Interplay of Matrix Stiffness and CellâCell Contact in Regulating

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Interplay of Matrix Stiffness and Cell−Cell Contact in Regulating Differentiation of Stem Cells Kai Ye, Luping Cao, Shiyu Li, Lin Yu, and Jiandong Ding* State Key Laboratory of Molecular Engineering of Polymers, Collaborative Innovation Center of Polymers and Polymer Composite Materials, Department of Macromolecular Science, Advanced Materials Laboratory, Fudan University, Shanghai 200433, China S Supporting Information *

ABSTRACT: Stem cells are capable of sensing and responding to the mechanical properties of extracellular matrixes (ECMs). It is well-known that, while osteogenesis is promoted on the stiff matrixes, adipogenesis is enhanced on the soft ones. Herein, we report an “abnormal” tendency of matrix-stiffness-directed stem cell differentiation. Well-defined nanoarrays of cell-adhesive arginine-glycine-aspartate (RGD) peptides were modified onto the surfaces of persistently nonfouling poly(ethylene glycol) (PEG) hydrogels to achieve controlled specific cell adhesion and simultaneously eliminate nonspecific protein adsorption. Mesenchymal stem cells were cultivated on the RGD-nanopatterned PEG hydrogels with the same RGD nanospacing but different hydrogel stiffnesses and incubated in the induction medium to examine the effect of matrix stiffness on osteogenic and adipogenic differentiation extents. When stem cells were kept at a low density during the induction period, the differentiation tendency was consistent with the previous reports in the literature; however, both lineage commitments were favored on the stiff matrices at a high cell density. We interpreted such a complicated stiffness effect at a high cell density in two-dimensional culture as the interplay of matrix stiffness and cell−cell contact. As a result, this study strengthens the essence of the stiffness effect and highlights the combinatory effects of ECM cues and cell cues on stem cell differentiation. KEYWORDS: matrix stiffness, stem cell differentiation, cell−cell contact, cell density, RGD nanopattern, PEG hydrogel Discher group about one decade ago,35 it has been widely acknowledged that stem cell fate is directed by matrix stiffness, with a stiff matrix beneficial for differentiation toward cell types of hard tissues such as osteoblasts, and a soft one favorable for differentiation toward cell types of soft tissues such as adipocytes. Recently, we observed an “abnormal” matrix-stiffness-related cellular phenomenonboth osteogenesis and adipogenesis of stem cells were promoted on a stiff matrix. Such an abnormal phenomenon happened only in the case of a high cell density in our two-dimensional in vitro cell culture and stem cell induction. Both stiffness dependences of osteogenic and adipogenic inductions of stem cells were normal at a low cell density. We further interpreted the “abnormal” phenomenon of stem cell differentiation at a high cell density as the interplay of matrix stiffness and cell−cell contact. Therefore, our results are essentially consistent with the classic stiffness effect on stem cell differentiation reported in the literature35,36,40,41 and, mean-

1. INTRODUCTION The latest decade has witnessed the booming development of stem cell research from bench to bedside.1,2 Because of the abilities to self-renew and differentiate into a wide range of specialized cell types, stem cells have become the most versatile and promising cell source for tissue regeneration.3,4 A stem cell lives in a microenvironment, where the neighboring cells, the surrounding extracellular matrixes (ECMs), and the soluble factors together constitute the stem cell niche.5 While most biologists focus on how the various soluble factors influence cell behaviors, biomaterial scientists pay more attention to the “insoluble” factors, namely, ECM cues and cell cues.6−9 Stem cell fate is modulated by the different aspects of ECMs and their mimics, such as chemical composition,10,11 surface topography,12−15 and so forth.16 In addition, cell size,17,18 cell shape,19,20 cell−cell contact,21−23 and cell density22,24 play crucial roles in regulating stem cell differentiation. The mechanical properties of ECMs, usually indicated as stiffness, have profound impact on cell adhesion,25−27 spreading,28−30 migration,31,32 proliferation,33,34 and differentiation.35−38 The relationship between matrix stiffness and stem cell phenotype receives particular concerns. Stem cells are able to feel and respond to the stiffnesses of ECMs, probably through a traction-force-mediated inside−outside-in mechanotransduction pathway.39 Since the famous work published by © XXXX American Chemical Society

Special Issue: Interfaces for Mechanobiology and Mechanochemistry: From 2-D to 3-D Platforms Received: October 13, 2015 Accepted: November 13, 2015

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DOI: 10.1021/acsami.5b09746 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

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Figure 1. Preparation of RGD-nanopatterned PEG hydrogels with the same surface chemistry but different hydrogel stiffnesses. (a) Synthesis routes of a PEG hydrogel. PEG reacted with acryloyl chloride in the presence of TEA, which acted as the proton scavenger, and the resultant PEG-DA macromer was photopolymerized to obtain the chemically cross-linked PEG hydrogel. (b) Fabrication processes of an RGD nanopattern on a PEG hydrogel. Gold nanodots were prepared on glass via block copolymer micelle nanolithography, then transferred to the PEG hydrogel via transfer lithography, and finally grafted with cyclic RGD peptides to form an RGD-nanopatterned PEG hydrogel. (c) Schematic illustration of the acquisition of RGD-nanopatterned PEG hydrogels with the same RGD nanospacing but different hydrogel stiffnesses by taking advantage of the same equilibrium swelling ratio of the two hydrogels.

