Cross-Linkable Gelatin-CMC Hydrogels Designed for Rapid

May 6, 2016 - Vascular endothelial cells encapsulated in the gelatin-CMC hydrogels were .... hydrogels was estimated by a live/dead fluorometric assay...
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Article pubs.acs.org/journal/abseba

In Situ Cross-Linkable Gelatin-CMC Hydrogels Designed for Rapid Engineering of Perfusable Vasculatures Tatsuto Kageyama,† Tatsuya Osaki,‡ Junko Enomoto,† Dina Myasnikova,† Tadashi Nittami,† Takuro Hozumi,§ Taichi Ito,§ and Junji Fukuda*,† †

Graduate School of Engineering, Yokohama National University, 79-5 Tokiwadai, Hodogaya-ku, Yokohama, Kanagawa 240-8501, Japan Graduate School of Pure and Applied Sciences, University of Tsukuba, 1-1-1 Tennodai, Tsukuba, Ibaraki 305-8573, Japan § Center for Disease Biology and Integrative Medicine, The University of Tokyo, 7-3-1 Hongo, Bunkyo-ku, Tokyo 113-8656, Japan ‡

S Supporting Information *

ABSTRACT: Hydrogels that can be rapidly cross-linked under physiological conditions are beneficial for the engineering of vascularized 3-dimensional (3D) tissues and organs, in particular when cells are embedded at a high cell density or tissues are fabricated using bottom-up processes, including bioprinting and micromolding. Here, we prepared a gelatin-carboxymethylcellulose (CMC) hydrogel that cross-linked rapidly (1800

86 ± 16 240 ± 47 426 ± 19 N/A

LOW MID HIGH COL

the 96-well plate were suspended in 0.5 mL of gelatin-ADH solution (∼24 spheroids/mL) and mixed with 0.5 mL of CMC−CHO solution using a double-barreled syringe in a 35 mm culture dish. After 7 days of culture, confocal laser scanning microscopic images were acquired to visualize the sprouting of GFP-HUVECs (LSM 700; Carl Zeiss, Jena, Germany). Image processing software (IMALIS; Carl Zeiss) was used to quantify the average length of sprouting that extended from a spheroid surface. Electrochemical Transfer of Cells to the Gelatin-CMC Hydrogel. We previously showed that cells adhered onto the gold substrate through the zwitterionic oligopeptide monolayer could be detached along with desorption of the monolayer by applying an electrochemical potential.23 To prepare the substrates, flat glass substrates (24 × 24 mm) were sputter-coated with a few nanometers of chromium and ∼40 nm gold, followed by modification with the oligopeptide CGGGKEKEKEKGRGDSP by immersing the substrates in 100 μM oligopeptide solution overnight. After rinsing with distilled water, the substrates were

sterilized with 70% ethanol and placed in a 6-well plate. GFP-HUVECs were seeded on the substrate at a density of 1.0 × 105 cells/mL in 2 mL of culture medium. After 2 days of culture, the cells were washed with PBS and covered with the MID hydrogel. Next, an electric potential of −1.0 V vs Ag/AgCl electrode was applied for 5 min using a potentiostat (HA-151; Hokuto-Denko, Tokyo, Japan). The hydrogel was then peeled away from the substrate. To quantify cell transfer to the hydrogel, the number of cells present on the substrates and hydrogels were counted. The experiment was repeated without oligopeptide modification for comparison (cells were attached directly to the gold surface, and an electrical potential was applied). Rapid Fabrication of Endothelialized Microchannels. Electrochemical cell transfer using cylindrical rods was used to fabricate perfusable vascular-like structures (Figure 1d). Gold-coated glass rods were modified with the oligopeptide as described above. GFP-HUVECs were seeded on the rods in a 35 mm noncell-adhesive culture dish at a density of 2.0 × 106 cells/dish. After 24 h of culture, excess cells were C

DOI: 10.1021/acsbiomaterials.6b00203 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX

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ACS Biomaterials Science & Engineering removed by exchanging the culture medium. Culture was continued for 2−3 days until the cells reached confluence. A culture chamber was composed of poly(methyl methacrylate) plates with holes for fixing the rods in a spatially controlled manner. The rods with the coated cells were placed in the chamber, and 0.5 mL of MID hydrogel solution was poured and gelled in the chamber. The cells were transferred from the rods to the hydrogel by applying −1.0 V vs the Ag/AgCl reference electrode for 5 min, and the rods were carefully extracted. The chamber was connected to a microsyringe pump, and the culture medium was perfused at 10 μL/min. The cells were fixed at 4 days of culture and stained with rhodaminephalloidin and DAPI to visualize F-actin filaments and the cell nuclei, respectively.

different concentrations were prepared in the same mixing ratios (Table 1). By mixing the two hydrogel component solutions through a double-barreled syringe, transparent hydrogels were formed under all three conditions (Figure 1c). The gelation process is based on the spontaneous dehydration and condensation of the hydrazide and aldehyde groups, which leads to the formation of multiple hydrizone bonds between gelatin and the polysaccharide chains. The reaction occurs at room temperature and only produces water (H2O) as a byproduct (Figure 1b). Gelation time was dependent on the hydrogel concentration and occurred within 3 min, even for the LOW hydrogel (Table 1). These results indicate that our gelation process was at least 10-fold faster than that of collagen (typically >30 min). The mechanical properties of the hydrogels were characterized using a texture analyzer (Figure S3), revealing that the compressive modulus increased when the hydrogel concentration was increased (Table 1). Collagen gel could not be tested using the analyzer because of its excessively low elasticity, even at the maximum concentration (0.24 w/v) in the commercial products. We previously described gelatin-hyaluronic acid (HA) hydrogels as rapidly cross-linkable cell-laden hydrogels.17 However, these hydrogels swelled significantly and degraded within a few days. Thus, in the present study, we used CMC, which is a semisynthetic natural polymer obtained from the carboxymethylation of natural cellulose with well-documented biocompatible properties, low toxicity, and a low degradation rate compared to those of other polysaccharides.21,27,28 These properties may be beneficial for modulating the swelling and degradation of gelatincomposite hydrogels. Swelling of the gelatin-CMC hydrogels was significantly reduced compared to that of gelatin-HA hydrogels, and the MID and HIGH hydrogels maintained their shapes for at least 7 days in culture (Figure S1). The pore size of the hydrogels, which is important in determining the cross-linking density, was measured by freeze-drying the hydrogels, preparing cross-sections, and acquiring scanning electron microscope images (Figure 2a−d). The pore sizes of the hydrogels decreased with increasing concentration (Figure 2e). Hydrogels with pore sizes lower than 100 μm were too rigid to allow endothelial cell migration and sprouting.29 Since the pore size of HIGH hydrogels was lower than this value, the HIGH hydrogel may be inappropriate for our purposes; HUVECs did not migrate in the HIGH hydrogel (see below). Cell Behaviors on and in Gelatin-CMC Hydrogels. HUVECs were seeded onto the hydrogels to evaluate their initial cell attachment and subsequent behaviors in the microenvironment

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RESULTS AND DISCUSSION Synthesis of Gelatin-ADH and CMC−CHO. The in situ cross-linkable hydrogels composed of two biocompatible materials, gelatin and CMC, were synthesized for the rapid fabrication of vasculatures. Gelatin is an irreversibly hydrolyzed form of collagen and provides cell-binding motifs such as RGD residues.24,25 The modified gelatin displayed a newly formed amide carbonyl group that was detected using TNBS and Fourier transform infrared spectroscopy (FT-IR) measurements. The TNBS assay, which is based on a reaction with primary amines,26 revealed that ∼30% of the carboxy groups in the gelatin were substituted with hydrazide groups. A small peak for the amide carbonyl group appeared in the FT-IR difference spectrum at the corresponding wavenumbers (Figure S2b). CMC is a chemically versatile natural material with abundant hydroxyl and carboxyl groups, and its biocompatibility has been demonstrated in various in vivo applications, including the prevention of postoperative peritoneal adhesion.21 We predicted that CMC could provide physical strength to the hydrogel. The modification of CMC led to the formation of aldehyde groups through an oxidation reaction with sodium periodate (Figure 1a). The successful synthesis of CMC−CHO was confirmed by FT-IR measurements (Figure S2c) and 1H NMR analyses. The modification rate was ∼40% based on 1H NMR data. Characterization of Gelatin-CMC Hydrogels. Compared to physical gelation hydrogels such as collagen, chemical crosslinking hydrogels are more flexible, which offers an advantage for readily optimizing their characteristics such as gelation time, cross-linking density, and mechanical strength in accordance with their intended uses. These characteristics are closely related and are readily adjustable by modulating factors such as the concentrations of hydrogel materials, compositional ratios, and chemical modifications. In this study, hydrogels with three

