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Microporous methacrylated glycol chitosan-montmorillonite nanocomposite hydrogel for bone tissue engineering

Microporous methacrylated glycol chitosan-montmorillonite nanocomposite hydrogel for bone tissue... ARTICLE https://doi.org/10.1038/s41467-019-11511-3 OPEN Microporous methacrylated glycol chitosan- montmorillonite nanocomposite hydrogel for bone tissue engineering 1,2,5 3,5 3 2,3 4 2,3 Zhong-Kai Cui , Soyon Kim , Jessalyn J. Baljon , Benjamin M. Wu , Tara Aghaloo & Min Lee Injectable hydrogels can fill irregular defects and promote in situ tissue regrowth and regeneration. The ability of directing stem cell differentiation in a three-dimensional micro- environment for bone regeneration remains a challenge. In this study, we successfully nanoengineer an interconnected microporous networked photocrosslinkable chitosan in situ- forming hydrogel by introducing two-dimensional nanoclay particles with intercalation chemistry. The presence of the nanosilicates increases the Young’s modulus and stalls the degradation rate of the resulting hydrogels. We demonstrate that the reinforced hydrogels promote the proliferation as well as the attachment and induced the differentiation of encapsulated mesenchymal stem cells in vitro. Furthermore, we explore the effects of nanoengineered hydrogels in vivo with the critical-sized mouse calvarial defect model. Our results confirm that chitosan-montmorillonite hydrogels are able to recruit native cells and promote calvarial healing without delivery of additional therapeutic agents or stem cells, indicating their tissue engineering potential. 1 2 Department of Cell Biology, School of Basic Medical Sciences, Southern Medical University, Guangzhou, Guangdong 510515, China. Division of Advanced Prosthodontics, University of California Los Angeles, 10833 Le Conte Avenue, Los Angeles, CA 90095, USA. Department of Bioengineering, University of California Los Angeles, 420 Westwood Plaza, Los Angeles, CA 90095, USA. Division of Diagnostic and Surgical Sciences, University of California Los Angeles, 10833 Le Conte Avenue, Los Angeles, CA 90095, USA. These authors contributed equally: Zhong-Kai Cui, Soyon Kim. Correspondence and requests for materials should be addressed to Z.-K.C. (email: zhongkaicui@smu.edu.cn) or to M.L. (email: leemin@ucla.edu) NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications 1 1234567890():,; ARTICLE NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 issue engineering exploits a combination of cells, engi- include any live cells during fabrication of those scaffolds with neering and materials methods, along with proper bio- the conventional methods to create porous structure, such as Tchemical and physiochemical factors to improve or replace freeze-drying, porogen leaching. After the establishment of the biological tissues, including the skin, cartilage, bladder, blood scaffolds, cells were loaded leading to nonuniform distribution. vessels, and bone . The decisive contribution of the tissue engi- Considering this, MMT is of interest for the formation of porous neering matrix in ensuring the regeneration potential of stem and structure, and to be investigated further in vivo for TERM as an 2–4 progenitor cells has been emphasized increasingly . Progress in application. cell selection, cell culture, and new material formulations has led Hydrogels derived from natural products are an appealing to the development of more effective therapies for tissue engi- three-dimensional biomaterial for tissue engineering. Compared neering and regenerative medicine (TERM). Various materials with hard scaffolds that require pre-shaping, soft hydrogels are have been explored for regenerative applications, including injectable and can fill any irregular shaped defects in a minimally 5,6 naturally occurring products and synthetic materials . Generally invasive manner. Recently, we have developed a photoinducible speaking, the intrinsic properties of naturally derived materials, hydrogel system of methacrylated glycol chitosan (MeGC) using such as collagens, are definitely attractive; however, the com- riboflavin as a photoinitiator . MeGC hydrogels supported plexities of purification, immunogenicity, mechanical properties, proliferation and extracellular matrix deposition of encapsulated 25,26 and pathogen transmission significantly limit their application. In mesenchymal stem cells ; however, the hydrogel itself showed light of this, greater control over material properties and tissue minimal bone-forming ability without osteogenic factors, bioac- responses has garnered more attention based on designable tive molecules, or encapsulated cells. One of the possible reasons modern material science. is lacking of porous structure in the MeGC hydrogel. Porosity is Biomaterials are having great impact on the medical treatments reported to be necessary for new tissue formation as it allows the in helping to solve clinical problems. In the field of medicine, cells to migrate, infiltrate, and proliferate in a 3D environment, as drug-eluting stent coated with polymers and controlled drug well as for the vascularization, differentiation, and mass trans- release systems constituted with biomaterials have been saving port . Therefore, it is of interest to develop a hydrogel system 7,8 hundreds of thousands of lives each year . In the field of TERM, that could recruit native cells and facilitate bone formation with a the combination of biomaterial scaffold and certain cells is pos- microporous interconnected structure. sible to replace biological tissues . In the field of medical devices, In this study, we introduce MMT to the photopolymerizable biomaterials are also playing a significantly important role as the MeGC hydrogel system to fabricate an injectable highly osteo- core components in surgical sutures, bioadhesives, and dental conductive in situ-forming biomaterial for bone regeneration. We implants . Clays and clay minerals are emerging materials for hypothesize that MMT can not only improve the microstructure biomaterial design to provide new strategies for TERM. The usage but also enhance the mechanical properties of cured MeGC of inorganic layered nanomaterials for medical purpose dates hydrogel. Therefore, we started with optimizing the proportion of back to ancient time, such as for wound healing and hemorrhage MMT in hydrogels and investigated the osteogenesis of encap- inhibition . Nowadays, clays and clay minerals are being applied sulated mesenchymal stem cells in vitro. We further evaluated the in pharmaceuticals as active ingredients or excipients, and in ability of the nanocomposite hydrogels to recruit native cells and cosmetics as creams, powders, and emulsions . The interactions improve bone formation in a critical-sized mouse calvarial defect between clay nanoparticles and drugs as well as other biological model. The treatment of bone defects remains one of the largest molecules have been well investigated and therefore exploited for challenges in musculoskeletal TERM, thereby our developed 11,12 controlled delivery , and moreover, the addition of clay into nanosilicate-loaded MeGC hydrogel may represent a new mate- polymers enhances the mechanical properties owing to the for- rial design in the broadly interesting area of growth factor free 13,14 mation of nanocomposites . and cell-free strategies. This bioactive nanocomposite hydrogel Montmorillonite (MMT) is a major component of Bentonite, can provide an effective treatment for bone defects. which is already approved by the FDA as an additive in various medicinal products . Studies on the acute and chronic toxicities of MMT have confirmed the absence of any negative effects, Results 15,16 even on embryos of pregnant Sprague–Dawley rats . MMT Characterization of MeGC-MMT nanocomposite hydrogels. is a layered silicate [(Na,Ca) (Al,Mg) Si O (OH) � nH O], Nanocomposite hydrogels including various amount of MMT 0.33 2 4 10 2 2 belonging to the smectite group of minerals, with high (0.5-4% w/v) were prepared as illustrated in Fig. 1a. We already 2 −1 specific surface area (up to 600 m g ) and aspect ratio. The tested higher amount of MMT up to 10% in MeGC hydrogels. It repeating structural unit of MMT consists of one alumina octa- became too viscous to mix with poor handling, as the MMT ratio hedral sheet sandwiched in between two silicon tetrahedral layers. increased over 4%. Therefore, we have selected the MMT ratio The overall surface charge is light negative because of the dom- between 0.5 and 4% for the investigation. The electrostatic ination of the oxide anions, which facilitates mixing with cationic interactions between the heterogeneously distributed charges of agents. The MMT particles are typically in a plate-shape, ren- discotic MMT (overall negative charge) and MeGC hydrogel dering ~1 nm in thickness and 0.2–2 μm in diameter. The bio- matrices (positive charge) enhanced the Young’s modulus with an compatibility, availability, and feasibility of this particular mineral increasing amount of MMT in the nanocomposite hydrogels from has garnered