while, significantly deepen the corresponding understanding. The present paper is aimed to report such a complicated stiffness effect and the interpretation. In light of material science, it is extremely difficult to fix other material parameters when one adjusts matrix stiffness. The key to the present report is a well-designed material platform. In this study, arginine-glycine-aspartate (RGD)-nanopatterned poly(ethylene glycol) (PEG) hydrogels were employed as the platform for the investigation of the effect of matrix stiffness on the lineage specifications of stem cells. Compared to other stiffness-tunable background materials, such as polyacrylamide (PA) gels, polydimethylsiloxane (PDMS), hyaluronic acid (HA) hydrogels, and alginate hydrogels, PEG hydrogels are

able to achieve very strong and persistent resistance to protein adsorption and nonspecific cell adhesion for more than one week.42−45 The interference of the effect of surface proteins adsorbed from the culture media could thus be thoroughly excluded from the effect of matrix stiffness in the present research. Specific cell adhesion can be enhanced by surface modification with cell-adhesive RGD peptides, which are the minimal integrin-binding sequences in certain natural ECM proteins, for instance, fibronectins.46−48 The transmembrane receptor integrin possesses a size of ∼10 nm and serves as an important signaling protein involved in cell-matrix interactions.49 Therefore, our well-controlled RGD nanoarrays with RGD binding sites of the similar sizes enabled the control of B


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dosage of 0.05% (w/w) of PEG-DAs. The macromers were polymerized under the protection of a nitrogen atmosphere by initiation of ultraviolet irradiation (365 nm) and D2959, and the polymerization lasted for 1 h. The resultant PEG hydrogels were collected and cut into round slices for further characterization. Measurements of linear swelling ratios and compressive moduli were conducted for the as-synthesized PEG hydrogels. Three parallel samples were tested for each type of PEG hydrogel. The diameter of a specimen before swollen was recorded as D1. After swollen to an equilibrium state in phosphate buffer saline (PBS) at 37 °C, the diameter was recorded as D2. Linear swelling ratio was calculated as D2 divided by D1. The compressive moduli of the swollen PEG hydrogels were obtained via the traditional stress−strain method. The original force−gap data were measured via unconfined submersion compression testing in a rheometer (Kinexus Pro, Malvern, U.K.) equipped with a parallel plate (diameter of 8 mm) at 37 °C. The stress−strain data could be calculated from the original force−gap data, and the compressive modulus was then determined from the slope of the stress−strain curve at low strains (≤5%). 2.3. Preparation of RGD-nanopatterned Poly(ethylene glycol) Hydrogels. RGD nanopatterns on nonfouling PEG hydrogels were generated as schematically depicted in Figure 1b. First, gold nanopatterns were prepared on glass surfaces via block copolymer micelle nanolithography.50 Block copolymer of poly(styrene-block-2vinylpyridine) (PS-b-P2VP, Polymer Source) was dissolved in toluene to form reversed micelles, with poly(2-vinylpyridine) (P2VP) segments as cores and polystyrene segments as coronae. A gold precursor hydrogen tetrachloroaurate(III) hydrate (HAuCl4·3H2O, Alfa Aesar) was added into the micelle solution and selectively bound to P2VP, and stable polymer−metal complexes were obtained. Then, glass was dipped into the micelle solution and pulled up with a specific velocity, and thus a self-assembled monolayer was coated onto the glass surface. Afterward, an oxygen plasma treatment was performed to remove the block copolymer template, and the gold precursor was reduced to gold after placed in the air. Thus, a quasi-hexagonal array of gold nanodots on glass was attained. Subsequently, Au nanopatterns were transferred from glass to PEG hydrogels via transfer lithography.51−53 Specimens were first immersed in the linker solution and grafted with a hetero-bifunctional linker N,N′-bis(acryloyl) cystamine (Sigma). Then, photopolymerization was performed to form chemically cross-linked PEG hydrogels as described above, resulting in the successful transfer of gold nanoarrays to the PEG surfaces. Finally, before cell seeding, gold-nanopattered PEG hydrogels were soaked in 25 μM c(−RGDfK−)−thiol (f: D-phenylalanine, K: L-lysine; Peptides International) solution at 4 °C for 4 h, and the cyclic oligopeptides were grafted onto the gold nanodots via S−Au covalent bonds. Excess peptides were removed by thorough rinse, and RGDnanopatterned PEG hydrogels were eventually acquired. 2.4. Morphological Observations of Nanopatterns. A transmission electron microscope (TEM, Tecnai G2 20 TWIN, FEI, U.S.A.) was utilized to observe the block copolymer micelles loaded with chloroauric acid. Gold nanopatterns on glass or PEG hydrogels were characterized in a field-emission scanning electron microscope (FE-SEM, Ultra 55, Zeiss, Germany). The water-containing PEG-4k hydrogels were dehydrated step by step in graded ethanol before the SEM observations. The nanoarrays on glass were also observed in an atomic force microscope (AFM, Multimode 8, Bruker, U.S.A). The diameters of the nanodots were analyzed from the AFM images by the software NanoScopeAnalysis, and the interparticle spacings were calculated from the FE-SEM images using the software ImageJ (freely available at http://www.nih.gov). Four parallel specimens were analyzed and averaged to attain the mean lateral spacing of the nanodots. 2.5. Isolation and Cultivation of Mesenchymal Stem Cells. MSCs were derived from tibias and femurs of 7 d old neonatal Sprague−Dawley (SD) rats. Primary cells were seeded in culture flasks. The growth medium was composed of low-glucose Dulbecco’s modified Eagle medium (DMEM, Gibco) supplemented with 10% fetal bovine serum (FBS, Gibco), 100 U/ml penicillin (Gibco), 100

