Construction of Injectable Double-Network Hydrogels for Cell Delivery

Article

Construction of injectable double-network hydrogels for cell delivery Yan Yan, Mengnan Li, Di Yang, Qian Wang, Fuxin Liang, Xiaozhong Qu, Dong Qiu, and Zhenzhong Yang Biomacromolecules, Just Accepted Manuscript • Publication Date (Web): 30 May 2017 Downloaded from http://pubs.acs.org on May 31, 2017

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Construction of injectable double-network hydrogels for cell delivery Yan Yana, Mengnan Lib, Di Yanga, Qian Wangb, Fuxin Liangb, Xiaozhong Qua*, Dong Qiu and Zhenzhong Yangb* a

College of Materials Science and Opto-Electronic Technology, University of Chinese Academy

of Sciences, Beijing 100049, China. b State Key Laboratory of Polymer Physics and Chemistry, Institute of Chemistry, Chinese Academy of Sciences, Beijing 100190, China. KEYWORDS. Injectable hydrogel, double-network, dynamic crosslinks, 3D cell encapsulation, tissue repair.

ABSTRACT. Herein we present a unique method of using dynamic crosslinks, i.e. dynamic covalent bonding and ionic interaction, for the construction of injectable double-network (DN) hydrogels, with the objective of cell delivery for cartilage repair. Glycol chitosan and dibenzaldhyde capped poly(ethylene oxide) formed the first network while calcium alginate formed the second one, and in the resultant DN hydrogel, either of the networks could be selectively removed. The moduli of the DN hydrogel were significantly improved compared to that of the parent single-network hydrogels and were tunable by changing the chemical components. In situ 3D cell encapsulation could be easily performed by mixing cell suspension

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to the polymer solutions and transferred through a syringe needle before sol-gel transition. Cell proliferation and mediated differentiation of mouse chondrogenic cells were achieved in the DN hydrogel extracellular matrix.

INTRODUCTION Injectable hydrogels are promising for application in drug delivery and tissue engineering because of their carrier property together with easy-operation perspective.1-3 It is known that hydrogel is one of ideal materials as extracellular matrix (ECM) for three-dimensional (3D) encapsulation of cells which could provide not only a highly hydrated tissue-like environment,4 but also the capacity of incorporation of bioactives, such as growth factors, to mimic the tissuespecific signals and thus influence the embryonic development, proliferation and even the differentiation of the cells.5 A 3D encapsulation of cells requires in situ crosslinking property of the gelation system, which is also a necessary ability of injectable gels. More importantly, for cell encapsulation the sol-gel transition should be managed to occur under mild condition to ensure the viability of the entrapped cells. Meanwhile, the injectability of the gelation system is expected for clinic application due to expected low invasive procedure and the ability to fill irregular shaped voids.6-8 To date, many efforts have been made on the construction of injectable in situ forming gels, including physical and chemical crosslinking methods.2 However, physical crosslinking relies on secondary forces like electrostatic interactions, stereo-complexation, hydrophobic interactions, and hydrogen bonding which frequently result in low stability and weak mechanical property of the gelation system.9-11 In contrast, chemical crosslinking regimes via prominent covalent bonding would introduce organic compounds for initiation and/or the


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trigger of UV light,12,13 which probably lead to uncertainty on safety issue or the handicap on injectability such as to deep tissue with the utility of photopolymerization.14 Recently, injectable hydrogels formed through dynamic interactions including dynamic covalent bonding have been developed.15-17 For example, gelation via Schiff base between polymers or macromolecules with amine and benzaldehyde groups can be obtained under physiological condition, i.e. pH 7.4 and 37 °C, without the addition and production of small compound, and the resultant gels exhibited high modulus similar to that of chemical crosslinked hydrogels.18,19 By the formation of benzoic-imine linkage, injectable glycol chitosan (GC) hydrogels showed a storage moduli (G’) of ~104 Pa by using di-benzaldehyde capped linear poly(ethylene oxide) derivatives (OHC-PEO-CHO) as a macromolecular crosslinker. Such hydrogels have been tested as an implant for the delivery of anti-tumor drugs and cells.20 However, like classic chemical crosslinked hydrogels, the dynamic covalent bonded GC gels is also brittle, which influences its application in tissue engineering applications that demand certain mechanical strength for the scaffold. For instance, human hyaluronic cartilage is a kind of hydrogel-like tissue containing more than 70% of water with chondrocytes dispersed in a matrix mainly formed by collagen type II (Col II) and glycosaminoglycan (GAG) including chondroitin sulfate and keratan sulfate.21 Such structure allows the tissue to be high load-bearing but is difficult to regenerate upon damage due to the lack of blood supply and innervation. Although hydrogel has been recognized as a potential candidate for cartilage repair, the material with capacity of cell encapsulation, injectability as well as desirable mechanical properties is rare.22 The formation of double-network (DN) is an efficient way to improve the mechanical property of hydrogels.23,24 DN hydrogel is a combination of rigid and ductile networks which significantly enhance the fracture energy against compressive or tensile stress.25 Nevertheless, since the rigid