Figure 2. Characterization of gelatin-CMC hydrogels. (a−d) Scanning electron microscope images of lyophilized gelation-CMC hydrogels with LOW (a), MID (b), and HIGH (c) concentrations and COL hydrogel (d). (e) Pore sizes quantified by image analysis of (a−d). Error bars indicate standard deviations from four independent experiments. Numerical variables were evaluated by one-way ANOVA, and *p < 0.05 was considered significant. D

DOI: 10.1021/acsbiomaterials.6b00203 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX

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Figure 3. Cytocompatibility of gelatin-CMC hydrogels. (a) Morphology of HUVECs on the hydrogels. (b) Changes in the number of cells on the hydrogels. (c) Viability of GFP-HUVECs exposed to gelatin-CMC hydrogels at 24 h of culture. Error bars indicate standard deviations calculated from six independent experiments.

In Vitro Angiogenesis Assay Using HUVEC Spheroids. We further examined network formation in the hydrogels using HUVEC spheroids. GFP-HUVECs seeded into a noncell-adhesive round-bottomed 96-well plate formed spheroids that were 150−200 μm in diameter (Figure 5a). The spheroids were then suspended in a gelatin-ADH solution at 25 spheroids/mL and subsequently mixed with a CMC−CHO solution using a doublebarreled syringe in a 35 mm culture dish. In LOW hydrogel, sprouting was observed at 3 days of culture, but the hydrogel degraded and collapsed by 5 days of culture (Figure 5b,c). The HIGH hydrogel displayed robust stability and maintained its gel form for at least 7 days of culture, but the cells showed little migration and sprouting from the spheroids over time (Figure 5b,c). The MID hydrogel maintained its gel form for at least 7 days of culture and allowed HUVECs to sprout from the spheroids (Figure 5b,c). The average length of sprouting reached ∼150 μm at 3 days in the MID hydrogel (Figure 5b,c), which is comparable to that observed in the COL hydrogel and in a previous report.22 Sprouting was stabilized and further extended to ∼250 μm at 7 days of culture (Figure 5c,d). On the basis of these results, the MID hydrogel was rapidly cross-linkable, sufficiently robust to hold its shape, and suitable for vascularization of endothelial cells. However, the sprouting structures had ruptured by 14 days of culture in both the MID and collagen hydrogels, suggesting that additional strategies such as coculture with pericytes and other types of vascular cells are necessary to further stabilize the vascular networks.11

(Figure 3a). After 24 h of culture, a larger number of cells was attached on the MID and COL hydrogels than on the LOW and HIGH hydrogels (Figure 3b). Particularly, the number of cells on the MID hydrogel was 2-fold higher than that on the HIGH hydrogel after 24 h of culture. In subsequent culture, attached cells grew vigorously on the MID and COL hydrogels, but slow and low growth were observed on the LOW and HIGH hydrogels, respectively. Since such significant differences were observed between these hydrogels, the two hydrogel components were ejected directly onto HUVECs adhered to a culture dish through a double-barreled syringe to evaluate the cytocompatibility of the hydrogels. Although there was no significant difference between LOW and COL hydrogels, cell viability decreased as hydrogel concentration increased (Figure 3c). This is likely because of the adverse effects of the CHO groups, which may have inhibited growth on the HIGH hydrogel (Figure 3b). The slow growth on the LOW hydrogel may have occurred not because of the CHO groups but was related to swelling and degradation, as the hydrogel had fragmented within 3 days in PBS (Figure S1). Although an ∼20% decrease in cell viability was observed in the MID hydrogel, cells actively grew and began to form network-like structures (Figure 3a,b). We further examined the network formation of HUVECs in the MID hydrogel (Figure 4). HUVECs represented network-like structures both on and in the MID hydrogel, which resembled the structures observed in Matrigel rather than those in the COL hydrogel.30 The results suggest that the MID hydrogel provides a better microenvironment for the engineering of vascularized tissues. E