integrin distribution and thus specific focal adhesion on the molecular level. Herein, we generated RGD-nanopatterned PEG hydrogels with two different compressive moduli but the same RGD nanospacing. This is not trivial, because the polymerized networks are usually swollen in aqueous environments such as cell culture media, and the two cross-linking densities usually lead to varied swelling ratios and thus different eventual RGD nanospacings. We prepared two PEG hydrogels via polymerization of the corresponding macromers, namely, poly(ethylene glycol)-diacrylates (PEG-DAs) of two different molecular weights and initial macromer concentrations (initial water content). The macromer with a longer chain length was grouped with a higher initial water content. In this way, we eventually achieved the same swelling ratio of the two hydrogels. Thus, the two hydrogels differed merely in the hydrogel stiffness, while the surface chemistry was kept constant. This enabled the determination of the matrix stiffness effect by ruling out other possible interferential factors. Because of the ease of isolation and cultivation, and the capacity for the lineage-specific commitments, mesenchymal stem cells (MSCs) derived from bone marrow of rats were used as the stem cell model. MSCs were seeded onto the two substrates and incubated in the osteogenic medium, the adipogenic medium, or the mixed medium for one week. Stem cells were seeded at a high density of 7000 cells per square centimeter and a low density of 1000 cells per square centimeter. The stiffness effects on both osteogenic and adipogenic differentiations of MSCs were examined at different cell densities.

2. MATERIALS AND METHODS 2.1. Synthesis and Characterization of PEG-DA Macromers. PEG-DA macromers with two number-average molecular weights (Mn values of 700 and 4000) were used in this study. While the former was purchased from Aldrich and used directly without further purification, the latter was synthesized in our lab according to the procedures schematically shown in Figure 1a. First, fresh acryloyl chloride was obtained by atmospheric distillation of benzoyl chloride (120 mL, 1 mol, Sigma-Aldrich) and acrylic acid (40 mL, 0.6 mol, Aldrich). PEG4k (20 g, 5 mmol, Aldrich) was dissolved in toluene (100 mL), and the solvent was distilled off to eliminate the trace amount of water. After it cooled to room temperature, the residue was redissolved in dried dichloromethane (DCM, 120 mL) under an argon atmosphere. Triethylamine (TEA, 5.6 mL, 40 mmol) and acryloyl chloride (1.6 mL, 20 mmol) were added under the conditions of ice-bath and protection from light, with the latter added dropwise within 3 h. The reaction lasted overnight at room temperature. The resultant mixture was filtered through a diatomaceous earth bed to remove triethylamine hydrochloride, and the filtrate was concentrated by vacuum rotary evaporation. The concentrated products were added slowly into diethyl ether under vigorous stirring and precipitated at −20 °C for 24 h. The precipitate was collected and washed with diethyl ether. The solid was dried under vacuum for at least 48 h. The final products were stored at −20 °C under vacuum before use. PEG-DA macromers were confirmed by 1H NMR spectra measured in a nuclear magnetic resonance (NMR) spectrometer (DMX 500, Bruker, U.S.A). Samples were dissolved in CDCl3 at a concentration of 5 mg/mL, and tetramethylsilane (TMS) was used as the internal reference. 2.2. Formation and Characterization of Poly(ethylene glycol) Hydrogels. Chemically cross-linked PEG hydrogels were obtained by photopolymerization of PEG-DA macromers (Figure 1a). Briefly, PEG-4k-DA was dissolved in Milli-Q water to achieve a concentration of 16.7% (w/w) before polymerization, while PEG-700DA was directly used. The photoinitiator 2-hydroxy-4′-(2-hydroxyethoxy)-2-methylpropiophenone (D2959, Aldrich) was added at a C