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network serves as sacrificial frame dissipating the loading stress, it normally needs to be highly crosslinked by chemical bonds.26 This characteristic limited the formation of the second network synchronously, due to possible steric hindrances causing by the combination of two synchronous crosslinking processes.27 Although in situ forming DN hydrogels were developed recently,27-29 the construction of injectable DN gels is still challenging, especially for biomedical aims like cell delivery which requires mild gelation, and meanwhile needs fast transition sooner after the polymer solution flowed through the needle capillary in order to hold the integration of the implant. So far only a few examples on injectable DN and interpenetration (IPN) hydrogels were reported.30-32 Among them, to get a tough system, in situ chemical crosslinking, such as by Michael addition and thiol-disulfide exchange reaction, was exploited to build one of the polymer networks in the DN gel while ionic interaction was used to construct the second one.30,33 Nevertheless, these reactions require basic condition at pH 8.5 or take long duration up to 12 h for equilibrium. Therefore, more works on injectable DN hydrogels are necessary in accordance with the biomedical requirements. The major objective of this work is to obtain injectable DN hydrogels with dynamic covalent crosslinks to gain both cytocompatiblity and high mechanical property for the engineering of soft tissues like articular cartilage. We propose that the hydrogel formed via dynamic covalent bonding, e.g. the GC/OHC-PEO-CHO hydrogel (GC gel), could serve as the first network to incorporate a soft network that is crosslinked by weak secondary forces like electrostatic interaction. For instance, calcium alginate hydrogel (Alg gel) can be selected as the soft network as it has been extensively studied as a building block of DN and IPN hydrogels.34,35 Besides, the characteristics of alginate scaffolds for tissue engineering were also well-documented including the transplantation of chondrogenic cells to restore damaged cartilage in animal models.36,37 It


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was presumed that the ionically-crosslinked alginate network would provide an energy dissipation mechanism upon compressive strain, whilst the dynamic covalently bonded glycol chitosan network could allow for crack bridging and hence maintenance of mechanical integrity once the ionic crosslinks are broken. Besides, because both networks are formed by dynamic forces, in situ cell encapsulation and proliferation of cells in the DN hydrogel matrix could be favorably achieved. Herein we will show the synthesis, characterization, in vitro and in vivo performances of the GC/Alg DN hydrogels.

EXPERIMENTAL SECTION Materials. Poly(ethylene oxide) (PEO) (MW = 2 kDa), glycol chitosan (GC, MW ~250 kDa) and sodium alginate (MW ~150 kDa, M/G ratio ~1.6) were purchased from Sigma-Aldrich (St. Louis, US). Benzaldehyde capped poly(ethylene oxide) (CHO-PEO-CHO) and silica nanoparticles (average diameter 12 nm) were synthesized according to our previous report.18,19,38 RPMI1640 medium, DMEM medium, pancreatic enzymes, fetal bovine serum (FBS), penicillinstreptomycin solution and phosphate buffered saline (PBS) were purchased from Gibco (Grant Island, US). Organic solvents and other compounds were all obtained from Beijing Chemical Reagents Company (China), and were used as received. Formation of GC/OHC-PEO-CHO and calcium alginate double-network hydrogel. The typical method for preparing the double-network (DN) hydrogels was first to prepare two aqueous solutions containing GC and CaCl2, sodium alginate and OHC-PEO-CHO, respectively, at calculated concentrations, and then pour equal volume of the solutions in a mold to reach the desired concentrations the components, ranged from 0.1 to 5 wt%, followed by mixing under violent stirring for 1 min and then standing for 30 min at room temperature to allow gelation. For


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injectable samples, PBS solutions of GC and CaCl2, sodium alginate and OHC-PEO-CHO were made at designed concentrations. The solutions were loaded into two 2.5 mL syringes at equal volume which were then mounted to a dual syringe kit for an injection through a 24 G needle. Mechanical property test. Mechanical tests were performed using an INSTRON Model 5567 machine at room temperature in air. For compression property test, the samples were molded in cylinder shape of 14 mm in diameter (d) and 10 mm in height (h). The rate of compression was fixed at 100 N and 1 mm/min with a maximum compressive strain set at 90%. For loading and unloading test, the compressive strain was set at 35 % with compression rate of 5 mm/min. The compression stress (σ) was obtained by dividing the force (F) by the cross-sectional area (A = π×d2/4) and the compressive strain (ε) was obtained by dividing the deformed height (h’) by the original height (h). The experiments were conducted in triplicate. To test the interpenetration of the double-network, one the polymer networks was selectively moved from the prepared DN hydrogels. 3 pieces of DN gels were dipped into HCl (pH = 5) to wipe off GC/OHC-PEO-CHO gel, while other 3 pieces were dipped into 10% EDTA to chelate Ca2+ in calcium alginate network. Then the treated DN gels were immersed in and washed with water to remove free molecules for further mechanical property test. Rheology characterization. GC, sodium alginate, OHC-PEO-CHO and CaCl2 solutions were mixed together in a beaker with violent stirring for 30 s, and then the mixture was poured onto the bottom plate of stress-controlled rheometer (Thermo Haake MARS) equipped with parallel plates of 35 mm in diameter for time sweeping and frequency sweeping which were carried out at 37 °C with a gap of 1 mm to monitor the change of the elastic (G’) and loss (G’’) moduli at a shear strain of 1% which was pre-determined in the linear strain range. During the measurement, paraffin was used to cover the edges of the sample to avoid water evaporation.


