fibrils play an important role in influencing the cell program. One can infer that mechanical cues are involved. However, because the changes in topology that accompany the elimination of large fibrils could potentially affect the response of cells, we examined a method for altering fibril mechanical properties without grossly changing topology or chemistry. Water plays an important role in the structure and mechanics of collagen. Water bridges occur within and between the collagen chains [109–112], and hydration layers maintain the spacing within collagen fibrils [113]. Rheological studies suggest that loss of water from collagen fibrils results in tighter packing of fibrils and enhanced mechanical rigidity due to increased inter-fibril attractive forces [114]. By allowing fibrillar films of collagen to dry in air for several hours, an apparently permanent change in the stiffness of the fibrils can be effected. Quantitative AFM and nanoindentation measurements were used to determine that the mechanical properties of dehydrated collagen fibrils were distinctly different from fibrils that were allowed to remain fully hydrated, and that the change in the contact stiffness of

collagen fibrils after dehydrating was an order of magnitude greater than that for fully hydrated fibrils [91]. Fig. 4A shows that cell spreading also increases in response to dehydrated collagen films, presumably due to the increased stiffness of the fibrils. Mean cell area is shown for vSMC on fibrillar collagen allowed to dehydrate for different lengths of time. After drying of fibrillar collagen films for 24 h, cells spread as much as they spread on nonfibrillar collagen formed at low concentration, or on denatured collagen. Interestingly, the loss of matrix flexibility, either due to dehydration or to the absence of fibrils, appears to provide a quantitatively similar phenotype with respect to cell spread area. Furthermore, under these conditions where beta 1 integrins are engaged, the extent of spreading is identical as for cells on denatured collagen where alpha v beta 3 integrins are ligated. These results suggest that under these conditions, cell spreading is apparently not determined by the nature of the integrin ligated as much as by the mechanics of the matrix. Many observations have led to the suggestion that cell spreading is required for proliferation, starting with the seminal study of and Folkman and Moscona [115]. Cells can be constrained from spreading using a variety of methods such as decreasing matrix ECM density [2], changing matrix composition [116], or plating cells on small spatial patterns of matrix [4]; these treatments all lead to an arrest in the cell cycle. Indeed Fig. 4B shows that A10 vSMC that spread more proliferate more than cells that don't spread well. Cells plated on the more rigid dehydrated collagen fibrils or on denatured collagen spread and proliferate, and cells that are poorly spread on flexible collagen fibrils do not proliferate. The data in Fig. 4B show that cells on denatured collagen approximately triple in number over 96 h in culture in the presence of 2% serum [93]. Interestingly, cells on dehydrated collagen fibrils spread as well as they do on denatured collagen, but appear to have a lower rate of proliferation. Whether this observed difference in proliferation rate is quantitatively influenced by the identity of the integrin ligated is yet to be determined. 2.4.3. Enzymatic crosslinking of collagen fibrils It is also possible to increase the mechanical stiffness of polymerized collagen thin films by covalent crosslinking, providing yet another way to tune the mechanical properties of these substrates. Physiologically, collagen fibril flexibility is influenced by chemical crosslinking catalyzed by lysyl oxidase and transglutaminase (Tgase), among other mechanisms [117–119]. At least seven forms of Tgase have been identified and shown to play important roles in physiologic process, such as wound healing, keratinocyte differentiation, and tumor biology [119]. The physiological role of type 2 Tgase is not fully understood, but it is present in a variety of tissues and is implicated in a number of human pathologies including tumor progression, cartilage calcification, and tissue fibrosis [119–121]. We have used treatment with transglutaminase as a means to tune collagen fibril flexibility. Transglutaminase treatment of collagen fibrils results in increased spreading by vSMC. Initial results (Fig. 5) show that increasing the time allowed for the enzyme to cross-link collagen fibrils results in an increase in cell spreading. Cell spreading increases from an area of 2655 ± 209 μm2 for cells on uncross-linked collagen films to 5417 ± 132 μm2 for cells on collagen films that had been cross-linked with 2.5 × 10− 5 U type 2 transglutaminase for 48 h. This observed increased spreading is likely the response of the vSMCs to the increased stiffness of crosslinked collagen fibrils [122]. 3. Exploring the mechanisms of mechanical signaling The forces involved in the interaction of cells with extracellular matrix are transduced through focal complexes, focal adhesions and fibrillar adhesions, which consist of integrins and a large array of

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Fig. 5. Effect of transglutaminase treatment of type 1 collagen fibrils on A10 vSMC spreading after 24 h in culture on (A) thin film of collagen fibrils; (B) thin film of collagen fibrils crosslinked with 2.5 × 10− 5 U transglutaminase type 2 for 48 h. Scale bar = 140 μm. (C) Mean cell areas determined after staining with Texas Red for accurate segmentation of cell edges. Cells areas were determined by quantification of at least 400 cells per replicate per treatment. SD bars are determined from 3 replicate experiments.