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Figure 2. Characterization of nanoarrays. (a) TEM image of the Au-loaded block copolymer micelles. The micelle suspension was dropped onto the copper meshes and dried under infrared lamp before the TEM observations. (b) AFM image of Au nanopattern on glass (left) and the corresponding height profile of the section along the white dashed line (right). (c) FE-SEM image of Au nanopattern on glass. μg/mL streptomycin (Gibco) and 2 mM L-glutamine (Gibco). Cell culture was maintained at 37 °C in a humidified atmosphere containing 5% CO2, and cells were subcultured upon 70−80% confluence. 2.6. Stem Cell Seeding on the Nanopatterned Surfaces. Prior to cell seeding on RGD-nanopatterned PEG hydrogels, the substrates were sterilized by 70% ethanol, put into 12-well tissue culture plates (TCPs), and rinsed thoroughly with PBS. MSCs of the second passage were seeded on the nanopatterned surfaces and utilized for the following cell experiments. 2.7. Cell Adhesion Assays. MSCs were allowed to adhere onto RGD-nanopatterned PEG hydrogels in the growth medium for 24 h after seeded at a density of 5000 cells per square centimeter. Then, all samples were fixed with 4% paraformaldehyde for 15 min and permeabilized with 0.1% Triton X-100 for 10 min. Specimens were blocked with 5% bovine serum albumin (BSA) in PBS to avoid nonspecific protein staining, and then incubated with 1:100 dilution of mouse monoclonal antivinculin primary antibody (Sigma) at 4 °C overnight, followed by triplicate rinse with PBS. Afterward, cells were labeled with 1:200 dilution of Alexa fluor 488-conjugated goat antimouse secondary antibody (Invitrogen) at room temperature for 2 h. To visualize filamentous actins (F-actins), samples were treated with 1 μg/mL phalloidin-tetramethyl rhodamine B isothiocyanate (Phalloidin-TRITC, Sigma) for 30 min. At last, cells were incubated with 2 μg/mL 4′,6-diamidino-2-phenylindole (DAPI, Sigma-Aldrich) for 8 min to stain cell nuclei, followed by thorough rinse with Milli-Q water. The fluorescently stained cells were observed in a laser scanning confocal microscope (LSCM, C2+, Nikon, Japan). Single cells were outlined using the software ImageJ to provide information about cell spreading areas and the fluorescent intensities of F-actins. The fluorescent intensity was calculated by the mean of pixel gray levels within the outline of a single cell minus the mean of its nearby background and then times the spreading area of the single cell, namely, the number of pixels within the cell outline. 2.8. Live/Dead Assays. Stem cells were cultured on the nanopatterned surfaces in the growth medium for 8 d and conducted with live/dead staining, using LIVE/DEAD Viability/Cytotoxicity Kit (Invitrogen) per manufacturer’s instructions. Cells were incubated with a mixture of ethidium homodimer-1 (4 μM) and calcein AM (2 μM) in PBS at 37 °C for 30 min. Then, the staining solution was replaced by PBS, and microscopic observations of the fluorescently stained cells were conducted in an inverted microscope (Axiovert 200, Zeiss, Germany) mounted with a charge coupled device (CCD, AxioCam HRC, Zeiss, Germany). 2.9. Stem Cell Induction. Osteogenic induction, adipogenic induction, or coinduction started following 1 d of cell adhesion onto the nanopatterned hydrogels, and the growth medium was replaced by the induction medium. For adipogenic differentiation, stem cells were incubated in the adipogenic induction medium for 3 d. The medium was composed of high-glucose DMEM, 10% FBS, 1 μM dexamethasone (Sigma), 200 μM indomethacin (Sigma), 0.5 mM 3-isobutyl-1methylxanthine (Sigma) and 10 μg/mL insulin (Sigma). Next, cells were incubated in the adipogenic maintenance medium (high-glucose

DMEM, 10% FBS, 10 μg/mL insulin) for 2 d, and in the adipogenic induction medium for another 2 d. For osteogenic differentiation, stem cells were exposed to the osteogenic induction medium for 7 d, which was composed of high-glucose DMEM, 10% FBS, 100 nM dexamethasone, 50 μM ascorbic acid-2-phosphate (Sigma), and 10 mM β-glycerophosphate (Sigma). The medium was changed on the fourth and sixth days of induction to keep pace with the change of the adipogenic medium. We also tried osteogenic and adipogenic coinduction for 7 d. Stem cells were exposed to the mixed medium, which was a 1:1 mixture of the osteogenic induction medium and the adipogenic induction/maintenance medium.24 In the stem cell differentiation assays, cells were seeded at two densities, a high density of 7000 cells per square centimeter, and a low density of 1000 cells per square centimeter. To inhibit cell proliferation, 0.5 μg/mL Aphidicolin (Sigma) was added to the induction medium chronically. 2.10. Determination of Cell Differentiation Extents. After 7 d of induction, cells were conducted with alkaline phosphatase (ALP) or/and oil-droplet staining. The former and the latter were indicators of osteoblasts and adipocytes, respectively. After osteogenic induction, cells were fixed with 4% paraformaldehyde for 15 min and then incubated in the aqueous solution of Fast Blue RR salt (Sigma) supplemented with naphthol AS-MX phosphate (Sigma) for 30 min. Adipogenically induced cells were fixed and treated with 3 mg/mL Oil Red O (Sigma-Aldrich) in 60% isopropanol. Cells were counterstained with DAPI to enable the counting of cell numbers. For the coinduced specimens, we stained ALP first, then oil droplets, and finally cell nuclei. The stained cells were photographed for the later data analysis. Osteoblasts were stained in blue with Fast Blue RR, and cells with red lipid vacuoles were set as adipocytes. Otherwise, the cells were considered as undifferentiated ones. The percentages of osteoblasts and adipocytes were calculated to determine osteogenic and adipogenic differentiation extents, respectively. 2.11. Statistical Analysis. For cell adhesion assays, 600 cells from three independent experiments for each group were included in data analysis and box-and-whisker plot mapping. For stem cell differentiation assays, four independent experiments participated in statistical analysis for each group, and the final results were demonstrated as mean ± standard deviation using the common histogram mapping. Data were analyzed by unpaired two-sample Student’s t-tests to examine the difference between the two relevant samples. A difference was regarded as significant when p < 0.05. Statistically significant difference was denoted as “*” (p < 0.05, significant difference), “**” (p < 0.01, significant difference), or “***” (p < 0.001, very significant difference).

