Collagen gels find widespread application as three-dimensional substrates in cell culture assays,1 drug delivery,2 and tissue engineering.3 At the macroscale, boundary constraints influence cell-laden collagen gel anisotropy;4 at shorter length scales, composites of collagen gels with microfabricated materials5–7 raise questions concerning how fibrillogenesis itself may be influenced by the geometry of such microstructures. In particular, collagen gel morphology imparted by compartmentalization within microfabricated materials could impact functional performance parameters (e.g., cell mobility, shape, or alignment;8–14 drug diffusion;2 hierarchical engineered tissue mechanics15–17) of such composite devices.
In biomimetic tissue engineering,18 collagen gels have been used for generating functional myocardium from heart cells.3 Collagen gels are capable of promoting cell alignment under boundary constraint4, 10 or cyclic loading,19 however they tend to be mechanically inferior to myocardium.8 We generated tissue engineered myocardium by cultivating heart cells on an accordion-like honeycomb (ALH) scaffold rendered by laser microablation of poly(glycerol sebacate) (PGS).20 The ALH scaffold provided cardio-mimetic anisotropic elastic properties and a capacity to guide preferential cell alignment. Toward enhancing heart cell-mediated contractility, we developed periodic finite element simulations for investigating changes in ALH scaffold geometry21 and investigated improving heart cell seeding efficiency via Matrigel.22 Matrigel, however, did not promote cardiomyocyte elongation.
Based on observations of directional collagen fibrillogenesis in collagen-doped microfluidic devices23 and that elongated scaffold pores can promote cell-secreted collagen alignment,9 we speculated that directional collagen fibrillogenesis might manifest within ALH pores.18 Anisotropic collagen fibrillogenesis could potentially overcome the limited heart cell elongation observed in Matrigel-ALH composites.22
To elucidate if ALH scaffolds can induce anisotropic collagen fibrillogenesis, the three-dimensional fibril organizations within ALH (Figure 1A) and square diamond (Figure 1B) pores were imaged by confocal reflectance microscopy (Figure 1C,D) and compared to collagen gelled unconstrained on glass slides (Figure S1, Supporting Information). ALH scaffolds preferentially oriented fibrils along the long axis of the pore (Figure 1C,E), with an orientation index OIALH = 17.75 ± 6.55% significantly higher than that measured for glass slides (OIglass = 1.45 ± 0.40%; p < 0.05). Indeed, neither glass slides nor square diamond scaffolds (Figure 1D,E; OIsquare = 0.26 ± 0.16%) induced preferential fibril orientation along the pore long axis. Providing a measure of both the density and homogeneity of the collagen gel, the inter-fibril distance distributions were quantified (Figure 1F). The mean interfibril distance was 5.06 ± 0.05 μm for collagen gelled on glass slides; values for the ALH (4.075 ± 0.5 μm) and square diamond (4.08 ± 0.3 μm) scaffolds were lower (p < 0.05). Hence, the organization of fibrils tended to be denser upon compartmentalization within the pores of ALH and square diamond scaffolds than on glass slides. The entropy value calculated for ALH pores (ϵALH = 3.8 ± 0.23) was significantly lower than that calculated for glass slides (ϵglass = 4.6 ± 0.10; p < 0.05), demonstrating that the organization of fibrils was more ordered in ALH pores. Square diamond pores exhibited an intermediate value ϵsquare = 4.18 ± 0.12.