significant attention in recent years. Extensive 10 kPa to over 60 kPa (Fig. 1b). Equilibrium water content (EWC) research has been carried out to investigate for the purpose of inversely represents the cross-linking density and the mechanical 11,17,18 28 drug and gene delivery with natural or modified MMT .In properties of a hydrogel . Introduction of more than 0.5% MMT addition, several reports have demonstrated that fabrication of in MeGC hydrogels significantly decreased the EWC of corre- scaffolds by introduction of MMT in natural biomaterials, sponding nanocomposite hydrogels (Fig. 1c), indicating enhanced 19,20 21 22 23,24 including gelatin , collagen , silk , and chitosan , mechanical properties. The dry weights of the MeGC-MMT improved cell-scaffold interactions, cell proliferation, and nanocomposite hydrogels were recorded for 42 days (Fig. 1d). As enhanced cell differentiation. Although all those hard scaffolds observed in the degradation profiles, the degradation rate of the were reported with porous structure under high vacuum and dry 1.5% and 3.0% MMT incorporated MeGC hydrogels was sig- conditions, besides those experiments were limited only in vitro, nificantly decreased compared with the control MeGC hydrogels which are not ideal to be applied in TERM, as it is not possible to by ~40%. TGA degradation profiles confirmed the presence of 2 NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 ARTICLE ab Visible blue light MMT Riboflavin MeGC MeGC-MMT 0 0.5 1 1.5 2 3 4 MMT weight (%) cd 90 0 0 7 14 21 28 35 42 0 0.5 1 1.5 2 34 MMT weight (%) Time (Day) Fig. 1 Characterization of photocross-linked MeGC-MMT nanocomposite hydrogels. a Schematic illustration of MeGC-MMT hydrogel cured by visible blue light cross-linking with riboflavin as the photoinitiator in the presence of MMT. b Young’s modulus of MeGC-MMT nanocomposite hydrogels with various amount of MMT. c Equilibrium water content of MeGC-MMT nanocomposite hydrogels. d Degradation profiles of MeGC-MMT nanocomposite hydrogels in the presence of 0% (square), 1.5% (circle), and 3% (triangle) MMT in PBS at 37 °C for 42 days. Error bars indicate standard deviation (n = 5). *p < 0.05, **p < 0.01, and ***p < 0.001 compared with the control MeGC group with 0% MMT (ANOVA followed by Tukey’s post hoc test) MMT enhanced the thermal stability of MeGC hydrogels (Sup- confirm the interconnected microporous structure in the 1.5 and plementary Fig. 1). 3% MMT groups, and the highest number of cells was observed in The microstructure of MeGC and MeGC-MMT hydrogels the 1.5% MMT group, while MeGC hydrogel exhibited a smooth observed by SEM (Fig. 2) confirms the formation of a surface (Fig. 4). In addition, round-shaped cells were observed in microporous and interconnected network when MMT exceeds the MeGC image, while spread cell morphology was seen in the 1.5%, while a smooth continuous surface was acquired for the MMT containing nanocomposite hydrogels. Taken together, the control MeGC group, and a coarse continuous surface appeared 1.5% MeGC-MMT nanocomposite hydrogel group presented for 0.5 and 1% MMT groups. Porosity was quantified as an index the best support for cell proliferation. of surface area occupied by the pores in the SEM images. The pore sizes were around 115 ± 40 μm for the 1.5% MMT group. Bioactivity of MeGC-MMT nanocomposite hydrogels. Differ- With the increasing MMT content, the pore sizes reached the entiation of BMSCs to osteoblasts typically can be evaluated by maximum of 150 ± 50 μm for the 3.0% MMT group. The the expression of early markers, such as ALP and the ultimate involvement of MMT thoroughly changed the microstructure of calcium deposition . The BMSCs encapsulated in the MeGC- MeGC hydrogels. In addition, MMT was well distributed in the MMT (0%, 1.5 and 3%) were cultured for various duration. ALP MeGC hydrogel matrices in all the nanocomposite groups staining and ALP activity at day 3, 7 and alizarin red S staining without aggregation, which was confirmed by EDS (Supplemen- and its quantification at day 14, 28 were carried out (Fig. 5). The tary Table 1) and TEM observation (Supplementary Fig. 2). 1.5% MeGC-MMT group exhibited the most intensified staining of ALP at both days 3 and 7 compared with the other two groups Cytocompatibility of MeGC-MMT nanocomposite hydrogels. (Fig. 5a), and ALP activity also showed the highest values for Interconnected porous microstructure was created in the MeGC- 1.5% MeGC-MMT group at both time points (Fig. 5b); while MMT nanocomposite hydrogels, the viability of cells inside which comparable mineralization to the 3% MeGC-MMT group was further investigated to optimize the composition of MMT for (Fig. 5c). bone tissue engineering. BMSCs were encapsulated inside of qRT-PCR was employed to evaluate the differentiation of various MeGC-MMT nanocomposite hydrogels and MeGC BMSCs encapsulated in the MeGC-MMT nanocomposite hydro- hydrogels were employed as control. After 24-h culture, the cell gels at a gene level. The gene expression of ALP, an early viability of the 1% and 1.5% MeGC-MMT groups was sig- osteogenic marker, and Runx2, one of the most specific nificantly higher compared with the other groups (Fig. 3; Sup- osteogenic differentiation markers in the earlier stage and OCN, plementary Fig. 3). Considering the initial characterization of the a late osteogenic marker, was examined at days 7 and 14, composite biomaterial, including the morphology, porosity, and respectively, and these results are presented in Fig. 6. Consistent cell viability, we chose the 1.5 and 3% MMT groups for further results confirmed that the 1.5% MeGC-MMT exhibited the most investigation. The hematoxyline and eosin (H&E) staining images powerful ability to promote the differentiation of BMSCs with a NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications 3 Equilibrium water content (%) Gel weight (%) Young’s modulus (kPa) ARTICLE NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 MeGC MeGC + 0.5% MMT MeGC + 1% MMT MeGC + 1.5% MMT MeGC + 2% MMT MeGC + 3% MMT MeGC + 4% MMT 50 0 0.5 1 1.5 2 34 MMT weight (%) Fig. 2 Characterization of interconnected microporous structure. SEM images of photocross-linked MeGC-MMT nanocomposite hydrogels. Various amount of MMT ranged from 0 to 4% (scale bar = 100 μm). Percentage porous area covered was quantified by SEM image analysis. Error bars indicate standard deviation tissue along with remaining hydrogels inside the defects. Sections were stained with H&E in Fig. 7; overt native cell recruitment was N.S. observed. There was highly improved cell infiltration with thick 300 granulation tissue formation composed of inflammatory cells and collagen-producing fibroblasts in the defect treated with MeGC- MMT nanocomposite hydrogel compared with the pure MeGC hydrogel group, while poor cell infiltration was observed in the blank control group. As time progressed to week 6 post surgery, all mice were euthanatized for tissue collection. Ex vivo high resolution μCT was employed to evaluate the status of bone healing. The size of the remaining defect treated with 1.5% MeGC-MMT nanocom- posite hydrogel was remarkably smaller than that of the other two groups, representative images are shown in Fig. 8a. The relative new bone growth surface area, bone volume/tissue volume (BV/ −1 TV %), and the trabecular number (Tb. N., mm ) were extracted 0 0.5 1 1.5 2 34 from the μCT images. The normalization was based on the MMT weight (%) original 3 mm defect area (Fig. 8b). The defects treated with Fig. 3 Assessment of cell cytotoxicity. The viability of encapsulated cells in MeGC and 1.5% MeGC-MMT hydrogels were covered by new bone at 38 and 69%, respectively, whereas the defects left empty photocross-linked MeGC-MMT nanocomposite hydrogels with various amount of MMT ranging from 0 to 4%. The hydrogels were cultured for exhibited a minimal healing 6 weeks post surgery (10%). The −1 BV/TV and Tb. N. rose up to 46% and 6.2 mm for the 1.5% 24 h. Error bars indicate standard deviation, *p < 0.05, **p < 0.01, ***p < 0.001, and NS, no significance (ANOVA followed by Tukey’s post hoc test) MeGC-MMT group, considerably higher compared with that of the MeGC group (16% and 1.5) or the blank group (7% and 0.9). The 1.5% MeGC-MMT group resulted in the most effective 5.5-, 3.4-, and 4.5-fold increase of the gene expression of ALP, bone repair in the absence of any exogeneous growth factor or Runx2, and OCN compared with the pure MeGC group. For the stem cells. 