3. RESULTS 3.1. Preparation of Poly(ethylene glycol) Hydrogels of Two Stiffnesses. After the reaction of PEG and fresh distilled acryloyl chloride using TEA as the proton scavenger, the protons of the terminal hydroxyl groups of PEG were substituted with the acryloyl groups to obtain PEG-DA D


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ACS Applied Materials & Interfaces macromer, which was confirmed by the characteristic peaks at δ 6.4−6.5, 5.8−5.9, and 6.1−6.2 ppm in the 1H NMR spectra of the resultant PEG-DA, as shown in Supporting Information Figure S1. The as-synthesized PEG-DA macromers were photopolymerized to form PEG hydrogels. By modulating the molecular weight and the concentration of PEG-DA (Supporting Information Table S1), PEG hydrogels of two different stiffnesses were acquired. The relatively softer one owned a compressive modulus of ∼130 kPa, and the stiffer one had a modulus of ∼3170 kPa. After swollen in PBS at 37 °C, both hydrogels gave an equilibrium swelling ratio of 1.15. 3.2. Preparation of RGD-Nanopatterned Poly(ethylene glycol) Hydrogels with the Same RGD Nanospacing but Different Hydrogel Stiffnesses. RGDnanopatterned PEG hydrogels were prepared by combining the nanopatterning technique and the transfer strategy, as schematically presented in Figure 1b. The successful loading of chloroauric acid into the cores of the block copolymer micelles and the uniform dispersion of the reversed micelles in the toluene solution were confirmed by TEM observations (Figure 2a). According to the AFM image, the sizes of the nanodots on the glass surface were determined to be ∼7 nm (Figure 2b), which was smaller than the size of an integrin (∼10 nm), and thus guaranteed the one-to-one binding of RGDs to integrins in the later cell experiments. The mean lateral spacing of the nanodots was 43 nm, as analyzed and averaged from several FE-SEM images (Figure 2c). The gold nanoarrays were then successfully transferred to PEG hydrogels of two different stiffnesses, as shown in Supporting Information Figure S2. As mentioned above, these two hydrogels possessed the same linear swelling ratio in PBS at 37 °C. Thus, after the nanopatterns were transferred from glass to PEG hydrogels and grafted with RGD peptides, the RGD nanospacing would be the same for these two hydrogels even after experiencing swollen in the cell culture media at 37 °C (Figure 1c). The average RGD nanospacing after swollen was 49 nm for both hydrogels (Supporting Information Table S2). Consequently, the RGD-nanopatterned PEG hydrogels were of the same surface chemistry but different hydrogel stiffnesses. 3.3. Cell Adhesion on RGD-Nanopatterned Poly(ethylene glycol) Hydrogels of Two Stiffnesses. Stem cells were cultured on RGD nanopatterns and immunofluorescently stained to investigate the matrix stiffness effect on cell adhesion. Representative fluorescence images of vinculins, Factins, cell nuclei, and the merged ones are shown in Figure 3. Cells adopted a more spread morphology on the stiff hydrogels. Vinculins were remarkably observed to be densely distributed around the cell periphery on the stiff hydrogels, indicating the robust cell-material interactions between the transmembrane receptors and the ECM ligands.54 Clearer and better-aligned actin filaments were seen on the stiff hydrogels as well, which suggested higher cell tension. Quantitative results of cell adhesion are displayed in Figure 4. This kind of box-and-whisker plots gives a visual presentation of the dispersion of the statistical data. Differently from the common histogram, which presents the mean value and standard deviation, a box-and-whisker plot provides information about the mean, median, quartiles, minimum, and maximum, and specifically identifies outliers (Supporting Information Figure S3). As seen from Figure 4a, cell spreading area was significantly larger on the stiff hydrogels. Analogously,

