Injectable Polypeptide Hydrogels with Tunable Microenvironment for

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Injectable Polypeptide Hydrogels with Tunable Microenvironment for 3D Spreading and Chondrogenic Differentiation of Bone Marrow-derived Mesenchymal Stem Cells Kaixuan Ren, Haitao Cui, Qinghua Xu, Chaoliang He, Gao Li, and Xuesi Chen Biomacromolecules, Just Accepted Manuscript • DOI: 10.1021/acs.biomac.6b00884 • Publication Date (Web): 24 Oct 2016 Downloaded from http://pubs.acs.org on October 25, 2016

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Injectable Polypeptide Hydrogels with Tunable Microenvironment for 3D Spreading and Chondrogenic Differentiation of Bone Marrow-derived Mesenchymal Stem Cells

Kaixuan Ren, †, ‡ Haitao Cui, †, ‡ Qinghua Xu, †, ‡ Chaoliang He, *, † Gao Li, **, † Xuesi Chen†

†

Key Laboratory of Polymer Ecomaterials, Changchun Institute of Applied Chemistry,

Chinese Academy of Sciences, Changchun 130022, P. R. China ‡

University of Chinese Academy of Sciences, Beijing 100039, P. R. China

ABSTRACT: Bone marrow-derived mesenchymal stem cells (BMSCs) possess vast potential for tissue engineering and regenerative medicine. In this study, an injectable hydrogel comprising poly(L-glutamic acid)-graft-tyramine (PLG-g-TA) with tunable microenvironment was developed via enzyme-catalyzed crosslinking, and used as artificial extracellular matrix (ECM) to explore the behaviors of BMSCs during three dimensional (3D) culture. It was found that the mechanical property, porous structure as well as degradation process of the hydrogels could be tuned by changing the copolymer concentration. The PLGg-TA hydrogels showed good cytocompatibility in vitro. After being subcutaneously injected into the back of rats, the hydrogels degraded gradually within 8 weeks and exhibited good biocompatibility in vivo. BMSCs were then encapsulated in the polypeptide-based hydrogels with different copolymer concentration to investigate the influence of 3D matrix microenvironment on stem cell behaviors. It is intriguing to note that the BMSCs within the 2% 1 ACS Paragon Plus Environment

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hydrogel showed a well-spread morphology after 24 h and a higher proliferation rate during 7 days of culture, in contrast to a rounded morphology and lower proliferation rate of BMSCs in the 4% hydrogel. Furthermore, the hydrogels with different microenvironment also regulated the matrix biosynthesis and the gene expression of BMSCs. After incubation in the 2% hydrogel for 4 weeks, the BMSCs produced more type II collagen and expressed higher amounts of chondrogenic markers, compared to the cells in the 4% hydrogel. Therefore, the PLG-g-TA hydrogels with tunable microenvironment may serve as an efficient 3D platform for guiding the lineage specification of BMSCs. KEYWORDS: polypeptide hydrogels, enzymatic crosslinking, mesenchymal stem cells, matrix microenvironment, chondrogenic differentiation 1. INTRODUCTION Mesenchymal stem cells (MSCs) are one of the most commonly used stem cells for biomedical applications, which can self-renewal and undergo differentiation into multiple lineages, such as chondrocytes, osteoblasts, myoblasts, adipocytes and neurons.

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In stem

cell niches, the biochemical and biophysical cues, including soluble factors, matrix properties and cell-cell interactions, played crucial roles in manipulating stem cell behaviors.

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For

example, MSCs underwent specific-lineage differentiation towards neurons, myoblasts and osteoblasts if the cells were cultivated on collagen-coated polyacrylamide gels with elasticities mimicking those of nerve, muscle, and bone tissues, respectively.