To predict the mechanical stiffnesses and anisotropy of ALH scaffold-collagen composites, compare these with native heart muscle, and to investigate the ramifications of anisotropic collagen fibrillogenesis in the ALH scaffold, periodic FE simulations21 were conducted. FE predicted effective stiffnesses EPD and EXD, and anisotropy ratio r = EPD/EXD, and Voigt (Equation 4) and Reuss (Equation 5) elastic bounds of the composite initially assumed the collagen was isotropic using upper (24.3 kPa15) and lower (5 kPa24) bounds of collagen stiffness. The Reuss bound (range 7.3–35.1 kPa) was dictated by the most compliant component of the composite (i.e., the collagen); the Voigt bound (aka the “rule-of-mixtures”; range 265–278 kPa) was dictated by the stiffer component, and therefore varied only slightly with the stiffness of the collagen. By contrast, FE predicted effective stiffnesses EPD (range 108–215 kPa) and EXD(range 62.2–173 kPa) depended strongly on the stiffness of the collagen and were comparable to values measured by uniaxial tensile testing (Figure S3, Supporting Information). As expected, FE predicted and measured values of EPD and EXD with 3 mg mL−1 collagen gelled within the pores were higher than those reported for the ALH scaffold itself (EPD = 83 kPa and EXD = 31 kPa20). The FE predicted anisotropy ratios r = 1.2 and r = 1.7 associated with the upper and lower collagen stiffness bounds, respectively, were significantly lower than that predicted for the ALH scaffold without collagen (2.5)21. Recognizing that the collagen within the ALH scaffold was itself anisotropic (Figure 1), we simulated collagen anisotropy via an orthotropic material model targeting left ventricular stiffnesses in the circumferential (157 kPa) and longitudinal (84 kPa) directions (r = 1.87)20 and solved for the requisite collagen gel stiffnesses in the PD and XD directions. FE simulations predicted collagen gel stiffnesses = 47.0 kPa and = 26.5 kPa (rcoll = 1.77). Finally, comparing the spatial distribution of equivalent von Mises strain within the collagen matrix for isotropic (Figure 2A) and orthotropic (Figure 2B) assumptions, orthotropic yielded a more homogeneous strain distribution (range 0.04–0.05) along the central PD axis versus isotropic collagen gel (range 0.03–0.07). Hence, FE simulations can be used to predict collagen gel properties required to match ALH scaffold-collagen composite stiffnesses and anisotropy to native heart muscle.
Composite devices comprised of microfabricated materials and collagen gels offer the prospect of controlled bridging between the micro-to-nanometer length scales, potentially yielding novel in vitro cell culture assays,1 drug delivery systems,2 and engineered tissues.3 Hierarchically, engineered tissues formed by seeding cells onto microfabricated scaffolds evolve through the structural–mechanical interplay between scaffold, cells, and extracellular matrix. Without accounting for extracellular matrix, we demonstrated that heart cell-seeded ALH scaffolds could mimic aspects of cardiac anisotropy.20 Extracellular collagen structures, however, play important roles in myocardium.18 We demonstrated three key findings regarding collagen gelled within the ALH pore versus unconstrained on a glass slide: 1) increased order of the fibril distribution (i.e., decreased entropy; ϵALH = 3.8 ± 0.23 versus ϵglass = 4.6 ± 0.10; p < 0.05), 2) increased fibril density (i.e., decreased mean inter-fibril distance; dALH = 4.075 ± 0.5 μm versus dglass = 5.06 ± 0.05 μm; p < 0.05), and 3) increased fibril alignment along the reference angle defined by the ALH pore long axis (OIALH = 17.75 ± 6.55% versus OIglass = 1.45 ± 0.40%; p < 0.05). For comparison, Bayan et al. reported entropy values ranging from 6.37–6.5 and OI values ranging from 9.45–13.46% in similar acellular collagen gels.25 Of note, Bayan et al. did not detect significant differences in OI when comparing 1, 2, and 3 mg mL−1 collagen gels. Further, when gelled on a glass slide and compared with the 3 mg mL−1 gel, we did not detect any difference in the collagen orientation distribution in a 6 mg mL−1 collagen gel (Figure S4, Supporting Information). Of note, the degree of collagen fibril alignment mediated by the ALH pore geometry alone (OI = 17.75 ± 6.55%) was less than that observed by Bayan et al. in a 3 mg mL−1 cell laden gel cultivated for 12 days (OI = 30.86 ± 14.76%).25
A combination of mechanisms may have contributed to the anisotropic collagen fibrillogenesis observed herein. For example, when 3 mg mL−1 collagen solution was flowed into and gelled within the channels of a collagen-doped alginate microfluidic device, Gillette et al. observed that a number of collagen fibrils appeared to bridge contiguously, in straight lines, from the collagen-doped alginate (i.e., the channel walls) into the collagen gelled within the channel.23 Coupled with the preference for collagen fibril tip growth predicted by diffusion limited aggregation models by Parkinson et al.,26 the results from Gillette et al. suggest that collagen fibrils can grow in a straight line from the tips of collagen fibrils exposed at a surface into the bulk of a collagen solution. In the present study collagen solution was gelled in direct contact with the PGS structural elements of the ALH scaffold. In a previous study, Sales et al. demonstrated that type I collagen can adsorb to a PGS foam scaffold from dilute solutions, reaching a maximum surface concentration from solutions as dilute 20 μL mL−1 collagen.27 We thus expect that the surfaces of the PGS struts were saturated with adsorbed collagen under the conditions tested herein, and that upon exhausting the available PGS strut surface area, the growing tips of the collagen fibrils would tend to progress outward from the struts into the bulk collagen solution filling the pore. Indeed, proximal to the collagen gel-PGS strut interfaces, confocal reflectance micrographs qualitatively revealed that collagen fibrils were arranged not in parallel, but rather at finite angles or roughly perpendicular to the PGS struts (Figure 1C,D). In the case of the square diamond pore, in which the PGS struts were oriented at opposing angles of ±45° and at equal distances from each other, essentially equal fractions of the collagen fibrils were oriented at ±45°, yielding no single preferential angle of alignment (Figure 1E). By contrast, while the PGS struts were likewise oriented at ±45° in the ALH scaffold, the distances between opposing struts were longer along the PD versus XD direction, thereby offering a longer path for extension of collagen fibrils along the PD direction of the ALH pore.