Figure 3. Fluorescence micrographs of adherent cells on RGDnanopatterned PEG hydrogels of the indicated two stiffnesses. Stem cells were allowed to adhere onto the substrates for 24 h, followed by immunofluorescent staining to visualize vinculins (green), F-actins (red), and cell nuclei (blue).

the fluorescent intensity of F-actins was higher on the stiff hydrogels as well (Figure 4b). In general, the stiff hydrogels were conducive to cell spreading and actin polymerization, which resulted in more cell tension. 3.4. Long-Term Cell Viability on the Nanopatterned Surfaces. Cell viability was examined by live/dead assays. Live cells after 8 d of culture in the growth medium were stained in green with calcein AM, and dead ones were stained in red with ethidium homodimer-1. The vast majority of the cells were alive, and dead cells could scarcely be seen on both the soft and the stiff hydrogels (Supporting Information Figure S4). This result reveals the good cytocompatibility of our RGDnanopatterned PEG hydrogels of two different stiffnesses, which makes the later differentiation results comparable between these two groups. 3.5. Persistent Antiadhesion of Poly(ethylene glycol) Hydrogels as the Backgrounds of Nanopatterns. To E


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Typical micrographs of cells after coinduction are presented in Figure 5a. When stem cells were exposed to the mixed

Figure 4. Quantitative adhesion parameters of MSCs on hydrogels of the indicated two stiffnesses. Cell spreading area (a) and normalized Factin intensity (b) are presented in the form of box-and-whisker plots (n = 600), with the meaning of the symbols presented in Supporting Information Figure S3. F-actin intensity was normalized to the mean value of the cells on the soft hydrogels. “***”: p < 0.001, very significant difference. The p values of Student’s t-tests of the data are listed in Supporting Information Tables S3 and S4.

Figure 5. Stem cell differentiation on RGD-nanopatterned PEG hydrogels of the indicated two stiffnesses in the mixed osteogenic and adipogenic medium. (a) Representative bright-field images of cells and the corresponding fluorescence images of cell nuclei after osteogenic and adipogenic coinduction. Osteoblasts were stained in blue with Fast Blue RR, adipocytes were stained in red with Oil Red O (upper), and all cell nuclei were stained with DAPI (lower). (b) Quantitative cell differentiation extents with respect to the two hydrogel stiffnesses. Mean values and standard deviations from four independent experiments are shown. “***”: p < 0.001, very significant difference. The p values of Student’s t-tests of the data are listed in Supporting Information Tables S5 and S6.

exclude the interference of nonspecific protein adsorption onto the substrates throughout the induction period, the persistent resistance of PEG hydrogels to cell adhesion must be guaranteed. This property was well-demonstrated by the dipline, a borderline dividing the RGD-grafted area from the nude PEG background (Figure 1b). Even after 8 d of adhesion and induction, cells were still restrictedly localized in the RGDnanopatterned regions, and the PEG backgrounds did not support any significant cell adhesion, as shown in Supporting Information Figure S5. The diplines were well-identified, and cells did not step over the diplines to the PEG backgrounds. This selective cell adhesion demonstrates that the PEG hydrogels are highly potent and persistent in resisting nonspecific cell adhesion. Hence, the subsequent results of the stiffness effect on stem cell differentiation are very convincing. 3.6. “Abnormal” Stem Cell Differentiation on RGDNanopatterned Poly(ethylene glycol) Hydrogels of Two Stiffnesses at a High Cell Density. To assess the effect of hydrogel stiffness on osteogenic and adipogenic differentiations of stem cells, MSCs were seeded on RGD-nanopatterned PEG hydrogels at a density of 7000 cells per square centimeter and underwent osteogenic, adipogenic, or coinduction.

osteogenic and adipogenic medium, the stiff hydrogels were beneficial for osteogenesis, and the soft hydrogels were favorable for adipogenesis (Figure 5b). This tendency is consistent with our previous study based upon coinduction of MSCs.38 Representative micrographs of osteogenically or adipogenically induced cells are shown in Figure 6a,b. The differentiation tendency differed from that of coinduction. While osteogenesis was, as expected, promoted on the stiff hydrogels (Figure 6c), we found that the stiff hydrogels were conducive to adipogenesis as well (Figure 6d). This interesting phenomenon, especially the abnormal tendency of adipogenesis, seems to contradict with the well-recognized matrix stiffness effect on the surface, and it deserves to be further investigated and discussed. F


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Figure 6. Stem cell differentiation on RGD-nanopatterned PEG hydrogels at a high cell seeding density of 7000 cells per square centimeter. (a, b) Representative optical micrographs of cells after osteogenic (a) or adipogenic (b) induction. (c, d) Quantitative osteogenic (c) and adipogenic (d) differentiation extents with respect to the two hydrogel stiffnesses. Mean values and standard deviations from four independent experiments are presented. “**”: p < 0.01, significant difference; “***”: p < 0.001, very significant difference. The p values of Student’s t-tests of the data are listed in Supporting Information Tables S7 and S8.