DOI: 10.1021/acsbiomaterials.6b00203 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX

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Figure 4. Confocal microscopic images of GFP-HUVECs cultured on/in hydrogels for 48 h. (a−d) Cells on/in MID hydrogel. (e−h) Cells on/in COL hydrogel. The three-dimensional images (b,d,f,h) were reconstructed from the corresponding images (a,c,e,g).

Figure 5. Sprouting of HUVECs from spheroids embedded in hydrogels. (a) Scanning electron microscope image of a spheroid at 2 days of culture. (b) Average length of sprouting from a spheroid surface. Error bars represent standard deviations calculated from 4 independent experiments for each condition. (c) Comparisons of sprouting 3 days after seeding in the hydrogels. (d) Sprouting in MID hydrogel at 7 days of culture. (i) Confocal microscope image of GFP, (ii) 3D image of sprouting reconstructed (i) using image analysis software and (iii) from a confocal microscopic image of sprouting with rhodamine-phalloidin and DAPI staining.

hydrogel after applying −1.0 V vs the Ag/AgCl reference electrode for 5 min (Figure 6b). In contrast, more than 95% of the cells were electrochemically transferred from the oligopeptide-modified substrate onto the hydrogel (Figure 6a,b), which is consistent with our previous report using different hydrogels.31

Fabrication of Perfusable Vasculatures. HUVECs were seeded on an oligopeptide-modified gold substrate and then electrochemically transferred to a MID hydrogel. Without oligopeptide modification, nearly all cells remained on the substrate (Figure 6a), and less than 10% of cells were transferred to the F

DOI: 10.1021/acsbiomaterials.6b00203 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX

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Figure 6. Electrochemical cell transfer and fabrication of perfusable vascular-like structures. (a) Electrochemical cell transfer to the MID hydrogel from a gold surface modified with (Pep+) or without (Pep−) oligopeptides. (b) Number of cells electrochemically transferred with or without oligopeptide modification. (c,d) HUVECs adhered via an oligopeptide layer on a gold-coated needle. (e,f) HUVECs transferred from a needle to an internal surface of a microchannel in MID hydrogel. (g−j) HUVECs sprouted from a microchannel at 4 days of perfusion culture.

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CONCLUSIONS We demonstrated that by mixing two hydrogel component solutions, gelation could be achieved in less than 30 s through the formation of stable hydrazone bonds. HUVECs seeded on or encapsulated in the MID hydrogel migrated and sprouted in the following culture. A combination of a promptly cross-linking hydrogel and electrochemical cell detachment was used for the rapid fabrication of microchannels, in which the internal surface was lined with HUVECs. The transferred HUVECs migrated into the hydrogels and formed luminal structures. This approach can be used to engineer vascularized tissues and organs.

Using electrochemical cell transfer, we fabricated a perfusable vasculature in the MID hydrogel (Figure 1d). HUVECs were seeded onto an oligopeptide-modified rod and grew to completely cover the surface (Figure 6c,d). The rod was placed in the culture chamber, and the chamber was then filled with the MID hydrogel. By applying −1.0 V vs the Ag/AgCl reference electrode for 5 min, the cell layer was transferred to the internal surface of the microchannels in the hydrogel (Figure 6e,f). After the rod was extracted from the chamber, the culture medium was perfused through the HUVEC-enveloped microchannel. After 4 days of perfusion culture, the cells migrated and sprouted from the surface to the hydrogel (Figure 6g−i). Some of the cells showed a sprouting length of ∼200 μm (Figure 6j). In our previous study using collagen gels, HUVECs began sprouting at 4 days of perfusion culture and then bridged at 500 μm intervals between the two microchannels at 7 days.32 However, at least 30 min was required to gelate the collagen, which is critical when organ cells such as hepatocytes are encapsulated at high cell densities to engineer solid organs because oxygen shortage may pose tremendous stress on the cells. MID hydrogels gelated in