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Stem cell-

based therapeutic strategies have been widely investigated as approaches to differentiate into chondrocytes for cartilage regeneration under the guidance of specific physicochemical and/or biological cues. 10 It has been reported that high cell seeding density and some soluble growth factors, including transforming growth factor-β (TGF-β), bone morphogenetic protein (BMP) and insulin-like growth factor (IGF), showed markedly induced ability for 2 ACS Paragon Plus Environment

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chondrogenic differentiation of MSCs.

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Besides, scaffolds with suitable properties could

contribute to chondrogenesis and cartilage reconstruction.

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Compared to 2D substrates,

the three dimensional (3D) artificial extracellular matrix (ECM) can provide more appropriate circumstance that mimic the cellular physiological environment in vivo. Besides, the effect of 3D artificial ECM on stem cell fate may cause different results due to the space restriction to the cells by the surrounding matrix. Recently, progressive attention has been paid to injectable hydrogels acting as 3D artificial ECM for the cultivation of stem cells, due to their physical properties resembling native ECM. Injectable hydrogels allow the permeation of nutrients and stem cell-laden hydrogels can be easily formed in situ, preventing the undesirable diffusion of precursor solutions and ensuring the efficient survival of stem cells. These properties render them vast potential as scaffolds for in situ stem cell culture and tissue reconstruction.

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Meanwhile, the physicochemical

properties of injectable hydrogels, which can convert into biochemical signals through mechanotransduction, strongly influence stem cell behaviors.

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For example, the MSCs

encapsulated in the MMP-sensitive hyaluronic acid (HA) hydrogels switched to a more spread morphology and displayed enhanced chondrogenesis compared with MMP-insensitive HA hydrogels.

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Increased crosslinking of the HA hydrogels resulted in a reduced cartilage

matrix content and promoted the hypertrophic differentiation of MSCs.

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Hence, in order to

achieve stem cell lineage specification, injectable hydrogels are expected not only to support stem cell survival, but also to guide stem cell differentiation through interactions with cell surface receptors. In the present work, an injectable polypeptide hydrogel comprising poly(L-glutamic acid)graft-tyramine (PLG-g-TA) was prepared through enzyme-mediated crosslinking in the presence of horseradish peroxidase (HRP) and hydrogen peroxide (H2O2), which was used as a scaffold to explore the effect of 3D matrix microenvironment on the stem cell behaviors. 3 ACS Paragon Plus Environment

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The polypeptide hydrogels were designed owing to their excellent biocompatibility and biodegradability, as well as their unique similarity to natural protein-based ECM.

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The

hydrogels with different microenvironments were controlled by changing the copolymer concentration and then served as 3D artificial ECM to investigate the spreading and chondrogenetic differentiation of bone marrow-derived mesenchymal stem cells (BMSCs) in vitro. This study may give important insights to design appropriate synthetic scaffolds for stem cell therapy and cartilage tissue engineering. 2. EXPERIMENTAL SECTION 2.1. Preparation of the Poly(L-glutamic acid)-graft-tyramine (PLG-g-TA) Hydrogels. The PLG-g-TA copolymer was first synthesized according to the procedure in the supporting information (Section S1.2). Then, the PLG-g-TA hydrogels were prepared via enzymecatalyzed crosslinking in the presence of HRP and H2O2 in PBS (0.01 M, pH 7.4). The PLGg-TA copolymer solution (200 µL) at a certain copolymer concentration was mixed with HRP solution (50 µL, final concentration: 1.6 units/mL) and H2O2 solution (50 µL, molar ratio of H2O2:TA = 0.4). If the sample was not flowing within 30 s after inverting the vial, it was regarded as a gel. The final concentrations of PLG-g-TA copolymer were 2%, 4% and 6% (w/v), and the corresponding hydrogels were abbreviated as 2% gel, 4% gel and 6% gel. The final concentrations of H2O2 were 4.1, 8.2 and 12.3 mM for 2%, 4% and 6% gel, respectively. The physicochemical properties, cytocompatibility in vitro, degradation and biocompatibility in vivo of PLG-g-TA hydrogels, were measured according to the procedures in the supporting information (Section S1.3-S1.8). 2.2. 3D Spreading of BMSCs within the Hydrogels In Vitro. BMSCs (1.0 × 105 cells for each well) were mixed homogeneously with the precursor solutions (300 µL for each well) containing PLG-g-TA copolymer, HRP and H2O2, followed by immediate seeding on the