3.7. Normal Stem Cell Differentiation on RGD-Nanopatterned Poly(ethylene glycol) Hydrogels of Two Stiffnesses at a Low Cell Density. We hypothesized that the above abnormal trend might arise from the enhanced cell− cell contact effect on cell differentiation due to more cell spreading on the relatively stiffer matrices, and the inherent straightforward stiffness effect on stem cell differentiation might be better revealed at a low cell density without obvious cell− cell contact. So we further examined the matrix stiffness effect on osteogenic and adipogenic differentiations of stem cells at a low density of 1000 cells per square centimeter. Representative optical images of cells after induction are demonstrated in Figure 7a,b. While enhanced osteogenesis was found on the stiff hydrogels (Figure 7c), adipogenesis was favored on the soft ones (Figure 7d). On this occasion, the differentiation tendency agrees with the previous reports in the literature.35,36,40,41 The present study also reports the significant cell-density dependence of the effects of matrix stiffness on differentiation of stem cells after spreading on the twodimensional surfaces.

enhance osteogenesis and reduce adipogenesis, owing to the high cell tension. The distinctive findings reported in the present paper, however, complicate the effect of matrix stiffness on the lineage commitments of stem cells. In this study, we successfully prepared cell-adhesive RGD nanopatterns on nonfouling PEG hydrogels of two different stiffnesses. RGD peptides were employed, because specific focal adhesion could be initiated by the binding of RGD and its transmembrane receptor integrin.59−63 Owing to the same equilibrium swelling ratio of the two hydrogels in the cell culture media at 37 °C due to our careful design and try, the average RGD nanospacing was the same for both hydrogels in our study, as schematically presented in Figure 1c. Although it is well-known that RGD peptides induce osteogenic differentiation of MSCs,64−66 the interference of the effect of RGD peptides on MSC differentiation to the effect of matrix stiffness could be ruled out in the present study, due to the same RGD nanospacing of our two hydrogels. Furthermore, the persistent resistance of the PEG hydrogels to protein adsorption led to the elimination of nonspecific cell adhesion even after 8 d (Supporting Information Figure S5). Therefore, matrix stiffness would be the only variable, since the surface chemistry remained the same for the two hydrogels. The discrepant differentiation extents of stem cells would be exclusively derived from the stiffness difference. Our experimental results of cell adhesion confirmed better cell spreading and more cell tension on the stiff hydrogels (Figure 4), which was consistent with the previous findings of cell adhesion on other substrates.67 It is believed that more cell

4. DISCUSSION Matrix stiffness has been known to be one of the pivotal biophysical signals of ECMs to regulate stem cell differentiation since the latest decade.35,38,41,55,56 A mechanistic understanding of matrix-stiffness-directed stem cell differentiation is that stem cells exert traction forces on ECMs, gauge the mechanical resistance of ECMs to the forces, and make decisions on their lineage specifications.57,58 The stiff matrices are believed to G


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Figure 7. Stem cell differentiation on RGD-nanopatterned PEG hydrogels at a low cell seeding density of 1000 cells per square centimeter. (a, b) Representative optical micrographs of cells after osteogenic (a) or adipogenic (b) induction. (c, d) Quantitative osteogenic (c) and adipogenic (d) differentiation extents with respect to the two hydrogel stiffnesses. Mean values and standard deviations from four independent experiments are presented. “***”: p < 0.001, very significant difference. The p values of Student’s t-tests of the data are listed in Supporting Information Tables S9 and S10.

To further figure out how matrix stiffness influenced stem cell differentiation at a low cell density, we decreased the cell seeding density from 7000 cells per square centimeter to 1000 cells per square centimeter. Under this circumstance, the stiff hydrogels were conducive to osteogenesis (Figure 7c), and the soft ones were beneficial for adipogenesis (Figure 7d), which was in conformity with the aforementioned publications.36,41 By this token, cell density would indeed influence the matrix stiffness effect on stem cell differentiation. Then, what could be the essential cue apart from matrix stiffness at a high cell density? We speculated that cell−cell contact might play an important role. The cellular microenvironments are actually dependent upon cell density. There exists a critical contact density dcontact, below which cells are isolated, and beyond which cell−cell contact forms; and dcontact is negatively correlated to cell spreading area in twodimensional culture.22 Cell−cell contact has been found to promote osteogenesis, adipogenesis, and chondrogenesis of stem cells.21,23 Cell−cell contact is ubiquitous in the cell culture media at a high cell density, which makes it the strongest candidate besides matrix stiffness to influence stem cell differentiation in our research system. In general, the interplay of matrix stiffness and cell−cell contact gives rise to the discrepant differentiation tendencies of stem cells at different cell densities, as schematically illustrated in Figure 8. At a low density, cells are isolated from each other. The stiff matrices bring about more cell tension, which results