3% MeGC-MMT group, statistically significant elevation of all The quality of new bone formation was further examined by three gene expression was observed as well; however, compared histological evaluation with H&E, Masson trichrome staining with the 1.5% MMT group, the ability of osteogenesis was (Fig. 9). The defect treated with 1.5% MeGC-MMT nanocompo- significantly limited. site hydrogel was occupied with newly formed bone, and thick soft tissue and bone-like tissue connected the edges, 6 weeks post surgery. The blank and MeGC groups presented very little bone In vivo bone regeneration of MeGC-MMT nanocomposites. tissue only on the edges of the defects and very thin soft tissue We took a further step to translate our 1.5% MeGC-MMT connecting the defects. Masson trichrome staining revealed an nanocomposite hydrogel to the in vivo critical-sized calvarial osteoid matrix formed on the edge of defects treated with the defect model of mice to evaluate bone regeneration. First, MeGC 1.5% MeGC-MMT nanocomposite hydrogel, whereas defects and 1.5% MeGC-MMT hydrogels were injected in the defects, without any treatment or treated with the MeGC hydrogel were and mice were euthanized 10 days post surgery to harvest the 4 NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications Cell viability (%) Porous area covered (%) NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 ARTICLE 50 µm MeGC MeGC + 1.5% MMT MeGC + 3% MMT Fig. 4 Hematoxyline and eosin (H&E) staining of encapsulated cells in photocross-linked MeGC-MMT nanocomposite hydrogels with various amount of MMT (0%, 1.5%, 3%). The hydrogels were cultured for 7 days (n = 3), scale bar = 50 μm, arrows indicate the cells a b MeGC + MeGC + ALP activity MeGC MMT 1.5 % MMT 3 % Day 3 ALP staining Day 7 Day 3 Day 7 0 1.5 3 MMT weight (%) MeGC + MeGC + MeGC MMT 1.5 % MMT 3 % Day 14 Day 28 Alizarin red S staining Day 14 Day 28 0 1.5 3 MMT weight (%) Fig. 5 Osteoconductivity in vitro. The bioactivity of encapsulated cells in photocross-linked MeGC-MMT nanocomposite hydrogels with various amount of MMT (0%, 1.5%, 3%). The hydrogels were cultured in the osteogenic medium for various time. a ALP staining and b ALP activity were performed at days 3 and 7. c Alizarin red S staining was carried out at days 14 and 28, and their quantification was also evaluated. Error bars indicate standard deviation (three independent cultures, n = 6). *p< 0.05, **p < 0.01, and ***p < 0.001 compared with the control MeGC group with 0% MMT (ANOVA followed by Tukey’s post hoc test) only filled with fibrous soft tissue with minimal bone healing. We transport in 3D constructs . However, current technology of pre- have also performed immunohistochemical staining for the paring such interconnected microporous hydrogels by using sacri- osteogenic markers Runx2 and OCN (Supplementary Fig. 8). ficing porogen requires severe porogen removal steps employing Strong staining was observed in the 1.5% MeGC-MMT acid or base, high temperature, solvent that may have significant nanocomposite hydrogel group presenting osteoblastic cells, toxicity to encapsulating cells. This study demonstrated a proof-of- while weak immunostaining for Runx2 and OCN was detected concept of creating interconnected microporous structure in in situ- within the fibrous tissue induced with the blank and MeGC forming hydrogels by exploring intercalation chemistry in polymer groups. No signs of inflammatory responses were noted in all chains with nanoclay to facilitate cell filtration, adhesion, pro- treated groups. liferation, and differentiation for tissue engineering . The positively charged MeGC polymers were well-mixed with weakly negatively charged MMT. The delamination of the individual layers of MMT evenly distributed inside the cured nanocomposite hydrogel, Discussion probably led to the interconnected microporous structure. The Injectable formulations of living cells and bioactive molecules using hydrogels would be an ideal route of administration to the target unique structure of smectites contributes to the microstructure change through interactions between nanoclay and polymer chains area without any surgical incisions. In particular, high porosity and creating interlay-pores and inter-particle spaces. This discovery is of pore interconnectivity are critical for cell ingrowth and mass NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications 5 Relative mineralization quantification ALP activity/DNA (mM/ng) ARTICLE NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 ALP Rumx2 OCN 7 4 0 0 0 1.5 30 1.5 3 0 1.5 3 MMT weight (%) MMT weight (%) MMT weight (%) Fig. 6 Gene markers evaluation of osteogenesis. The bioactivity of encapsulated cells was evaluated with qRT-PCR in photocross-linked MeGC-MMT nanocomposite hydrogels with various amount of MMT (0%, 1.5%, 3%). The hydrogels were cultured in the osteogenic medium for various time. ALP and Runx2 were examined at day 7, and OCN was measured at day 14. Error bars indicate standard deviation (three independent cultures, n = 6), *p < 0.05, **p < 0.01, and ***p < 0.001 (ANOVA followed by Tukey’s post hoc test) Blank MeGC MeGC + 1.5% MMT 4× 500 µm 10× 200 µm 20× 100 µm Fig. 7 In vivo cell infiltration assessment. Histological analysis of native cell infiltration in the hydrogels with hematoxyline and eosin (H&E) staining in calvarial defects, 10 days post surgery considerable importance in the development of the next generation promoted bone regeneration without any growth factor, small of in situ-forming hydrogels with high porosity, because the process molecular drug, or gene. of tissue formation is a highly orchestrated set of temporal and Chitosan, a naturally occurring polymer, is appealing for spatial events that involve infiltration and proliferation of the stem/ tissue-engineering applications due to its high biocompatibility progenitor cells, matrix deposition, and vascularization. and hydrophilicity. MeGC hydrogels can be prepared under mild In recent years, our group has developed a series of chitosan- conditions with visible blue light. The MeGC hydrogel system based injectable hydrogels, which were designed for bone tissue supported proliferation and extracellular matrix deposition of 25,26,32,35 engineering. For example, Arg–Gly–Asp (RGD) and phospho- encapsulated mesenchymal stem cells as well . MMT serine (PS)-conjugated MeGC (PS–RGD–MeGC) hydrogel was belongs to the smectites with a unique structure and high aspect synthesized to improve cell adhesion by the RGD motifs and ratio, no harm to the environment or cells (Fig. 3) and low cost. promote osteogenesis by enhancing cell–matrix interactions and Typically, the mechanical properties of hydrogels restrict their hydroxyapatite nucleation through PS . Polysulfonate MeGC application on hard tissue engineering, such as bone. The lit- was rationally designed to sequester and stabilize endogenous erature reported that the inclusion of nanoclay in the polymer bone morphogenic protein-2 (BMP-2), therefore to achieve matrices significantly improve the mechanical and thermal 36–38 osteogenesis without applying its exogenous supraphysiological properties . Young’s modulus and EWC (Fig. 1b, c) were 31 32 dosage . In addition, the bone-forming sterosomes , the small analyzed and, with the increment of MMT, the mechanical molecular drug, i.e. phenamil, and siRNA knocking down noggin properties were remarkably improved. The Young’s modulus 33,34 expression were embedded in MeGC hydrogels to enhance was increased more than six times with the 4% MeGC-MMT bone regeneration. All the efforts were proven to be effective nanocomposite hydrogel compared with that of MeGC hydrogel in vivo in the calvarial defects of mice to a large extent, however, (<10 kPa), and it was around 30 kPa for the 1.5% MeGC-MMT none of the formulation alone has achieved fully satisfied results nanocomposite hydrogel. within the experimental time yet, typically 6 weeks. One of the MMT has been used as a filler in hard scaffolds, such as 24 23 phenomena in all those injectable chitosan-based hydrogels is chitosan–MMT–hydroxyapatite , gelatin–MMT–cellulose scaf- that they all bear a continuous microstructure without any pores folds and silk biomaterials for bone tissue formation. To the best (Figs. 2, 4), which probably limits their osteoconductive cap- of our knowledge, it is the first report to utilize MMT to create ability. The current MeGC-MMT formulation significantly interconnected microporous structure in chitosan-based in situ- 6 NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications Fold change Fold change Fold change NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 ARTICLE Blank MeGC MeGC + 1.5% MMT 80 10 NS 0 0 0 Fig. 8 Evaluation of in vivo bone regeneration. a Microcomputed tomography images of calvarial defects treated with hydrogels or left empty, 6 weeks post surgery, scale bar = 1 mm. b μCT quantification of bone regeneration in calvarial defects. Relative bone growth surface area, bone volume density −1 (BV/TV%), and the trabecular number (Tb.N., mm ). *p < 0.05, ***p < 0.001, and NS, no significance (ANOVA followed by Tukey’s post hoc test). Error bars indicate standard deviation (n = 6) Blank MeGC MeGC + 1.5% MMT 500 µm 100 µm 100 µm Fig. 9 Histological analysis of bone regeneration in calvarial defects, 6 weeks post surgery. Hematoxyline and eosin (H&E) staining (scale bar = 500 μm), and magnified images of H&E and Masson trichrome staining (red boxes represent the magnified areas, scale bar = 100 μm; arrowheads indicate the new bone formation) forming hydrogel just simply by intercalation chemistry (Supple- 1.5% w/v, to exhibit the strongest conductive ability in BMSCs, mentary Fig. 4). We confirmed the existence of pores in the wet confirmed with ALP and