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coverslip in a 6-well culture plate. The BMSC-laden hydrogels were rapidly formed at 37 oC without sedimentation of cells, followed by adding culture medium (1 mL) and incubating at 37 oC for 24 h. Subsequently, 4% paraformaldehyde and 0.1% Triton X-100 solution were used to fix and permeate the BMSCs within the polypeptide hydrogels, respectively. The cell nuclei and F-actin filaments were stained with 0.1% DAPI and 0.4% Alexa Fluor 488 phalloidin, respectively. The BMSC-laden hydrogels were washed with PBS repeatedly after each step. BMSCs seeded on the coverslip without hydrogels were used as a control. All the stained samples were observed by a laser scanning confocal microscope (Zeiss LSM 780, Germany). 2.3. 3D Proliferation of BMSCs within the Hydrogels In Vitro. Cell counting kit-8 (CCK-8) method was applied to test the 3D proliferation of BMSCs within the PLG-g-TA hydrogels, according to the procedure in the supporting information (Section S1.10). [24] 2.4. 3D Chondrogenic Differentiation of BMSCs within the Hydrogels In Vitro. To explore the chondrogenic differentiation of BMSCs within the PLG-g-TA hydrogels, the BMSC-laden hydrogels were incubated in DMEM containing 0.1 µM dexamethasone, 50 µM ascorbic acid, 6.25 mg/L insulin and 10 µg/L TGF-β1 (denoted as the induction medium). DMEM without the induction agents was used as a control medium. The BMSC-laden hydrogels were prepared in a 24-well culture plate (300 µL, 1.0 × 106 cells for each well) according to the procedure above. The BMSC-laden hydrogels were cultured in control medium or induction medium (1 mL), respectively. After induction for 4 weeks, the cell survival and chondrogenic differentiation of BMSCs within the hydrogels were evaluated. 2.5. Cell Survival and Morphology. After cultivating for 4 weeks, we employed livedead cell staining kit to detect the cell survival of BMSCs within the PLG-g-TA hydrogels, according to the procedure in the supporting information (Section S1.11). [24] Meanwhile, the


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morphology of BMSCs encapsulated in the polypeptide hydrogels was observed by scanning electron microscope (SEM). The samples were prepared according to the procedure in the supporting information (Section S1.5). The surface microstructure of BMSC-laden 2% gel was observed as the 2% gel shrunk during the drying process and was hard to crosscut. 2.6. Isolation of DNA. After incubation for 4 weeks, the BMSC-laden hydrogels were lyophilized and papain solution was then added to digest the sample at 56 oC overnight as previously described.

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The content of DNA was quantified by SYBR Green I fluorescent

dye assay (BioTeke Corporation) using a fluorescent plate reader. We used Calf thymus DNA as a standard. All the measurements were performed in triplicate. 2.7. Quantitative Real-time Polymerase Chain Reaction (RT-qPCR). The gene expression level of chondrogenic markers was analyzed by RT-qPCR, according to the procedure in the supporting information (Section S1.12). 2.8. Immunohistochemical Staining. The production and distribution of type II collagen in the PLG-g-TA hydrogels were detected by immunohistochemical staining. Briefly, cellladen hydrogels were embedded in OCT compound and kept at -80 oC. The samples were cut cryostat section to 10 µm at -20 oC and mounted on histological slides, followed by airdrying the sections for 30 min. H2O2 solution (0.3%) was applied to block endogenous peroxidase activity, and 1% bovine serum albumin was then used to block non-specific binding of the sections. The samples were then incubated with primary antibody (rabbit collagen II polyclonal antibody, 1:300 in PBS) at 4 oC overnight. Subsequently, the samples were cultured with biotin-labeled goat anti-rabbit IgG, and then incubated with horseradish peroxidase-conjugated streptavidin. The sections were washed with PBS repeatedly after each step and ultimately colored with DAB detection kit, according to the kit instructions. 3. RESULTS AND DISCUSSION 6 ACS Paragon Plus Environment