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Scanning electron microscopy (SEM) and FTIR Characterizations. The morphology of hydrogel was observed using a Quanta-250 SEM. The samples were prepared by freeze-fracture method, i.e. immersing the hydrogel into liquid nitrogen, sliced up and sputter coated with gold for SEM observation. Infrared spectroscopy was recorded using a Bruker tensor 27 FT-IR spectrometer. Freeze-dried samples were pressed with KBr and scanned from 4000 to 400 cm-1. In vitro degradation studies. Degradation of the hydrogels was examined according to weight loss with time up to 4 weeks. Both DN and single-network hydrogels were prepared with water in cylinder shape and a volume of 2 mL. The hydrogels were weighted as W0 and then incubated with 30 mL of water at 37 °C. Weight loss of the hydrogels was monitored as a function of incubation time by replacing the incubation solution with equal value of water at a specified time interval. The incubation solutions collected at each time point were lyophilized and weighed (Wt). The weight remaining ratio was defined as 100%×(W0-ΣWt)/W0. The experiments were conducted in triplicate. Cytotoxicity. Cytotoxicity was assessed by CCK-8 assay on ATDC5 cell line. Hydrogels (0.2 mL) were immersed into 2 mL of DMEM culture medium at 37 °C for 72 h. The leach liquor was collected to incubate with ATDC5 cells. Medium solutions of GC, CHO-PEO-CHO solution and sodium alginate were used as controls. The cells were first seeded in 96-well plates at a density of 1 × 104 cell/well. After 24 h incubation, the culture media was removed by sample solutions, and the plates were incubated at 37 °C for 24 h. Then 100 µL of freshly prepared medium with 10% CCK-8 reagent solution was added to each well and incubated for 1 h before it was measured at the absorbance of 450 nm using a microplate reader (Thermo MULTISKAN MK3). Percentage cell viability was expressed relative to the negative control (untreated cells) and the positive control (10% CCK-8 in medium).


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In vitro 3D cell culture. Dispersion of ATDC5 cells in DMEM medium was generally mixed by sodium alginate solution to get desired polysaccharide concentration (8 or 12 wt%). 50 µL of the mixture was added to 50 µL of the medium solution of OHC-PEO-CHO (4 wt%) in 96-well plate, to which 100 µL of a solution containing GC (4 or 6 wt%) and CaCl2 (2 wt%) was pipetted, mildly stirred, and held to allow gelation to form a cylinder shaped gel with a diameter of ca. 6 mm and a height of ca. 7 mm. The final cell density was fixed at 2.5 × 105 cell/mL. DMEM medium supplemented with 10% fetal bovine serum (FBS), 1% penicillin/streptomycin (100 U/mL) were then added to the wells to immerse the hydrogel. The plate was maintained at 37 °C in 5% CO2, 95% humidiﬁed atmosphere. Cell proliferation was measured by CCK-8 assay at predicted time within 4 weeks. The optical density (OD) value for the number of seeded cells was set as 100%. The proliferation was then defined as the percentage of OD at the desired incubation time over that recorded from the amount of seeded cells. The cell condition was monitored using confocal laser scanning microscopy (CLSM). ATDC5 cell encapsulated DN hydrogels were prepared in 96-well plate. After incubated for predetermined days, the hydrogels were stained by 1 mL of acridine orange (AO, 2 µg/mL) and propidium iodide (PI, 3 µg/mL) for 20 min, followed by washing with PBS for three times. The samples were cut into slices and transferred to CLSM observation which was done using a FV 1000-IX81 confocal laser scanning microscope (Olympus, Japan) with an excitation wavelength of 488 nm and 559 nm for AO and PI, respectively. The emission of AO was set at 500 - 580 nm, while it set at 600 - 650 nm for PI. In vitro differentiation of ADTC5 in the hydrogels. To check the ability of chondrogenic differentiation for the ATDC5 cells in the hydrogel matrix, cell loaded single and doublenetwork hydrogels were prepared as described above, expect the medium was supplemented by