mineralization staining and their cor- state of hydrogel with low-vacuum SEM (Fig. 2) and also with responding quantification (Fig. 5), as well as specific gene frozen sections (Fig. 4) to avoid the ambiguity as the evaporation expression (Fig. 6). We tempted to culture BMSCs only in the of water in the hydrogel leading to the formation of pores. basal medium in the MeGC-MMT hydrogels, and osteogenic Further in vitro characterization of the combination of markers (ALP and Runx2) were upregulated on day 7 (Supple- MeGC and MMT showed that the optimal weight of MMT was mentary Fig. 6), indicating potential osteoinductive ability of this NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications 7 Blank MeGC MeGC + 1.5% MMT Blank MeGC MeGC + 1.5% MMT Blank MeGC MeGC + 1.5% MMT MT H&E Relative bone growth surface area % BV/TV % –1 Tb. N. (mm ) ARTICLE NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 composite. Further experiments of ectopic bone formation should fetal bovine serum (FBS, MT35010CV) was supplied by Mediatech Inc. (Manassas, VA, USA). The live/dead staining solution (L3224) and cDNA transcription kit be carried out to confirm this speculation. The most appealing (18080-051) were purchased from Invitrogen (Carlsbad, CA, USA). Trizol reagent character of this nanocomposite hydrogel is their ability in vivo to (15596018) and RNeasy Mini Plant Kit (74106) were supplied by Qiagen (Valencia, recruit native cells for de novo bone formation (Fig. 7). Satis- CA, USA). Pierce BCA Protein Assay Kit was obtained from Thermo Scientific factory results were obtained in the mouse calvarial defect model (23235, Rockford, IL, USA). The nude mice were purchased from Charles River Laboratories (Wilmington, MA, USA). All solvents and products were used as only by applying the material itself in the absence of cells or received. growth factors. This feature is of importance for translation to the clinical application, as cells were not necessary to be harvested Preparation of MeGC and MeGC-MMT nanocomposite hydrogels. In all, 2% and re-implanted back to the patients or expensive growth fac- (w/v) GC and GMA were mixed at 1:1 molar ratio of GMA to the amino groups in tors, such as BMP-2, to be applied. In addition, we have treated GC in Milli-Q water. The solution was adjusted to pH 9.0, and placed on a shaker the mouse calvarial defect with commercially available collagen at room temperature. After 40 h, pH was readjusted to 7.0, and the solution was 39 40 dialyzed with 50 kDa tubes against Milli-Q water for 16 h. After lyophilization and sponge and demineralized bone matrix (DBX) for 6 weeks as rehydration with phosphate-buffered saline (PBS), a 4% (w/v) MeGC solution was positive controls (Supplementary Fig. 7). The defects treated with −1 obtained. MMT was dispersed in Milli-Q water and yield 100 mg mL stock 1.5% MeGC-MMT hydrogels, collagen sponge, and DBX were dispersion. Various amount of MMT (w/v) was mixed with 4% (w/v) MeGC covered by new bone at 69, 23, and 26%, respectively. The BV/TV solution, and the hydrogels were cured under visible blue light with riboflavin as −1 and Tb. N. rose up to 46% and 6.2 mm for the 1.5% MeGC- the photoinitiator (final concentration 6 µM). The final concentration of MeGC is 2% (w/v) with various amount of MMT. MMT group, considerably higher compared with that of the collagen sponge group (6.4% and 2.7) or the DBX group (13.1% Characterization of MeGC-MMT nanocomposite hydrogels. A 400 μLof and 4.3). Compared with the commercially available materials, hydrogel mixture was cured for 80 s in a 48-well plate, and compressive modulus our newly developed MeGC-MMT hydrogel presents great was measured using a flat-ended indenter (1.6 mm in diameter) on an Instron potential for bone healing. Electro-Mechanical Testing Machins (Instron, Model 5564, Norwood, MA, USA). The surface charges and the exchange ability of MMT were The Young’s modulus was determined from the slope of linear portion of the 25,41,42 11 obtained stress-strain curve using a Poissons’ ratio of 0.25 . well studied for delivery purpose . Taking advantage of this, Hydrogels were equilibrated in PBS for 24 h and lyophilized for 16 h to obtain further development of this chitosan-based hydrogel system dry gels. EWC was calculated using the following equation specifically for bone tissue engineering is indispensable. The M  M combination and optimization of PS–RGD–MeGC or poly- w d EWC ¼ ; ð1Þ sulfonate MeGC and MMT hold great promise for bone regen- w eration. The critical defect fully filled with regenerated bone where M and M refer to the weight of wet and dry hydrogels, respectively. w d within experimental time will soon be achieved. The underlying The degradation of hydrogels was carried out for 42 days. Hydrogels were incubated at 37 °C in cell growth medium (DMEM, 10% FBS, and 1% P/S), which mechanism should be further studied to elucidate the strong was replaced every 7 days. At a pre-set time, hydrogels were lyophilized for weight osteoconductivity of this biomaterial and the type of tissue that measurement. The present residual weight of hydrogels was calculated using the replaces the hydrogel site with time progression needs to be following equation closely monitored as well in our future work. Our novel nanocomposite hydrogel system can be utilized not Residual dry gel weight% ¼ ´ 100%; ð2Þ only as 3D scaffolds for non-load-bearing injury site such as 0 cranial defects but also synthetic biological carriers for stem cells, where M and M refer to the weight of hydrogels at time 0 (hydrogels did not 0 t bioactive agents, or currently available tissue grafts for regen- undergo degradation) and t, respectively. Hydrogels were imaged in low vacuum using scanning electron microscopy erative medicine. We did not evaluate mechanical properties of with X-ray microanalysis (SEM/EDS, FEI Nova NanoSEM 230, Hillsboro, OR, the regenerated bone in this study, because the calvarial defect is USA) to observe the microstructure and chemical compositions. For Transmission small and the healing site is not load bearing. However, the Electron Microscope (TEM) analysis, hydrogels were embedded in sucrose-PVP further study of biomechanical properties will be needed in other and mounted on Cu EM grid after sectioning with glass 45 knife with 100 -nm fracture models such as large segmental defects. It is possible that thickness (Leica UC6/F6 −100 C), then imaged with JEM1200EX. The thermogravimetric analysis (TGA) was performed with the lyophilized hydrogels the hydrogel may not possess sufficient mechanical properties in o o o −1 from 50 C to 600 Cat 30 C min speed. The Powder X-ray Diffraction (XRD, the load-bearing bone defect sites. A promising alternative is to Bruker Corporation, Germany) pattern was collected from 2 to 10 (2θ range) with combine the hydrogel with solid particles or bone graft materials. ̊ a diffractometer using Ni-filtered CuKα X-ray radiation (λ = 1.5418 A). In conclusion, we have demonstrated that with the help of intercalation chemistry, proportional MMT mixing with MeGC 3D cell culture in MeGC-MMT nanocomposite hydrogels. BMSCs at a density 6 −1 leads to the formation of interconnected microporous structure, of 2 × 10 cells mL were mixed in MeGC-MMT dispersion. The hydrogel was cured by exposing 40 μL of the suspension to visible blue light (400–500 nm, which can promote native cell infiltration, proliferation, and −2 500–600 mW cm , Bisco Inc., Schaumburg, IL) in the presence of riboflavin as a in situ differentiation in the absence of any growth factors, small photoinitiator, (final concentration 6 µM). The resulting hydrogels were incubated molecular drugs, or genes. This biocompatible, bioactive, and in 1 mL of media accordingly. injectable nanocomposite material shows great promise being applied in a wide range of tissue regeneration. Cytotoxicity. Cytotoxicity of MMT in nanocomposite hydrogels was evaluated using alamar blue assay and live/dead staining (Invitrogen, Carlsbad, CA). Hydrogels with encapsulated BMSCs and MMT in various concentrations were Methods incubated at 37 C and 5% CO in fresh growth medium. Then the medium was Materials. Glycol chitosan (GC, ~100 kDa, 072-1581) was purchased from Wako replaced with 10% (v/v) alamar blue solution in growth medium at a pre- Chemicals USA, Inc. (Richmond, VA, USA). MMT (682659-500G), glycidyl determined time. After a 3-h incubation, the fluorescence intensity (F) of alamar methacrylate (GMA, 151238-100G), 1-ethyl-3-(3-dimethylaminopropyl)-carbodii- blue was measured at 585 nm with an excitation wavelength of 570 nm. For the mide (EDC, E1769-10G), tween-20 (p1379-500ML), p-nitrophenyl phosphate blank group, the 10% (v/v) alamar blue solution was added in an empty well and (N7653-100ML), β-glycerophosphate (G5422-100G), -ascorbic acid (A5960-25G), L incubated together with other samples. The relative cell viability (%) was calculated dexamethasone (D4902-25MG), Nitro Blue Tetrazolium (NBT, N5514-25TAB), 5- using the following equation bromo-4-chloro-3-indoxylphosphate (BCIP, B6149-50MG), alizarin red S (A5533), and ethylenediaminetetraacetic acid (EDTA, EDS-1KG) were supplied by Sigma- F  F s b Relative cell viability ¼ ´ 100%; ð3Þ Aldrich (St. Louis, MO, USA). The mouse bone marrow stromal cell line (BMSCs, F  F c b D1 ORL UVA [D1], D1 cell, CRL-12424) was obtained from American Type Culture Collection (ATCC, Manassas, VA, USA). High glucose Dulbecco’s Mod- where F , F , and F refer to the fluorescence intensity of the sample after incu- s c b ified Eagle’s Medium (DMEM, 11995-065), penicillin/streptomycin (100 U/mL, bation for 24 h, intensity of the corresponding sample without treatment, and the 15140122) were purchased from Life Technologies (Grand Island, NY, USA), and blank, respectively. 