spreading and cell tension are favorable for osteogenesis and unfavorable for adipogenesis.24 It seems that the stiff hydrogels would be beneficial for osteogenesis, and adipogenesis would be enhanced on the soft hydrogels. Nonetheless, our cell differentiation results exceeded the normal expectation. When sole osteogenic and sole adipogenic inductions of MSCs were examined, both osteogenesis and adipogenesis were favored on the stiff hydrogels (Figure 6c,d). In the osteogenic and adipogenic coinduction assays, there existed competitions between these two differentiation directions. Under this circumstance, osteogenesis was promoted, and adipogenesis was suppressed on the stiff hydrogels (Figure 5b), which is not difficult to be understood by the overwhelming competitive advantage of osteogenesis over adipogenesis. Yet, why do our results in the case of more standard in vitro induction (sole induction) differ from the “common senses”, especially the adipogenic differentiation tendency? Is there any additional cue under the apparent matrix stiffness effect? After consulting some prestigious literature concerning the relationship between matrix stiffness and MSC differentiation,35,41,68−70 particularly based upon the fundamental studies of the effects of cell−cell contact and cell density on differentiation of MSCs in our group,21−23 we realized that our observation of the abnormal phenomenon did not contradict with the common senses in this field at all and that the widely accepted stiffness effect was mostly based upon a relatively low cell density. H


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Figure 8. Schematic illustration of the interplay of matrix stiffness and cell−cell contact in regulating stem cell differentiation. (a) At a low cell density, stem cells can hardly form cell−cell contact. Cells spread better and generate more cell tension on the stiff matrices, and thus osteogenesis is favored. (b) At a high cell density, besides better cell adhesion and more cell tension, the larger cell spreading sizes bring about more cell−cell contact on the stiff matrices as well. Taking both cell tension and cell−cell contact into consideration, the stiff matrices lead to definitely more osteogenesis and possibly more adipogenesis. This complicated stiffness effect at a high cell density might result in the abnormal differentiation tendency for adipogenesis, if the stiffness-enhanced cell−cell contact effect outweighs the direct inherent stiffness effect in regulating stem cell differentiation.

induction. While stiff matrices were always beneficial for osteogenesis, soft matrices favored adipogenesis only conditionally, depending upon cell density. An unusual differentiation tendency was observed at a high cell density. This complicated stiffness effect was further attributed to the interplay of matrix stiffness and cell−cell contact. Besides the inherent and thus the direct effect of matrix stiffness on cell differentiation, the enhanced cell−cell contact due to more spreading of cells on the stiff matrices influenced fate decisions of stem cells on material surfaces as well. Our findings provide new insights into matrix-stiffness-directed stem cell differentiation and afford a comprehensive understanding of cellmatrix and cell−cell interactions in regulating stem cell fate.

in more osteogenesis and less adipogenesis (Figure 8a). This is the so-called normal stiffness effect. After the cell density is increased to an adequately high value, cell−cell contact starts to play an additional role in regulating stem cell differentiation. Apart from the mechanotransduction pathway as mentioned above in the normal stiffness effect, the stiff matrices also bring about larger cell spreading sizes; further, they lead to the decrease of dcontact and thus more cell−cell contact at a fixed high cell density, which promotes both osteogenesis and adipogenesis (Figure 8b). The effects of cell tension and cell−cell contact on osteogenesis of stem cells are synergistic; thus, the stiff matrices are undoubtedly beneficial for osteogenesis. In contrast, adipogenesis is confronted with the antagonistic effects of cell tension and cell−cell contact. Adipogenesis might possibly be enhanced or suppressed on the stiff matrices. Hence, it is the interplay of matrix stiffness and cell−cell contact that leads to the complicated stiffness effect at a high cell density and accounts for the abnormal differentiation tendency of stem cells reported in this paper. We would like to point out that cell−cell contact is omnipresent in most of the tissues of living organisms. The stem cell niche in vivo is extremely complicated, and apart from cell−matrix interactions, cell−cell interactions must be taken into consideration as well.71,72 Moreover, the in vivo stem cell niche is a three-dimensional environment, which makes the situation much more intricate.37,40,73

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ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsami.5b09746. 1 H NMR spectra of PEG-DA macromers, FE-SEM images of Au nanopatterns on PEG hydrogels, schematic illustration of the box-and-whisker plot, fluorescence micrographs of the cells after live/dead staining, phasecontrast micrographs of the cells adjacent to the diplines, parameters of the synthesized PEG hydrogels, parameters of the prepared nanopatterns, and the p values of Student’s t-tests. (PDF)

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5. CONCLUSIONS On the strength of material techniques, we acquired RGDnanopatterned PEG hydrogels with the same surface chemistry but different hydrogel stiffnesses, and we investigated the effects of matrix stiffness on the lineage specifications of MSCs at different cell densities in two-dimensional in vitro culture and

AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Phone: 86 21 65643506. Fax: 86 21 65640293. Notes

The authors declare no competing financial interest. I


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ACKNOWLEDGMENTS The authors are grateful for the financial support from NSF of China (Grant Nos. 51533002 and 51273046) and Science and Technology Developing Foundation of Shanghai (Grant No. 13XD1401000).

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