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10 µg/ml holo-transferrin, 3 × 10-8 M sodium selenite, 5% porcine bovine insulin, 50 µg/ml vitamin C.39 The cell loaded hydrogels were then incubated with the DMEM/F12 medium. After 1 - 4 weeks of incubation, the hydrogels were collected, fixed by paraformaldehyde, embedding in paraffin and cut into 4 µm sections for histopathology and immunohistochemistry stains under standard procedures. Subcutaneous injection of GC/Alg DN hydrogels. Male BALB/c mice (6 weeks of age, 20 ± 2 g) were purchased from Beijing HFK Bioscience Co., Ltd. (China). All animal experiments were performed in compliance with the relevant laws and institutional guidelines of the University of Chinese Academy of Sciences (UCAS) and approved by the Institutional Animal Care and Use Committee of UCAS. Polymer solutions of sodium alginate, OHC-PEO-CHO and GC, CaCl2 were prepared in DMEM medium, loaded to the syringes and mounted onto the dual syringe kit for DN hydrogel injection. Meanwhile, GC, OHC-PEO-CHO and sodium alginate, CaCl2 solution were also prepared respectively for the injection of single-network gels. Twenty seven mice were involved in the test which were divided into three groups for the three formulations. 0.2 mL of the polymer solution was injected subcutaneously into the right side of the dorsal region of each mouse at room temperature. At 2, 4 and 6 weeks post-administration, three animals in each group were sacrificed and the skin around the implantation site was carefully incised to isolate the skin tissue with residue of the implants. The tissue was immediately fixed in 10% formaldehyde for histopathological characterizations. Statistics analysis. Statistical significance was assessed using one-way ANOVA analysis. The difference was considered to be statistically significant if the probability value was less than 0.05 (p < 0.05).


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RESULTS AND DISCUSSION

Scheme 1. Illustration scheme of the formation of injectable GC/OHC-PEO-CHO and calcium alginate double-network hydrogel. Formation of GC/Alg double-network hydrogel. Four components were involved in the construction of the DN gel, i.e. glycol chitosan (GC), di-benzaldehyde functionalized poly(ethylene oxide) (OHC-PEO-CHO), sodium alginate and calcium chloride (CaCl2). GC was selected because of amine containing and good water solubility at neutral pH. The OHC-PEOCHO was synthesized according to our previous work,18,19 which could form benzoic-imine bond by Schiff’s reaction with GC to generate the first network. Sodium alginate is biocompatible as the chitosan derivative and poly(ethylene oxide), and was extensively studied for making hydrogel with CaCl2.40 The formation of GC/OHC-PEO-CHO (GC gel) and calcium alginate (Alg gel) double-network hydrogel (GC/Alg DN gel) was achieved by mixing aqueous solutions containing GC, CaCl2 and Alg, OHC-PEO-CHO respectively (Scheme 1). The utility of dynamic covalent bonding, i.e. benzoic-imine, and supramolecular interaction, i.e. ionic force, enabled desirable processing property of the DN gel. The injectability of the system was demonstrated in


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Figure 1, where it can be seen that the injection was easily performed using a dual syringe kit and the hydrogel was formed sooner after the polymer solutions were extruded out of the needle capillary, evidenced by the construction of three dimensional architecture. Moreover, the dynamic nature of the benzoic-mine bonding and electronic interaction endow a possible adhesion process of the DN gel pieces (Figure 1).

Figure 1. Demonstration of injectability of GC/Alg DN hydrogel (containing 1 wt % of GC, OHC-PEO-CHO, sodium alginate and CaCl2 respectively) through a 24 G syringe needle (up row), and the adhesion procedure of DN hydrogel pieces after the interface was activated using buffer of pH 6 (bottom row). The sol-gel transition was monitored using rheometer, and the results are shown in Figure 2a. It is found that the gelation time of DN gel depends on the polymer concentrations, which determines the gelation kinetics of the polymer networks. Due to different crosslinking mechanism, the gelatin time of calcium alginate is faster than that of the GC gel at similar polymer concentration under the experimental condition. For example, at an alginate concentration of 0.5 wt%, the sol-gel transition was around 2 mins after mixed the polymer with 0.1 wt% of CaCl2, as viewed at the overlap point of storage modulus (G’) and loss modulus (G”) curves (Figure 2a). Meanwhile, the gelation of GC gel containing 0.5 wt% of GC and 0.5 wt% of


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OHC-PEO-CHO was ca. 4 mins. Based on these components, the double-network formation required a period of ca. 3.3 mins. Upon an increasing of GC and OHC-PEO-CHO concentration to 1 wt%, the gelation time of the resultant DN gel decreased to 80 s, closed to that of the neat GC gel. The results indicated that sol-gel transition of DN gel was mainly driven by the network which had the relative faster gelation kinetics. Further evidences were given by increasing the alginate concentration to 1 wt% or higher, where the calcium alginate achieved a much rapid gelation within 1 min of the sample preparation time, so that no cross of G’ and G” can be seen in the time-sweeping profiles (Figure 2a). As a result, the G’ was larger than the G” at the beginning of the test for the DN gel as well, even though the gelation of the GC network was much lower. Gelation kinetics is key factor for the process of injectable hydrogels. Faster sol-gel transition rate normally required harder pressure for injection in order to avoid the gelation in the needle capillary which could cause the stuck of the materials. In contrast, much lower gelation would leave liquid phase in the implantation site, once used for biomedical aims, resulting in the loss of gelator and cargoes via diffusion. Considering that there are two gelation processes, suitable gelation kinetics is critical to the formation of DN gel.26,41 In this regime, we found that the formation of GC/Alg DN hydrogels could be achieved by adjusting the concentration of reagents, hence tuning the crosslinking kinetics of the two networks, and therefore getting a complete interpenetration rather than phase separation of the two polymer networks during the mixing process,42 which was normally in a time scale of serval minutes like shown in Figure 2a.