8 NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 ARTICLE ALP and alizarin red S staining and quantification. Gels were incubated in Data availability osteogenic media (growth medium was supplemented with 10 mM β-glyceropho- All relevant data are available within the article and Supplementary Information. The −1 sphate, 50 μgmL L-ascorbic acid, and 100 nM dexamethasone). At a pre- source data underlying Figs. 1B–D, 2, 3, 5B–C, and 66 and Supplementary Figs. 1, 4, 5, 6, determined time, gels were fixed in 10% formalin for 20 min, rinsed with PBS, and and 7B are provided as a Source Data file. Requests for other materials should be incubated in a solution consisting of NBT and BCIP stock solutions in an ALP addressed to the corresponding author. buffer (100 mM Tris, 50 mM MgCl , 100 mM NaCl, pH 8.5) for 2 h, at room temperature. The stained samples were observed with an Olympus SZX16 Ste- Received: 20 October 2018 Accepted: 1 July 2019 reomicroscope (Olympus, Tokyo, Japan). ALP expression appeared in blue. For the ALP activity assay, gels were rinsed with PBS, incubated in a lysis buffer (0.1% Tween-20 in PBS) at 4 °C for 5 min. ALP activity was determined colorimetrically using p-nitrophenyl phosphate as a substrate and measured at 405 nm. The ALP activity was normalized to total DNA contents measured with the picogreen assay (Thermo Scientific, Rockford, IL). Gels were fixed in 10% formalin for 20 min, rinsed with PBS, and incubated in References 2% alizarin red S solution for 5 min at room temperature. Then, the gels were 1. Langer, R. & Vacanti, J. P. Tissue engineering. Science 260, 920–926 (1993). washed with PBS under gentle shaking for 16 h, and the PBS was changed at least 2. Dvir, T., Timko, B. P., Kohane, D. S. & Langer, R. Nanotechnological strategies three times. The stained samples were observed with the Olympus SZX16 for engineering complex tissues. Nat. Nanotechnol. 6,13–22 (2011). Stereomicroscope. Calcium deposition appeared in red. The semi quantification of 3. Lutolf, M. P. & Hubbell, J. A. Synthetic biomaterials as instructive extracellular alizarin red S staining was carried out with acetic acid extraction and neutralization microenvironments for morphogenesis in tissue engineering. Nat. 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Mayer’s hematoxylin. 22. Mieszawska, A. J. et al. Clay enriched silk biomaterials for bone formation. Acta Biomater. 7, 3036–3041 (2011). Statistical analysis. Three independent experiments at least, unless otherwise 23. Hsu, S. H., Wang, M. C. & Lin, J. J. Biocompatibility and antimicrobial stated, were performed and data were presented as mean ± standard deviation. evaluation of montmorillonite/chitosan nanocomposites. Appl. Clay Sci. 56, Multiple comparisons were assessed using one-way or two-way analysis of variance 53–62 (2012). (ANOVA). The analysis of variances followed by Tukey’s post hoc test was 24. Katti, K. S., Katti, D. R. & Dash, R. Synthesis and characterization of a novel employed in this work and p < 0.05 was considered statistically significant. chitosan/montmorillonite/hydroxyapatite nanocomposite for bone tissue engineering. Biomed. Mater. 3, 12 (2008). 25. Hu, J. L. et al. Visible light crosslinkable chitosan hydrogels for tissue Reporting summary. Further information on research design is available in the Nature Research Reporting Summary linked to this article. engineering. Acta Biomater. 8, 1730–1738 (2012). NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications 9 ARTICLE NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 26. Choi, B. et al. Covalently conjugated transforming growth factor-beta 1 in 42. Choi, B. et al. Visible-light-initiated hydrogels preserving cartilage modular chitosan hydrogels for the effective treatment of articular cartilage extracellular signaling for inducing chondrogenesis of mesenchymal stem defects. Biomater. Sci. 3, 742–752 (2015). cells. Acta Biomater. 12,30–41 (2015). 27. Salerno, A., Di Maio, E., Iannace, S. & Netti, P. A. Tailoring the pore structure of PCL scaffolds for tissue engineering prepared via gas foaming of multi- Acknowledgements phase blends. J. Porous Mater. 19, 181–188 (2012). This work was supported by the grants from the National Natural Science Foundation of 28. Anseth, K. S., Bowman, C. N. & Brannon-Peppas, L. Mechanical properties of China 31800840, the National Institutes of Health R01 AR060213, R01 DE027332, the hydrogels and their experimental determination. Biomaterials 17, 1647–1657 Department of Defense W81XWH-18-1-0337, and the MTF Biologics. (1996). 29. Cui, Z.-K. et al. Design and characterization of a therapeutic non- phospholipid liposomal nanocarrier with osteoinductive characteristics to Author contributions promote bone formation. ACS Nano 11, 8055–8063 (2017). Z.-K.C. and M.L. conceived the ideas for experimental designs, analyzed the data, and 30. Kim, S. et al. Photocrosslinkable chitosan hydrogels functionalized with the wrote the paper. Z.-K.C., S.K., and J.J.B. conducted all the experiments and helped with RGD peptide and phosphoserine to enhance osteogenesis. J. Mater. Chem. B 4, paper preparation. T.A. performed the animal surgery. B.M.W. provided suggestions for 5289–5298 (2016). the project. Z.-K.C. and S.K. contributed equally to this work. 31. Kim, S., Cui, Z.-K., Kim, P. J., Jung, L. Y. & Lee, M. Design of hydrogels to stabilize and enhance bone morphogenetic protein activity by heparin mimetics. Acta Biomater. 72,45–54 (2018). Additional information 32. Cui, Z. K. et al. Design and characterization of a therapeutic non-phospholipid Supplementary Information accompanies this paper at https://doi.org/10.1038/s41467- liposomal nanocarrier with osteoinductive characteristics to promote bone 019-11511-3. formation. ACS Nano 11, 8055–8063 (2017). Competing interests: The authors declare no competing interests. 33. Cui, Z. K. et al. Simultaneous delivery of hydrophobic small molecules and siRNA using Sterosomes to direct mesenchymal stem cell differentiation for bone repair. Acta Biomater. 58, 214–224 (2017). Reprints and permission information is available online at http://npg.nature.com/ 34. Cui, Z.-K. et al. Delivery of siRNA via cationic Sterosomes to enhance reprintsandpermissions/ osteogenic differentiation of mesenchymal stem cells. J. Control Release 217, 42–52 (2015). Peer review information: Nature Communications thanks Oliver Betz and other 35. Park, H., Choi, B., Hu, J. & Lee, M. Injectable chitosan hyaluronic acid anonymous reviewer(s) for their contribution to the peer review of this work. hydrogels for cartilage tissue engineering. Acta Biomater. 9, 4779–4786 (2013). 36. Kojima, Y. et al. Synthesis of nylon 6–clay hybrid by montmorillonite Publisher’s note: Springer Nature remains neutral with regard to jurisdictional claims in intercalated with ϵ-caprolactam. J. Polym. Sci. Pol. Chem. 31, 983–986 (1993). published maps and institutional affiliations. 37. Sikdar, D., Katti, D., Katti, K. & Mohanty, B. Effect of organic modifiers on dynamic and static nanomechanical properties and crystallinity of intercalated clay-polycaprolactam nanocomposites. J. Appl. Polym. Sci. 105, 790–802 Open Access This article is licensed under a Creative Commons (2007). Attribution 4.0 International License, which permits use, sharing, 38. Yu, S. Z., Zhao, J. H., Chen, G., Juay, Y. K. & Yong, M. S. The characteristics of adaptation, distribution and reproduction in any medium or format, as long as you give polyamide layered-silicate nanocomposites. J. Mater. Process Technol. 192, appropriate credit to the original author(s) and the source, provide a link to the Creative 410–414 (2007). Commons license, and indicate if changes were made. The images or other third party 39. Yang, H. S. et al. Comparison between heparin-conjugated fibrin and collagen material in this article are included in the article’s Creative Commons license, unless sponge as bone morphogenetic protein-2 carriers for bone regeneration. Exp. indicated otherwise in a credit line to the material. If material is not included in the Mol. Med. 44, 350–355 (2012). article’s Creative Commons license and your intended use is not permitted by statutory 40. Townsend, J. M. et al. Effects of tissue processing on bioactivity of cartilage regulation or exceeds the permitted use, you will need to obtain permission directly from matrix-based hydrogels encapsulating osteoconductive particles. Biomed. the copyright holder. To view a copy of this license, visit http://creativecommons.org/ Mater. 13, 034108 (2018). licenses/by/4.0/. 41. Amsden, B. G., Sukarto, A., Knight, D. K. & Shapka, S. N. Methacrylated glycol chitosan as a photopolymerizable biomaterial. 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Microporous methacrylated glycol chitosan-montmorillonite nanocomposite hydrogel for bone tissue engineering