enzymatic and scaffolding proteins [123,124]. The time required for formation of initial focal complexes and the appearance of focal adhesions appears to be distinct, suggesting that the one structure

might mature to become the other [44]. Focal complexes resulting from the initial adhesion of cell receptor to matrix are reported to form within 60–90 s, while maturation by increase in size and

Fig. 6. Cell response to matrix stiffness and identity. (A) Time course for cell spreading on fully hydrated (flexible), dehydrated (stiff) fibrillar collagen thin films, and adsorbed fibronectin. At the times indicated, cells in replicate dishes were fixed, stained with Texas Red and their spread areas determined by automated microscopy. (B) Comparison of MLC phosphorylation on fully hydrated and dehydrated fibrillar collagen at 0.5 h and 3 h after plating. Cells were lysed at the corresponding time points and myosin light chain diphosphorylation, normalized to the amount of myosin heavy chain, was determined by western blotting. (C–F) Comparison of paxillin staining on fully hydrated (C, D) and dehydrated (E, F) fibrillar collagen at 1 h (C, E) and 24 h (D, F) after plating. Alexa-488 staining for paxillin is shown in green; Texas Red phalloidin staining for f-actin is in red; DAPI staining for nuclei is in blue. White arrow points to focal adhesions.

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recruitment of protein components in response to mechanical forces results in focal adhesions, which are reported to take on the order of 60 min to develop [44]. Myosin-driven contraction of actin filaments is likely the intracellular source of the mechanical tension with the extracellular matrix [51,125]. Myosin generated forces are associated with the movement of collagen fibrils by cells [126] and are required for the formation of focal adhesions [37,43]. The formation of vinculin-containing complexes has been observed to be reduced by the presence of an inhibitor of myosin light chain (MLC) kinase [51]. Contractility is regulated through myosin activation by phosphorylation involving the calcium dependent enzyme myosin light chain kinase (MLCK) and the calcium independent enzyme Rho dependent kinase (ROCK), which phosphorylate the myosin light chain (MLC) [127–129]. Quantitative cell by cell analysis has shown that cell spreading and extent of MLC phosphorylation are strongly correlated in individual cells plated on adsorbed fibronectin [130]. The ability to manipulate mechanical properties of the matrix in the absence of changes in topology or ligand density provides an opportunity for studying the effects of mechanics and the effects of receptor engagement independently, and to explore the possibility of integrin-specific differences in response to the mechanical environment. We have begun to explore how and when cells respond to their mechanical environment by probing cell spreading, myosin activation and focal adhesion formation in the temporal domain. In Fig. 6A we have compared A10 vSMC on fibrillar collagen films that are either fully hydrated or dehydrated, and followed changes during cell spreading compared to cell spreading on adsorbed fibronectin. At very early times of attachment and spreading (0.5 h and 1 h after seeding), cells on either fully hydrated flexible collagen fibrils or stiffer dehydrated collagen fibrils appear fairly isotropic and circular in appearance as they adhere to the collagen matrix, and their spread areas are the same. This early phase in which cells appear to be indifferent to the differences in matrix mechanical properties appears to last at least up to 1 h. After that time, cells on fully hydrated fibrils begin to diverge from a symmetrical shape (data not shown), and begin to assume a larger spread area, which reaches a maximum at about 6 h. Western blotting (Fig. 6B) shows that cells on the stiffer dehydrated collagen exhibit significantly more MLC phosphorylation after 3 h than cells on the flexible hydrated fibrils; however at early times in the spreading process, 0.5 h, there is no significant difference in MLC phosphorylation in cells on the two matrices. At about 1 h into the spreading process, paxillin staining in cells on either the hydrated (Fig. 6C) or dehydrated (Fig. 6E) collagen matrix also appears to be very similar, with cells showing little evidence of large focal adhesions. Spreading and activation of myosin on stiffened dehydrated collagen fibrils over time appear to be accompanied by maturation of focal adhesions. The appearance of paxillin staining in fully spread cells on dehydrated collagen fibrils (Fig. 6F) is quite different from cells on fully hydrated fibrils (Fig. 6D) after 24 h, suggesting that focal adhesion formation occurs only on the dehydrated substrate. While the temporal coincidence of cell spreading, MLC phosphorylation and focal adhesion formation are suggestive of causal relationships, these relationships bear further investigation. It is significant to note in Fig. 6A that the spreading of cells on adsorbed fibronectin is much more rapid than spreading of cells on either collagen matrix, and does not demonstrate the distinct lag phase in spreading that is observed by cells on either the flexible or stiffened collagen fibrils. These differences may reflect integrindependent responses, since fibronectin is primarily ligated through alpha 5 beta 1 and collagen through alpha 1 beta 1 and alpha 2 beta 1; alternatively, this observation may reflect differences in the mechanical stiffness of all three substrates. These possibilities will be explored in future experiments.

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