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Figure 2. (a) Time sweeping of storage modulus (G’) and loss modulus (G”) to monitor the solgel transition of GC/OHC-PEO-CHO (GC), calcium alginate (Alg) and the GC/Alg doublenetwork (DN) regimes with different polymer and CaCl2 concentrations at 37 °C. (b) Frequency sweeping of GC, Alg and DN hydrogels with different components at 37 °C. The numbers in each figure, from left to right, represent the weight concentration (wt%) of GC, OHC-PEO-CHO and/or sodium alginate and CaCl2 for the preparation of samples. The moduli of the DN gels are also concentration dependent. Higher G’ was obtained for the hydrogels with higher polymer concentration (Figure 2b). The G’ values of the three tested DN gels are significantly higher than that of the corresponding parent single network gels, i.e. the GC gels and the Alg gels. Besides, it is also noted that the shear frequency caused limited change of the moduli for both the single and the double-network gels (Figure 2b), inferring the elastic dominating property of the hydrogels. SEM images confirm the pore formation in the freezedried DN gel upon the sublimation of ice templates, similar to the GC and Alg gels (Figure 3). Meanwhile, FTIR spectra reveal the appearance of stretching vibration of carboxyl group at 1703


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cm-1 from the DN sample, apart from the carbonyl peak at 1630 cm-1 (Figure 3), implying the decrease of coupled alginate segments by calcium ions in the DN gel matrix when compared to that in the Alg gel, which is attributed to the penetration of polymer chains in the two networks by which the GC and PEO chains hampered the approach of partial alginate blocks. The combination of the two single networks resulted in condensed polymer network, stronger interaction and more entanglements in the DN gels and thus led to increase of the elastic modulus.

Figure 3. SEM images and FTIR spectra of freeze-dried GC, Alg and GC/Alg DN gels. Scale bar 100 µm. Mechanical property of DN hydrogel. The mechanical property of the hydrogels was then tested. Under compressive loading, it can be seen that the GC hydrogel, containing 3 wt% of GC and 1 wt% of OHC-PEO-CHO, exhibited a brittle rupture, while the Alg gel, containing 3 wt% of alginate and 1 wt% of CaCl2, exhibited ductile failure which were not recoverable upon the release of the stress (Figure 4). In contrast, by the combination of the two networks, the resultant DN gel, named as DN3131 (rheological performance see Figure S1), preserved the integrality


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against a compressive strain of 90% (Figure 4), although the stress-strain curve still shows a failure point at ca. 50% strain, similar to that of the GC gel (Figure 5a). This is because of the synergistic response of the two asymmetric networks, where the Alg network held the GC network while it was resisting an elastic deformation. The failure points indicated an internal fracture of the GC network. However, with strain less than the failure point, the stress of the DN hydrogel harvested significant increase of the compressive stress than the parent single-network hydrogels. A fracture energy of 410 J/m2 was obtained from the DN gel (Figure 5b), which is in the range of cartilage strength, i.e. 102-103 J/m2,43 and more than 10 times larger than that of GC gel, inferring that Alg network could dissipate the compression. Being a rigid chain, i.e. anionic polysaccharide, the deformation of the alginate macromolecules requires stronger force when compared to soft chains like polyamide or poly(acrylic acid) (PAA). The slip of alginate chains was locked by the calcium ions and further protected by the GC frame.44 Such structure resulted in combinational effect on the resistibility against compression, and led to the increase of modulus without the elongation of failure strain. Meanwhile, the DN network showed antifatigue feature in the loading-unloading tests at a given strain of 35% (Figure 5c). Hysteresis loops were seen to indicate the relaxation process of the polymer chains. The maximum stress of the hydrogel as well as the loops show no obvious drop within the test of 100 cycles. Tardy attenuation was expected due to possible damage in the networks which would not be avoided during the press loading. However, with dynamic cross linked strategy, the break at the cross linked points could be healed in both the GC and the Alg networks. Thus the DN hydrogel had only a little deformation with the compressing-decompressing cycles, but without influencing the compression modulus (Figure 5c).