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Science, Humanities and Social Sciences, multidisciplinary; Science, Humanities and Social Sciences, multidisciplinary; Science, multidisciplinary
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Abstract

ARTICLE https://doi.org/10.1038/s41467-019-11511-3 OPEN Microporous methacrylated glycol chitosan- montmorillonite nanocomposite hydrogel for bone tissue engineering 1,2,5 3,5 3 2,3 4 2,3 Zhong-Kai Cui , Soyon Kim , Jessalyn J. Baljon , Benjamin M. Wu , Tara Aghaloo & Min Lee Injectable hydrogels can fill irregular defects and promote in situ tissue regrowth and regeneration. The ability of directing stem cell differentiation in a three-dimensional micro- environment for bone regeneration remains a challenge. In this study, we successfully nanoengineer an interconnected microporous networked photocrosslinkable chitosan in situ- forming hydrogel by introducing two-dimensional nanoclay particles with intercalation chemistry. The presence of the nanosilicates increases the Young’s modulus and stalls the degradation rate of the resulting hydrogels. We demonstrate that the reinforced hydrogels promote the proliferation as well as the attachment and induced the differentiation of encapsulated mesenchymal stem cells in vitro. Furthermore, we explore the effects of nanoengineered hydrogels in vivo with the critical-sized mouse calvarial defect model. Our results confirm that chitosan-montmorillonite hydrogels are able to recruit native cells and promote calvarial healing without delivery of additional therapeutic agents or stem cells, indicating their tissue engineering potential. 1 2 Department of Cell Biology, School of Basic Medical Sciences, Southern Medical University, Guangzhou, Guangdong 510515, China. Division of Advanced Prosthodontics, University of California Los Angeles, 10833 Le Conte Avenue, Los Angeles, CA 90095, USA. Department of Bioengineering, University of California Los Angeles, 420 Westwood Plaza, Los Angeles, CA 90095, USA. Division of Diagnostic and Surgical Sciences, University of California Los Angeles, 10833 Le Conte Avenue, Los Angeles, CA 90095, USA. These authors contributed equally: Zhong-Kai Cui, Soyon Kim. Correspondence and requests for materials should be addressed to Z.-K.C. (email: zhongkaicui@smu.edu.cn) or to M.L. (email: leemin@ucla.edu) NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications 1 1234567890():,; ARTICLE NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 issue engineering exploits a combination of cells, engi- include any live cells during fabrication of those scaffolds with neering and materials methods, along with proper bio- the conventional methods to create porous structure, such as Tchemical and physiochemical factors to improve or replace freeze-drying, porogen leaching. After the establishment of the biological tissues, including the skin, cartilage, bladder, blood scaffolds, cells were loaded leading to nonuniform distribution. vessels, and bone . The decisive contribution of the tissue engi- Considering this, MMT is of interest for the formation of porous neering matrix in ensuring the regeneration potential of stem and structure, and to be investigated further in vivo for TERM as an 2–4 progenitor cells has been emphasized increasingly . Progress in application. cell selection, cell culture, and new material formulations has led Hydrogels derived from natural products are an appealing to the development of more effective therapies for tissue engi- three-dimensional biomaterial for tissue engineering. Compared neering and regenerative medicine (TERM). Various materials with hard scaffolds that require pre-shaping, soft hydrogels are have been explored for regenerative applications, including injectable and can fill any irregular shaped defects in a minimally 5,6 naturally occurring products and synthetic materials . Generally invasive manner. Recently, we have developed a photoinducible speaking, the intrinsic properties of naturally derived materials, hydrogel system of methacrylated glycol chitosan (MeGC) using such as collagens, are definitely attractive; however, the com- riboflavin as a photoinitiator . MeGC hydrogels supported plexities of purification, immunogenicity, mechanical properties, proliferation and extracellular matrix deposition of encapsulated 25,26 and pathogen transmission significantly limit their application. In mesenchymal stem cells ; however, the hydrogel itself showed light of this, greater control over material properties and tissue minimal bone-forming ability without osteogenic factors, bioac- responses has garnered more attention based on designable tive molecules, or encapsulated cells. One of the possible reasons modern material science. is lacking of porous structure in the MeGC hydrogel. Porosity is Biomaterials are having great impact on the medical treatments reported to be necessary for new tissue formation as it allows the in helping to solve clinical problems. In the field of medicine, cells to migrate, infiltrate, and proliferate in a 3D environment, as drug-eluting stent coated with polymers and controlled drug well as for the vascularization, differentiation, and mass trans- release systems constituted with biomaterials have been saving port . Therefore, it is of interest to develop a hydrogel system 7,8 hundreds of thousands of lives each year . In the field of TERM, that could recruit native cells and facilitate bone formation with a the combination of biomaterial scaffold and certain cells is pos- microporous interconnected structure. sible to replace biological tissues . In the field of medical devices, In this study, we introduce MMT to the photopolymerizable biomaterials are also playing a significantly important role as the MeGC hydrogel system to fabricate an injectable highly osteo- core components in surgical sutures, bioadhesives, and dental conductive in situ-forming biomaterial for bone regeneration. We implants . Clays and clay minerals are emerging materials for hypothesize that MMT can not only improve the microstructure biomaterial design to provide new strategies for TERM. The usage but also enhance the mechanical properties of cured MeGC of inorganic layered nanomaterials for medical purpose dates hydrogel. Therefore, we started with optimizing the proportion of back to ancient time, such as for wound healing and hemorrhage MMT in hydrogels and investigated the osteogenesis of encap- inhibition . Nowadays, clays and clay minerals are being applied sulated mesenchymal stem cells in vitro. We further evaluated the in pharmaceuticals as active ingredients or excipients, and in ability of the nanocomposite hydrogels to recruit native cells and cosmetics as creams, powders, and emulsions . The interactions improve bone formation in a critical-sized mouse calvarial defect between clay nanoparticles and drugs as well as other biological model. The treatment of bone defects remains one of the largest molecules have been well investigated and therefore exploited for challenges in musculoskeletal TERM, thereby our developed 11,12 controlled delivery , and moreover, the addition of clay into nanosilicate-loaded MeGC hydrogel may represent a new mate- polymers enhances the mechanical properties owing to the for- rial design in the broadly interesting area of growth factor free 13,14 mation of nanocomposites . and cell-free strategies. This bioactive nanocomposite hydrogel Montmorillonite (MMT) is a major component of Bentonite, can provide an effective treatment for bone defects. which is already approved by the FDA as an additive in various medicinal products . Studies on the acute and chronic toxicities of MMT have confirmed the absence of any negative effects, Results 15,16 even on embryos of pregnant Sprague–Dawley rats . MMT Characterization of MeGC-MMT nanocomposite hydrogels. is a layered silicate [(Na,Ca) (Al,Mg) Si O (OH) � nH O], Nanocomposite hydrogels including various amount of MMT 0.33 2 4 10 2 2 belonging to the smectite group of minerals, with high (0.5-4% w/v) were prepared as illustrated in Fig. 1a. We already 2 −1 specific surface area (up to 600 m g ) and aspect ratio. The tested higher amount of MMT up to 10% in MeGC hydrogels. It repeating structural unit of MMT consists of one alumina octa- became too viscous to mix with poor handling, as the MMT ratio hedral sheet sandwiched in between two silicon tetrahedral layers. increased over 4%. Therefore, we have selected the MMT ratio The overall surface charge is light negative because of the dom- between 0.5 and 4% for the investigation. The electrostatic ination of the oxide anions, which facilitates mixing with cationic interactions between the heterogeneously distributed charges of agents. The