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As illustrated in Scheme 1, the PLG-g-TA copolymer was synthesized by conjugating poly(L-glutamic acid) (PLG) with tyramine (TA). [30] The 1H NMR spectra of PLG and PLGg-TA copolymer are shown in Figure S2 (Supporting Information), and all the peaks are well assigned. The phenyl peaks of TA (-C6H4-) at 6.4-7.0 ppm were observed in Figure S2B. The grafting ratio of TA groups was estimated to be 8% by comparing the integral area of the phenyl resonance peak of TA (-C6H4-) at about 6.4-7.0 ppm with that of the resonance peak assigned to the methylene in glutamate unit (-CH2CH2C(O)-) at about 1.7-2.0 ppm. Additionally, UV-Vis spectroscopy was employed to test the grafting ratio of TA in the PLGg-TA copolymer. We obtained a linear calibration curve of tyramine by plotting the absorbance of tyramine to its concentration in deionized water (Figure S3B). The content of conjugated TA groups was evaluated to be 8.8 wt% by using the fitting formula as following: y = 7.48x-0.02, R2=0.9999 (Figure S3, Supporting Information). The data was consistent with the NMR result. Therefore, these results suggested that the PLG-g-TA copolymer was successfully synthesized.


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Scheme 1. Synthetic route of the PLG-g-TA copolymer and schematic illustration of the cellladen hydrogels.

The PLG-g-TA hydrogels were obtained via an enzyme-catalyzed crosslinking process of the copolymer (Scheme 1). The phenol groups in the PLG-g-TA copolymer reacted with each other to form intermolecular crosslinks by the catalysis of HRP in the presence of H2O2, leading to the gelation of the copolymer solution.

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In our previous work, an

enzymatically-crosslinked hydrogel of PLG grafted with poly(ethylene glycol) (PEG) and tyramine (TA) (PLG-g-PEG/TA) was prepared. The weight contents of PEG and TA were 79.1 wt% and 3.9 wt% respectively based on the NMR results. It was found that the physicochemical properties of the PLG-g-PEG/TA hydrogels can be well regulated by changing the concentrations of HRP and H2O2. However, the PLG-g-PEG/TA hydrogels


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displayed high swelling ratio and kept the integrity for only one week in vitro due to the high content of the hydrophilic PEG segments.

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In this study, the PLG-g-TA hydrogels were

expected to exhibit stronger mechanical properties and better stability because of the relatively higher content of conjugated TA groups (8 wt%) and the absence of PEG segments. Besides, Zhang etc. reported a kind of poly(γ-glutamic acid)-tyramine hydrogel as a delivery system for the controlled release of bovine serum albumin.

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It’s worth mentioning that the

poly(γ-glutamic acid) (γ-PGA) is an anionic polypeptide comprising of D- and/or L-glutamic acid via γ-amide linkages, which exhibits different properties from our poly(α,L-glutamic acid) (PLG).