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Figure 4. Performance of GC, Alg and GC/Alg DN gels (containing 3 wt % of GC, 1 wt % of OHC-PEO-CHO, 3 wt % of alginate and 1 wt % of CaCl2) against compressive strain where different modes of failure is seen. It was also shown that the stress of the DN gel could be adjusted by the chemical composition. An orthogonal investigation showed that increasing level of crosslinking and the alginate concentration improved the stress of the DN gel (Figure S2). And more content of alginate, i.e. the soft network, could prolong the failure strain. However, the largest failure strain of the DN gel was around 50 % in this regime possibly because of the rigid feature of the polysaccharides, especially for that forming the physical network, i.e. the calcium alginate. The tensile property confirm the rigid broken of the DN gel, with no yield point was observed, although the stress and break strain had been significantly improved by the formation of double network (Figure S3). Using soft polymer chain to make the second network was proven to be efficient to construct more ductile DN gel with the GC network. As an example, the GC/PAA DN gel, in which the PAA was crosslinked by iron ions, was made to demonstrate a compressive stress-strain curve without showing obvious failure point at a strain value larger than 80% (Figure S4a and b). On the other hand, more toughness DN gel could be gained by the addition of inorganic


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nanoparticles, e.g. SiO2 nanoparticles, which contained hydroxyl surface groups to further prohibit the movement of GC chains via the formation of hydrogen bonding,38 which endowed an enhancement of both compressive stress and the failure point (Figure S4c), and an increase of fracture energy by 90% than the parent GC/Alg DN gel (Figure 5b). These also confirmed the generalization of combining dynamic covalent bonded network with ionic crosslinked network to be efficient for making double-network hydrogel with injectability.

Figure 5. (a) Compressive stress-strain relationships of GC, Alg and GC/Alg DN gels. (b) Fracture energy and fracture strain of GC, Alg, DN gels and the DN gel containing 1 wt% of SiO2 nanoparticles (SiNP, diameter 12 nm). (c) Compressive stress-strain relationship of DN gel upon loading-unloading cycles. (d) Comparison of compressive stress-strain relationships of neat GC and Alg gels to the residue DN gels after the Alg or the GC network was removed by immersing in 10% EDTA solution or being treated at acidic pH of 5. The concentration of GC, OHC-PEO-CHO, alginate and CaCl2 for formation the hydrogels shown this figure was 3 wt%, 1 wt%, 3 wt% and 1 wt% respectively.


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Because of using impermanent crosslinks, the GC/Alg DN gel showed unique property that allowed the movement of one particular network from the DN gel. Figure 5d shows that after being treated by HCl solution at pH 5 or in 10% EDTA solution to break the GC or Alg network, the stress-strain curves of the leaving hydrogels are close to that of the neat Alg and GC gels, owing to the hydrolysis of the benzoic-imine bond or the dissociation of calcium complex in the DN gel and hence the removal of the free macromolecules via diffusion. The results indicated the sensitivity of the DN gel to acidic condition and ion chelators, and therefore the mechanical property of the DN gel can be tuned by the environmental parameters. Besides, it also confirmed the interpenetration structure of the two polymer networks within the GC/Alg DN gel. In vitro 3D cell culture in the DN hydrogel. DN hydrogels with GC, OHC-PEO-CHO, alginate and calcium concentration of 3 wt%, 1 wt%, 3 wt% and 1 wt%, i.e. DN3131, and 2 wt%, 1 wt%, 2 wt% and 1 wt%, i.e. DN2121, were used for encapsulating cells. The toxicity of the DN hydrogel was first examined. From Figure S5, it can be seen that the leach liquor from the DN3131 hydrogel after 72 h extraction reached a cell viability of larger than 80%, an indication of non-toxicity of the material. The in vitro degradation of the DN gels was then tested. Within the experimental duration of 4 weeks, the DN3131 hydrogel was monitored to have ca. 30 % of weight loss, whilst the 80% and 90% for the parent Alg and GC gels (Figure 6a). It is known that the main reason causing the degradation of GC gel in buffer is the hydrolysis of the imine linkage and then the diffusion of the polymer chains to the surrounding aqueous phase.18 Similarly, the dissociation of the calcium complex in the Alg gel results in the slimming of the gel in diluted condition.45 Compared to them, the DN gel exhibited much slower degradation rate due to the more compact networks, which reduced the diffusion rate of the polymer chains in case the crosslink points were broken by solvent molecules, and hence the


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chains would be able to re-crosslink owning to the reversibility of imine-bonding or calcium complexation. Figure 6 also shows that the degradation kinetics can be changed by varying the polymer concentration. Lower polymer content resulted in faster degradation in vitro.

Figure 6. (a) In vitro degradation curves, i.e. the weight loss of GC gel (containing 3 wt% of GC and 1 wt% of OHC-PEO-CHO), Alg gel (containing 3 wt% of alginate and 1 wt% of calcium) and the DN gels as function of time. (b) Optical image of ATDC5 cell loaded GC gel, Alg gel and DN3131 gel prepared in medium after being cultured for 1 and 21 days. (c) CLSM image of ATDC5 cells in the DN2121 gel for different time. Scale bar 100 µm. (d) Quantitative proliferation of ATDC5 cells in GC, Alg and DN gels. The cell viability without hydrogel was used as reference. n/a: No data was got for the GC gel and Alg gel on day 28 because of the hydrogels broke into pieces. * Significantly different, p < 0.05.