MMT particles are typically in a plate-shape, ren- discotic MMT (overall negative charge) and MeGC hydrogel dering ~1 nm in thickness and 0.2–2 μm in diameter. The bio- matrices (positive charge) enhanced the Young’s modulus with an compatibility, availability, and feasibility of this particular mineral increasing amount of MMT in the nanocomposite hydrogels from has garnered significant attention in recent years. Extensive 10 kPa to over 60 kPa (Fig. 1b). Equilibrium water content (EWC) research has been carried out to investigate for the purpose of inversely represents the cross-linking density and the mechanical 11,17,18 28 drug and gene delivery with natural or modified MMT .In properties of a hydrogel . Introduction of more than 0.5% MMT addition, several reports have demonstrated that fabrication of in MeGC hydrogels significantly decreased the EWC of corre- scaffolds by introduction of MMT in natural biomaterials, sponding nanocomposite hydrogels (Fig. 1c), indicating enhanced 19,20 21 22 23,24 including gelatin , collagen , silk , and chitosan , mechanical properties. The dry weights of the MeGC-MMT improved cell-scaffold interactions, cell proliferation, and nanocomposite hydrogels were recorded for 42 days (Fig. 1d). As enhanced cell differentiation. Although all those hard scaffolds observed in the degradation profiles, the degradation rate of the were reported with porous structure under high vacuum and dry 1.5% and 3.0% MMT incorporated MeGC hydrogels was sig- conditions, besides those experiments were limited only in vitro, nificantly decreased compared with the control MeGC hydrogels which are not ideal to be applied in TERM, as it is not possible to by ~40%. TGA degradation profiles confirmed the presence of 2 NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 ARTICLE ab Visible blue light MMT Riboflavin MeGC MeGC-MMT 0 0.5 1 1.5 2 3 4 MMT weight (%) cd 90 0 0 7 14 21 28 35 42 0 0.5 1 1.5 2 34 MMT weight (%) Time (Day) Fig. 1 Characterization of photocross-linked MeGC-MMT nanocomposite hydrogels. a Schematic illustration of MeGC-MMT hydrogel cured by visible blue light cross-linking with riboflavin as the photoinitiator in the presence of MMT. b Young’s modulus of MeGC-MMT nanocomposite hydrogels with various amount of MMT. c Equilibrium water content of MeGC-MMT nanocomposite hydrogels. d Degradation profiles of MeGC-MMT nanocomposite hydrogels in the presence of 0% (square), 1.5% (circle), and 3% (triangle) MMT in PBS at 37 °C for 42 days. Error bars indicate standard deviation (n = 5). *p < 0.05, **p < 0.01, and ***p < 0.001 compared with the control MeGC group with 0% MMT (ANOVA followed by Tukey’s post hoc test) MMT enhanced the thermal stability of MeGC hydrogels (Sup- confirm the interconnected microporous structure in the 1.5 and plementary Fig. 1). 3% MMT groups, and the highest number of cells was observed in The microstructure of MeGC and MeGC-MMT hydrogels the 1.5% MMT group, while MeGC hydrogel exhibited a smooth observed by SEM (Fig. 2) confirms the formation of a surface (Fig. 4). In addition, round-shaped cells were observed in microporous and interconnected network when MMT exceeds the MeGC image, while spread cell morphology was seen in the 1.5%, while a smooth continuous surface was acquired for the MMT containing nanocomposite hydrogels. Taken together, the control MeGC group, and a coarse continuous surface appeared 1.5% MeGC-MMT nanocomposite hydrogel group presented for 0.5 and 1% MMT groups. Porosity was quantified as an index the best support for cell proliferation. of surface area occupied by the pores in the SEM images. The pore sizes were around 115 ± 40 μm for the 1.5% MMT group. Bioactivity of MeGC-MMT nanocomposite hydrogels. Differ- With the increasing MMT content, the pore sizes reached the entiation of BMSCs to osteoblasts typically can be evaluated by maximum of 150 ± 50 μm for the 3.0% MMT group. The the expression of early markers, such as ALP and the ultimate involvement of MMT thoroughly changed the microstructure of calcium deposition . The BMSCs encapsulated in the MeGC- MeGC hydrogels. In addition, MMT was well distributed in the MMT (0%, 1.5 and 3%) were cultured for various duration. ALP MeGC hydrogel matrices in all the nanocomposite groups staining and ALP activity at day 3, 7 and alizarin red S staining without aggregation, which was confirmed by EDS (Supplemen- and its quantification at day 14, 28 were carried out (Fig. 5). The tary Table 1) and TEM observation (Supplementary Fig. 2). 1.5% MeGC-MMT group exhibited the most intensified staining of ALP at both days 3 and 7 compared with the other two groups Cytocompatibility of MeGC-MMT nanocomposite hydrogels. (Fig. 5a), and ALP activity also showed the highest values for Interconnected porous microstructure was created in the MeGC- 1.5% MeGC-MMT group at both time points (Fig. 5b); while MMT nanocomposite hydrogels, the viability of cells inside which comparable mineralization to the 3% MeGC-MMT group was further investigated to optimize the composition of MMT for (Fig. 5c). bone tissue engineering. BMSCs were encapsulated inside of qRT-PCR was employed to evaluate the differentiation of various MeGC-MMT nanocomposite hydrogels and MeGC BMSCs encapsulated in the MeGC-MMT nanocomposite hydro- hydrogels were employed as control. After 24-h culture, the cell gels at a gene level. The gene expression of ALP, an early viability of the 1% and 1.5% MeGC-MMT groups was sig- osteogenic marker, and Runx2, one of the most specific nificantly higher compared with the other groups (Fig. 3; Sup- osteogenic differentiation markers in the earlier stage and OCN, plementary Fig. 3). Considering the initial characterization of the a late osteogenic marker, was examined at days 7 and 14, composite biomaterial, including the morphology, porosity, and respectively, and these results are presented in Fig. 6. Consistent cell viability, we chose the 1.5 and 3% MMT groups for further results confirmed that the 1.5% MeGC-MMT exhibited the most investigation. The hematoxyline and eosin (H&E) staining images powerful ability to promote the differentiation of BMSCs with a NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications 3 Equilibrium water content (%) Gel weight (%) Young’s modulus (kPa) ARTICLE NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 MeGC MeGC + 0.5% MMT MeGC + 1% MMT MeGC + 1.5% MMT MeGC + 2% MMT MeGC + 3% MMT MeGC + 4% MMT 50 0 0.5 1 1.5 2 34 MMT weight (%) Fig. 2 Characterization of interconnected microporous structure. SEM images of photocross-linked MeGC-MMT nanocomposite hydrogels. Various amount of MMT ranged from 0 to 4% (scale bar = 100 μm). Percentage porous area covered was quantified by SEM image analysis. Error bars indicate standard deviation tissue along with remaining hydrogels inside the defects. Sections were stained with H&E in Fig. 7; overt native cell recruitment was N.S. observed. There was highly improved cell infiltration with thick 300 granulation tissue formation composed of inflammatory cells and collagen-producing fibroblasts in the defect treated with MeGC- MMT nanocomposite hydrogel compared with the pure MeGC hydrogel group, while poor cell infiltration was observed in the blank control group. As time progressed to week 6 post surgery, all mice were euthanatized for tissue collection. Ex vivo high resolution μCT was employed to evaluate the status of bone healing. The size of the remaining defect treated with 1.5% MeGC-MMT nanocom- posite hydrogel was remarkably smaller than that of the other two groups, representative images are shown in Fig. 8a. The relative new bone growth surface area, bone volume/tissue volume (BV/ −1 TV %), and the trabecular number (Tb. N., mm ) were extracted 0 0.5 1 1.5 2 34 from the μCT images. The normalization was based on the MMT weight (%) original 3 mm defect area (Fig. 8b). The defects treated with Fig. 3 Assessment of cell cytotoxicity. The viability of encapsulated cells in MeGC and 1.5% MeGC-MMT hydrogels were covered by new bone at 38 and 69%, respectively, whereas the defects left empty photocross-linked MeGC-MMT nanocomposite hydrogels with various amount of MMT ranging from 0 to 4%. The hydrogels were cultured for exhibited a minimal healing 6 weeks post surgery (10%). The −1 BV/TV and Tb. N. rose up to 46% and 6.2 mm for the 1.5% 24 h. Error bars indicate standard deviation, *p < 0.05, **p < 0.01, ***p < 0.001, and NS, no significance (ANOVA followed by Tukey’s post hoc test) MeGC-MMT group, considerably higher compared with that of the MeGC group (16% and 1.5) or the blank group (7% and 0.9). The 1.5% MeGC-MMT group resulted in the most effective 5.5-, 3.4-, and 4.5-fold increase of the gene expression of ALP, bone repair in the absence of any exogeneous growth factor or Runx2, and OCN compared with the pure MeGC group. For the stem cells. 