We undertook FE simulations to predict what stiffnesses the collagen matrix would need to manifest in order for the effective stiffnesses and anisotropy of the ALH-collagen composite to match those of native left ventricular myocardium. FE simulations predicted the collagen would need to exhibit = 47.0 kPa and = 26.5 kPa (rcoll = 1.77) in order for the composite to reach 157 kPa and 84 kPa (r = 1.87).20 Simulations suggested two potential routes toward matching ALH-based constructs to left ventricular mechanical properties. In the context of heart cell-seeding,20, 22 the stiffness of the collagen gel would be expected to increase as the gel is contracted by the seeded cells; a potential limitation, however, could be debonding of the collagen from the PGS scaffold upon cell-mediated gel contraction. Toward such approaches, we have demonstrated that cells and collagen can be retained within the ALH pore upon stretching the ALH scaffold (Figure S5, Supporting Information). A broad range of cell-seeded collagen gel stiffnesses have been reported ranging from ∼37 kPa (estimated from Figure 4 of Feng et al.8) to 5.33 ± 1.33 MPa.28 These studies suggest simulation predicted collagen stiffnesses of 26.5–47.0 kPa could be achieved by an appropriate combination of collagen gel concentration, cell seeding density, and cultivation time. An alternative approach could involve co-varying the ALH scaffold structure (e.g., strut width) and PGS curing conditions (i.e., PGS modulus).21 We demonstrated by FE simulations that two distinct values of strut width (w) are capable of yielding an anisotropy ratio equal to that of left ventricular myocardium (i.e., r = 1.87): w = 20 μm or w = 140 μm.21 The 20 μm strut width would be both feasible to microfabricate and provide allowance for increased collagen matrix stiffnesses associated with heart cell-mediated contraction. As collagen fiber alignment alone is not sufficient to explain the high degree of anisotropy observed in fibroblast-seeded collagen gels,29 we speculate that the anisotropic collagen fibrillogenesis demonstrated herein, while significant, represents only a starting point in understanding the interplay between pore geometry, collagen morphology, and cell morphology. In future studies, Voronoi tessellation-based models could be useful in coupling collagen gel morphology to mechanical behavior.30 Further, the evolution of collagen anisotropy demonstrated in the present study may be extendable to other hydrogels, such as fibrin and Matrigel. Of particular note, Bian et al. demonstrated that muscle cell-laden fibrin-based hydrogels can be spatially patterned into anisotropic tissue bundles by casting within microfabricated poly(dimethysiloxane) molds.31 More broadly, collagen gel-based cell culture assays and drug delivery systems may manifest and potentially exploit anisotropic collagen fibrillogenesis, in particular in miniaturized composites of collagen gel and microfabricated or microscale materials. For example, in a miniaturized aortic ring assay introduced by Reed et al., 30 μL of collagen solution was gelled within and supported by a nylon mesh ring (3 mm inside diameter) comprised of ∼50 μm diameter fibers arranged in a square lattice (∼125 × 125 μm inside pore dimensions).1 As such systems are further miniaturized for high throughput screening, collagen morphology induced by the system boundaries could potentially impact the directionality of capillary sprouting; similar phenomena could potentially be exploited in microfluidic collagen gels mimicking human microvascular networks.32 In the context of drug delivery, De Paoli et al. have reported on the effects of oscillating magnetic fields on drug release from magnetic collagen gels (i.e., collagen gels containing iron oxide particles of up to 3 μm diameter).33 In magnetic collagen gels, structural changes within the gel associated with oscillating magnetic fields were demonstrated to impact drug release kinetics; similar effects could potentially be mediated by collagen gelation within the compartments of microfabricated drug delivery devices.34 In concert with auxiliary biophysical and biochemical regulators, compartmentalizing collagen gels within microfabricated materials represents a promising strategy for controlling collagen fibril anisotropy and associated functional performance parameters in advanced cell culture assay, drug delivery, and tissue engineering applications.