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In this work, we developed PLG-g-TA hydrogels with tunable

physicochemical properties by varying the copolymer concentration and focused on the 3D spreading and chondrogenic differentiation of BMSCs within the hydrogel with different microenvironments in vitro. The PLG-g-TA hydrogels with varying physicochemical properties were developed by adjusting the copolymer concentration. The hydrogels with the copolymer concentrations of 2% or 4% (w/v) were denoted as 2% gel and 4% gel, respectively. It was found that both 2% gel and 4% gel showed a fast gelation process (< 5 min) in the presence of HRP and H2O2 (1.6 units/mL HRP, H2O2/TA = 0.4 (mol/mol)), suggesting the rapid formation of the crosslinking network (data not shown). The rheological properties of the polypeptide hydrogels were measured during the gelation process. As shown in Figure 1A, storage modulus (G’) of the hydrogels increased quickly at the beginning, and a plateau was observed within 20 mins, implying the completion of the crosslinking reaction. It is noteworthy that the 4% gel displayed a markedly higher storage modulus (~ 15.3 KPa) than that (~ 4.2 KPa) of the 2% gel, which was attributed to the fact that the hydrogel formed from a higher copolymer concentration had a more compact crosslinking network.

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It is worth mentioning that

there was no gel formed when the PLG-g-TA concentration was set to 1%. Meanwhile, the 6% 9 ACS Paragon Plus Environment

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gel exhibited a much faster gelation process in the presence of HRP and H2O2 (1.6 units/mL of HRP, 12.3 mM of H2O2, H2O2/TA = 0.4 mol/mol), and displayed a higher storage modulus (~ 26.3 kPa) than those of 2% and 4% gel (Figure S4, Supporting Information), suggesting that we can tune the viscoelastic property of the polypeptide hydrogels by varying the copolymer concentration. However, considering the relatively high concentration of H2O2 for 6% gel formation, which may lead to potential cytotoxicity, 2% gel and 4% gel were selected for the following experiments. The in vitro degradation of the polypeptide hydrogels was assessed by incubating the hydrogels in Tris-HCl buffer solutions with or without elastase. The 2% gel disintegrated completely in 8 days in the presence of elastase (5.0 units/mL), while the 4% gel maintained its integrity up to 22 days under the same conditions (Figure 1B). Compared to the relatively rapid degradation in the presence of elastase, the degradation of the hydrogels sustained for much longer time in the absence of enzyme. In the initial stage, an obvious mass loss of the hydrogels was observed after incubation in buffer solutions without enzyme for 4 days, likely due to the fast degradation and diffusion of the less crosslinked copolymer segments in the hydrogels. Thereafter, the hydrogels showed a relatively slow degradation in the buffer solution without enzyme. Meanwhile, the degradation rate of the 4% gel was slower than that of the 2% gel in the buffer without enzyme. It is well known that the change of gel mass during the degradation process in vitro was resulted from the combined effects of swelling and degradation. 42 The degradation time and rate were affected significantly by the degree of crosslinking of the hydrogels.

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Hence, the PLG-g-TA hydrogels with higher copolymer

concentration showed longer degradation time, due to the enhancement in the degree of crosslinking. Besides, the presence of elastase accelerated the degradation of the hydrogels, ascribed to the enzymatic hydrolysis of the polypeptide chains. These results demonstrated


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that we could regulate the degradation behavior of the polypeptide hydrogels by adjusting the copolymer concentration. Additionally, the microscopic morphology of the PLG-g-TA hydrogels was observed with SEM. As shown in Figure 1C and 1D, the PLG-g-TA hydrogels showed interconnected porous structure. The average pore size of the 2% gel was 92 ± 11 µm (Figure 1C). By contrast, the 4% gel displayed a more compact network structure with the average pore size of 54 ± 17 µm (Figure 1D). The reduction in the pore size with increasing copolymer concentration was caused by an enhancement in the degree of crosslinking, which was consistent with the results of rheological property and degradation in vitro. It is worth mentioning that the porous structure of hydrogels influenced the permeation of bioactive molecules, nutrients and metabolites, and thereby played an important role in 3D cell culture. 46, 47

Overall, the mechanical strength, porous structure and degradation behavior of the polypeptide hydrogels, which strongly depended on the crosslinking degree of the hydrogels, were well tailored by adjusting the copolymer concentration. Compared with our previous PLG-g-PEG/TA hydrogels,

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the PLG-g-TA hydrogels displayed a higher storage modulus

and maintained its integrity for much longer time. It’s notable that these properties of the hydrogels are obviously correlated with each other and it’s difficult to change a single parameter of the hydrogels without impacting the others. Besides, these parameters are dynamic during the swelling and degradation process of the hydrogels, rendering the stem cell behaviors more complicated within the hydrogels during the 3D culture.