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The cell encapsulation and proliferation were tested using ATDC5, a mouse chondrogenic cell line. The morphology of the cells in the formed DN hydrogel was observed using CLSM. The round shape of the cells was preserved and the cells were uniformly distributed in three dimension in the gel matrix by the in situ encapsulation (Figure 6c). Up to 21 days, the CLSM images confirmed the survival of majority of the cells by AO/PI staining. Besides, the images also reveal the increase of cell population and the formation of cell clusters, i.e. multicellular spheroids, with the prolongation of the incubation time (Figure 6c). Such phenomenon again inferred the favorable condition of the DN hydrogel as an ECM of the cells. The cell proliferation was quantitatively assessed by CCK-8 assay. Figure 6d showed that number of the ATDC5 cells in the DN gel increased by 6 and 10 times in an experimental period of 28 days in the DN3131 and DN2121 gels respectively. The results were reasonable because gel with looser density of the network provided more vacancies for faster proliferation of the cells. It also shows that the double-network did not seriously affect the cell proliferation in comparison with that in the single network hydrogels. The DN2121, with totally 5 wt% of polymer content, even gained faster cell growth than that of the Alg gel after 21 days of incubation (Figure 6d). Although the DN3131 had slower cell proliferation in the initial period, the difference became less obvious after 2 weeks of culture, attributed to the degradation of the matrix and also possible to the break of the dynamic crosslink points. Another point to be mentioned is that the measurement on cell proliferation in the GC and Alg gels was influenced by the degradation kinetics of the polymer network. A demonstration is shown in Figure 6b on the serious size change of the cell loaded GC and Alg hydrogels after 3 weeks’ incubation, which is consistent with the in vitro degradation results (Figure 6a). Faster weight loss of the singlenetwork gels could cause the leakage of cells, especially for the GC gel, which had led to more


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rapid increase of cell number on day 21 (Figure 6d), because of the proliferation of the released cells on the surface of well-plate. Furthermore after 4 weeks of incubation, the integration of both GC and Alg hydrogels could not be held and thus the proliferation in the gel matrix was unable to be calculated. In contrast, the lower level of weight loss from the DN gels will benefit their utility for long-term cell culture either in vitro or in vivo, which would be helpful for timeconsuming tissue engineering like cartilage tissue engineering.46

Figure 7. (a) Alcian blue staining of ATDC5 loaded GC gel (containing 2 wt% of GC and 1 wt% of OHC-PEO-CHO), Alg gel (containing 2 wt% of alginate and 1 wt% of calcium) and DN2121 gel after being in vitro cultured for 1 and 2 weeks. Scale bar 100 µm. (b) Immunohistochemical staining of collagen type II (Col II) on the cell loaded hydrogels incubated for 1, 2 and 4 (inset) weeks. Scale bar 100 µm. The hydrogels provide not only ECM for 3D cell proliferation, but also a carrier of bioactives to mediate the growth of the encapsulated cells. With loaded holo-transferrin, sodium selenite,


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bovine insulin and vitamin C,39 in vitro chondrogenic differentiation of ATDC5 cell in the DN gel as well as the parent single-network gels was investigated by histopathology and immunohistochemistry

characterizations.

Alcian

blue

has

been

used

to

stain

the

glycosaminoglycans (GAG) produced from the chondrogenesis of vary kind of cells in chitosan and alginate based scaffolds.47,48 Within 2 weeks, it shows that the ATDC5 loaded Alg and the DN gels display deeper blue color, indicating that the alginate may favor the chondrogenic differentiation of the cells (Figure 7a), probably due to the similarity of the alginate segments with the GAGs presented in articular cartilage.49 Meanwhile, the production of GAG was also revealed in the GC gel (Figure 7a). The immunohistochemical staining confirmed the synthesis of collagen type II (Col II) during the incubation of the ATDC5 cells in both the single-network and the double-network hydrogels in 2 weeks (Figure 7b). With well-preserved integrity, increased number of positive cells for Col II deposition could be observed from the immunohistochemical section of the DN gel after incubated in vitro for 4 weeks (Figure 7b). The results again infer favorable property of the DN gel as a scaffold for 3D encapsulation and the ability for further mediate the differentiation of chondrogenic cells in the matrix. Subcutaneous injection of GC/Alg DN hydrogels. Subcutaneous implantation of the GC/Alg DN hydrogel as well as GC and Alg gels was achieved on mouse model, by an injection using a dual syringe kit (Figure 8a). The H&E staining of the isolated tissues observed serious shrinkage of the Alg gel within 4 weeks in vivo, while the GC implant was nearly completely disappeared at 2 weeks post-implantation (Figure 8d). In contrast, the in vivo degradation of the DN gel was much slower. The residue of the implanted DN gel could be clearly found in the animals for up to 6 weeks (Figure 8b and c), which is in agreement with the in vitro data. Besides, the histopathological staining revealed the biocompatibility of the implants, in particular for the DN