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Figure 1. (A) Storage (G’) and loss moduli (G’’) of the PLG-g-TA hydrogels measured in a time-dependent model. The hydrogels formed from 2% or 4% (w/v) of PLG-g-TA copolymer (1.6 units/mL of HRP and molar ratio of H2O2:TA = 0.4) were denoted as 2% gel and 4% gel, respectively. (B) In vitro degradation curves of the PLG-g-TA hydrogels in Tris-HCl buffer (0.05 M, pH 7.4) with 5.0 units/mL elastase or without elastase (control) (n=3). (C, D) SEM images of the freeze-dried hydrogels: (C) 2% gel, (D) 4% gel. The scale bars represent 200 µm.

The in vitro cytocompatibility of PLG-g-TA hydrogels was evaluated by MTT assay against L929 mouse fibroblasts or BMSCs. Considering that the cells were mixed with the precursor solutions of the hydrogels for a very short time due to the fast gelation process, the cells were eventually exposed to the extracting leachable materials of the hydrogels, rather than the copolymer solution with high concentration. Therefore, to evaluate the cytotoxicity of the PLG-g-TA hydrogels, the cells were incubated with the leaching solutions of the hydrogels or the copolymer solution at different concentrations. As shown in Figure S5A and 12 ACS Paragon Plus Environment

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S5B (Supporting Information), L929 cells or BMSCs incubated with the PLG-g-TA copolymer solutions at the concentration up to 0.5 mg/mL showed high viability. Additionally, L929 cells exposed to the leaching solutions of the hydrogels displayed high viability (> 87%, Figure S5C, Supporting Information), suggesting no obvious cytotoxicity of PLG-g-TA copolymer after gel formation and good cytocompatibility of the hydrogels in vitro. To evaluate the degradation and biocompatibility of the PLG-g-TA hydrogels in vivo, we injected the precursor solutions containing PLG-g-TA copolymer (2% or 4% (w/v)), HRP and H2O2 in the back of rats. At 15 min after injection, the 2% gel or 4% gel was observed in the subcutaneous layer of rats (Figure 2), indicating the rapid formation of the hydrogels in vivo. Subsequently, the status of the hydrogels in the back of the rats was observed at different time intervals. It was found that both the 2% gel and 4% gel diminished gradually and completely disappeared at 8 weeks post-injection. Hence, the results indicated that the proteolytic enzymes or reactive species in the subcutaneous layer of rats led to the disintegration of the polypeptide hydrogels. Compared with the degradation behavior of the PLG-g-TA hydrogels in vitro, the 2% gel and 4% gel showed no significant difference when degraded in the subcutaneous layer of rats. This is likely due to the fact that the degradation process of hydrogels in vivo was more complicated, resulted from the combined effects of multiple inflammatory cells, enzymes and reactive species.

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Moreover, the degradation

products of the hydrogels in vivo were ultimately phagocytized and eliminated through metabolism process, rather than effectively diffused into the medium solution in vitro.


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Figure 2. In vivo hydrogel status of 2% gel and 4% gel in the subcutaneous lay of rats at different time intervals. Images were taken at different time points after the injection.