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gel, which caused limit immunological response of the surrounding skin (Figure 8d). The boarder between the implants and the dermis is clear, with no damage being found in the periimplant tissue. Only slight inflammations can be observed after the hydrogels were injected for 2 weeks, a typical period for the animals to convalesce the macrophage-mediated foreign body reaction.50 And at this time point, the distribution of monocytes was already less obvious although partial of the implant, i.e. in the DN gels, had been infiltrated probably by macrophages. Despite the GC gel shows faster degradation in vivo, it was known that chitosan derivatives have anti-bacteria and haemostatic potentials in implanted devices,51 which may contribute the reduction of inflammation level for the GC/Alg DN gel. After 4 weeks, organization can be seen in the implantation site of the DN gel (Figure 8d), owning to the general degradation of the hydrogel. The H&E section demonstrated the presence of haematopoietic foci in the hydrogel matrix near its interface to the dermis, which is a phenomenon normally accompanied by the blood vessel infiltration.52 Cartilage regeneration usually takes serval months.53,54 During this period, the scaffold should exhibit both mechanical resistance and gradual degradation to allow the generation of new tissue while affording the external stress to avoid the deformation filling area. Herein, by forming double network, the GC/Alg DN hydrogel displayed prolonged degradation duration, enhanced compressive modulus compared to the parent single-network gels, and meanwhile achieved in situ cell encapsulation and injectability for better operability in the future application. While future study is warranted for a systematic investigation on in vivo performance of the DN gels, a test of injecting ATDC5 cell loaded DN gel subcutaneously to BALB/c mice has observed the production of GAG in the implantation site and thus infers the potential of the DN hydrogel as a scaffold for cartilaginous tissue formation (Figure S6).


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(d)

Figure 8. Photograph of injection of the DN3131 gel subcutaneously to an anaesthetized BALB/c mouse (a) and the skin tissue with the DN3131 implant isolated after 4 (b) and 6 (c) weeks of implantation. Scale bar 5 mm. (d) Histological images (H&E staining) of implantation site and surrounding tissue after GC gel (containing 3 wt% of GC and 1 wt% of OHC-PEOCHO), Alg gel (containing 3 wt% of alginate and 1 wt% of calcium) and DN3131 gel were injected subcutaneously and harvested at 2, 4, and 6 weeks. To each sample, the scale bar is 200 µm (left column) and 50 µm (right column) respectively. CONCLUSIONS Double-network hydrogels with injectability under physiological condition were obtained by the combination of ionic crosslink with dynamic covalent bonding crosslink and using


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polysaccharides, i.e. GC and alginate, and poly(ethylene oxide) as building blocks. The GC/Alg DN gels showed much improved stress against compressive strain in comparison with the parent single-network hydrogels. Both the mechanical strength and degradation duration of the DN gels were tunable by changing the chemical composition. And the mild process procedure allowed in situ encapsulation, proliferation and controlled differentiation of mouse chondrogenic cells in the hydrogel matrix. Both in vitro and in vivo tests proved the biocompatibility of the GC/Alg DN gels, such as showing no tissue damage after the gels were injected subcutaneously into BALB/c mice. This work provides a facile way to provide injectable DN hydrogels with desired mechanical property as a scaffold for treating cartilage defects.

ASSOCIATED CONTENT Supporting Information. G’ and G” curves of DN3131 hydrogel, compressive stress-strain curves of different DN hydrogels, tensile stress-strain performance of GC, Alg and DN hydrogels, compressive stress-strain curves of GC/poly(acrylic acid) DN hydrogel, cytotoxicity of the leach liquor from different hydrogels and histopathological sections of isolated tissue with cell loaded hydrogel implants AUTHOR INFORMATION Corresponding Author * Email: [email protected] (X. Q.), [email protected] (Z. Y.). Author Contributions


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The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. Notes The authors declare no competing financial interest. ACKNOWLEDGMENT This work is financially supported by the National Natural Science Foundation of China (51473169, 21274155, 51233007). REFERENCES (1)

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For Table of Contents Use Only

Construction of Injectable Double-Network Hydrogels for Cell Delivery Yan Yana, Mengnan Lib, Di Yanga, Qian Wangb, Fuxin Liangb, Xiaozhong Qua*, Dong Qiu and Zhenzhong Yangb* a

College of Materials Science and Opto-Electronic Technology, University of Chinese Academy of Sciences, Beijing 100049, China. b State Key Laboratory of Polymer Physics and Chemistry, Institute of Chemistry, Chinese Academy of Sciences, Beijing 100190, China. * Corresponding authors. Email: [email protected] (X.Q.) and [email protected] (Z.Y.)


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Construction of Injectable Double-Network Hydrogels for Cell Delivery

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