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Moreover, the in vivo biocompatibility of the polypeptide hydrogels was assessed by H&E staining analysis of the skin tissues around the injection sites. Normally, the implantation of biomaterials perturbs the homeostasis of tissues and causes inflammatory response, therefore defense mechanisms to foreign body intend to eliminate the biomaterials and return to homeostasis. The inflammatory process is affected by the physicochemical properties of the biomaterials, determining the eventual success or failure of biomaterials. For the PLG-g-TA hydrogels, at the initial stage after injection (2 weeks), there was enhanced number of inflammatory cells in the skin tissues surrounding the 2% gel or 4% gel, implying an inflammatory reaction (Figure 3). It is worth noting that the number of inflammatory cells decreased and the inflammatory reaction eliminated gradually along with the disintegration of the hydrogels. The histology of the host tissues surrounding the hydrogels nearly restored to normal after injection for 8 weeks, suggesting the recovery of homeostasis. Compared with non-degradable implanted biomaterials, there was no obvious fibrous encapsulation or persistence of chronic inflammation reaction in the degradation process of the hydrogels. Besides, no obvious tissue edema and necrosis were observed during the experiment period. Hence, the results indicated an acceptable biocompatibility of the polypeptide hydrogels in vivo.


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Figure 3. H&E staining images of the skin tissues around the injection sites of 2% gel and 4% gel in the back of rats at different intervals.

BMSCs were encapsulated in the 2% and 4% gels to determine the effect of microenvironment on the 3D spreading and proliferation of BMSCs in vitro. After incubation


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within the 2% gel or 4% gel for 24 h, we stained the cell nuclei and F-actin filaments of BMSCs with DAPI and Alexa Fluor 488 phalloidin, respectively. BMSCs cultured on a 2D substrate were used as the control. As shown in Figure 4A, spindle-shaped and reticulate Factin filaments were observed for the BMSCs on the coverslip, suggesting cell adhesion and spread in a 2D monolayer culture. It is interesting to note that the BMSCs exhibited quite diverse morphologies within the PLG-g-TA hydrogels with different copolymer concentrations. Obvious filopodia and elongated actin fibers were observed for the BMSCs within the 2% gel after incubating for 24 h. Although the cells may be spatially hindered by the surrounding matrix in the 3D environment, the BMSCs within the 2% gel showed a wellspread morphology. In contrast, the BMSCs within the 4% gel displayed a spherical morphology without stress fibers. The cell adhesion to the ECM was mediated by focal adhesions and influenced by mechanical property, surface charge and wettability of the matrix.

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It has been reported that MSCs were able to extend within hydrogels with

relatively soft crosslinking networks.

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Therefore, the marked difference in cell

morphology within the 2% gel and 4% gel should be mainly attributed to the fact that the softer matrix and larger space of the 2% gel may facilitate the spreading of BMSCs during the 3D culture. Thus, the results clearly indicated that the polypeptide hydrogels with tunable microenvironment were capable of tuning the cell-matrix interactions and guiding the cell shape. Furthermore, we measured the proliferation of BMSCs within 2% or 4% gels by CCK-8 method. It was found that both the 2% gel and 4% gel supported the proliferation of BMSCs during the incubation for 7 days in vitro (Figure 4B). Moreover, a higher proliferation rate of BMSCs was observed in the 2% gel than in the 4% gel. It is well known that appropriate mitogens, sufficient nutrients as well as space should be available for cell proliferation.

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Compared to the 4% gel, the larger porous structure and better permeability for nutrients and 17 ACS Paragon Plus Environment

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metabolites of the 2% gel promoted cell proliferation. 53 In addition, the anchorage-dependent cells required adhesion and spreading for cell survival and proliferation.

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Hence, the cell

adhesion to the 2% gel and spreading of BMSCs in the 2% gel may also contribute to the faster cell proliferation.

Figure 4. (A) CLSM images of BMSCs within the 2% gel and 4% gel after incubation for 24 h. BMSCs cultured on the coverslip were used as a control. Alexa Fluor 488 phalloidin stained F-actin filaments and DAPI stained nuclei showed green and blue fluorescence, respectively. (B) The proliferation of BMSCs within the 2% gel and 4% gel measured by CCK-8 method (n=3) (**p

Injectable Polypeptide Hydrogels with Tunable Microenvironment for

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