3% MeGC-MMT group, statistically significant elevation of all The quality of new bone formation was further examined by three gene expression was observed as well; however, compared histological evaluation with H&E, Masson trichrome staining with the 1.5% MMT group, the ability of osteogenesis was (Fig. 9). The defect treated with 1.5% MeGC-MMT nanocompo- significantly limited. site hydrogel was occupied with newly formed bone, and thick soft tissue and bone-like tissue connected the edges, 6 weeks post surgery. The blank and MeGC groups presented very little bone In vivo bone regeneration of MeGC-MMT nanocomposites. tissue only on the edges of the defects and very thin soft tissue We took a further step to translate our 1.5% MeGC-MMT connecting the defects. Masson trichrome staining revealed an nanocomposite hydrogel to the in vivo critical-sized calvarial osteoid matrix formed on the edge of defects treated with the defect model of mice to evaluate bone regeneration. First, MeGC 1.5% MeGC-MMT nanocomposite hydrogel, whereas defects and 1.5% MeGC-MMT hydrogels were injected in the defects, without any treatment or treated with the MeGC hydrogel were and mice were euthanized 10 days post surgery to harvest the 4 NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications Cell viability (%) Porous area covered (%) NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 ARTICLE 50 µm MeGC MeGC + 1.5% MMT MeGC + 3% MMT Fig. 4 Hematoxyline and eosin (H&E) staining of encapsulated cells in photocross-linked MeGC-MMT nanocomposite hydrogels with various amount of MMT (0%, 1.5%, 3%). The hydrogels were cultured for 7 days (n = 3), scale bar = 50 μm, arrows indicate the cells a b MeGC + MeGC + ALP activity MeGC MMT 1.5 % MMT 3 % Day 3 ALP staining Day 7 Day 3 Day 7 0 1.5 3 MMT weight (%) MeGC + MeGC + MeGC MMT 1.5 % MMT 3 % Day 14 Day 28 Alizarin red S staining Day 14 Day 28 0 1.5 3 MMT weight (%) Fig. 5 Osteoconductivity in vitro. The bioactivity of encapsulated cells in photocross-linked MeGC-MMT nanocomposite hydrogels with various amount of MMT (0%, 1.5%, 3%). The hydrogels were cultured in the osteogenic medium for various time. a ALP staining and b ALP activity were performed at days 3 and 7. c Alizarin red S staining was carried out at days 14 and 28, and their quantification was also evaluated. Error bars indicate standard deviation (three independent cultures, n = 6). *p< 0.05, **p < 0.01, and ***p < 0.001 compared with the control MeGC group with 0% MMT (ANOVA followed by Tukey’s post hoc test) only filled with fibrous soft tissue with minimal bone healing. We transport in 3D constructs . However, current technology of pre- have also performed immunohistochemical staining for the paring such interconnected microporous hydrogels by using sacri- osteogenic markers Runx2 and OCN (Supplementary Fig. 8). ficing porogen requires severe porogen removal steps employing Strong staining was observed in the 1.5% MeGC-MMT acid or base, high temperature, solvent that may have significant nanocomposite hydrogel group presenting osteoblastic cells, toxicity to encapsulating cells. This study demonstrated a proof-of- while weak immunostaining for Runx2 and OCN was detected concept of creating interconnected microporous structure in in situ- within the fibrous tissue induced with the blank and MeGC forming hydrogels by exploring intercalation chemistry in polymer groups. No signs of inflammatory responses were noted in all chains with nanoclay to facilitate cell filtration, adhesion, pro- treated groups. liferation, and differentiation for tissue engineering . The positively charged MeGC polymers were well-mixed with weakly negatively charged MMT. The delamination of the individual layers of MMT evenly distributed inside the cured nanocomposite hydrogel, Discussion probably led to the interconnected microporous structure. The Injectable formulations of living cells and bioactive molecules using hydrogels would be an ideal route of administration to the target unique structure of smectites contributes to the microstructure change through interactions between nanoclay and polymer chains area without any surgical incisions. In particular, high porosity and creating interlay-pores and inter-particle spaces. This discovery is of pore interconnectivity are critical for cell ingrowth and mass NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications 5 Relative mineralization quantification ALP activity/DNA (mM/ng) ARTICLE NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 ALP Rumx2 OCN 7 4 0 0 0 1.5 30 1.5 3 0 1.5 3 MMT weight (%) MMT weight (%) MMT weight (%) Fig. 6 Gene markers evaluation of osteogenesis. The bioactivity of encapsulated cells was evaluated with qRT-PCR in photocross-linked MeGC-MMT nanocomposite hydrogels with various amount of MMT (0%, 1.5%, 3%). The hydrogels were cultured in the osteogenic medium for various time. ALP and Runx2 were examined at day 7, and OCN was measured at day 14. Error bars indicate standard deviation (three independent cultures, n = 6), *p < 0.05, **p < 0.01, and ***p < 0.001 (ANOVA followed by Tukey’s post hoc test) Blank MeGC MeGC + 1.5% MMT 4× 500 µm 10× 200 µm 20× 100 µm Fig. 7 In vivo cell infiltration assessment. Histological analysis of native cell infiltration in the hydrogels with hematoxyline and eosin (H&E) staining in calvarial defects, 10 days post surgery considerable importance in the development of the next generation promoted bone regeneration without any growth factor, small of in situ-forming hydrogels with high porosity, because the process molecular drug, or gene. of tissue formation is a highly orchestrated set of temporal and Chitosan, a naturally occurring polymer, is appealing for spatial events that involve infiltration and proliferation of the stem/ tissue-engineering applications due to its high biocompatibility progenitor cells, matrix deposition, and vascularization. and hydrophilicity. MeGC hydrogels can be prepared under mild In recent years, our group has developed a series of chitosan- conditions with visible blue light. The MeGC hydrogel system based injectable hydrogels, which were designed for bone tissue supported proliferation and extracellular matrix deposition of 25,26,32,35 engineering. For example, Arg–Gly–Asp (RGD) and phospho- encapsulated mesenchymal stem cells as well . MMT serine (PS)-conjugated MeGC (PS–RGD–MeGC) hydrogel was belongs to the smectites with a unique structure and high aspect synthesized to improve cell adhesion by the RGD motifs and ratio, no harm to the environment or cells (Fig. 3) and low cost. promote osteogenesis by enhancing cell–matrix interactions and Typically, the mechanical properties of hydrogels restrict their hydroxyapatite nucleation through PS . Polysulfonate MeGC application on hard tissue engineering, such as bone. The lit- was rationally designed to sequester and stabilize endogenous erature reported that the inclusion of nanoclay in the polymer bone morphogenic protein-2 (BMP-2), therefore to achieve matrices significantly improve the mechanical and thermal 36–38 osteogenesis without applying its exogenous supraphysiological properties . Young’s modulus and EWC (Fig. 1b, c) were 31 32 dosage . In addition, the bone-forming sterosomes , the small analyzed and, with the increment of MMT, the mechanical molecular drug, i.e. phenamil, and siRNA knocking down noggin properties were remarkably improved. The Young’s modulus 33,34 expression were embedded in MeGC hydrogels to enhance was increased more than six times with the 4% MeGC-MMT bone regeneration. All the efforts were proven to be effective nanocomposite hydrogel compared with that of MeGC hydrogel in vivo in the calvarial defects of mice to a large extent, however, (<10 kPa), and it was around 30 kPa for the 1.5% MeGC-MMT none of the formulation alone has achieved fully satisfied results nanocomposite hydrogel. within the experimental time yet, typically 6 weeks. One of the MMT has been used as a filler in hard scaffolds, such as 24 23 phenomena in all those injectable chitosan-based hydrogels is chitosan–MMT–hydroxyapatite , gelatin–MMT–cellulose scaf- that they all bear a continuous microstructure without any pores folds and silk biomaterials for bone tissue formation. To the best (Figs. 2, 4), which probably limits their osteoconductive cap- of our knowledge, it is the first report to utilize MMT to create ability. The current MeGC-MMT formulation significantly interconnected microporous structure in chitosan-based in situ- 6 NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications Fold change Fold change Fold change NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 ARTICLE Blank MeGC MeGC + 1.5% MMT 80 10 NS 0 0 0 Fig. 8 Evaluation of in vivo bone regeneration. a Microcomputed tomography images of calvarial defects treated with hydrogels or left empty, 6 weeks post surgery, scale bar = 1 mm. b μCT quantification of bone regeneration in calvarial defects. Relative bone growth surface area, bone volume density −1 (BV/TV%), and the trabecular number (Tb.N., mm ). *p < 0.05, ***p < 0.001, and NS, no significance (ANOVA followed by Tukey’s post hoc test). Error bars indicate standard deviation (n = 6) Blank MeGC MeGC + 1.5% MMT 500 µm 100 µm 100 µm Fig. 9 Histological analysis of bone regeneration in calvarial defects, 6 weeks post surgery. Hematoxyline and eosin (H&E) staining (scale bar = 500 μm), and magnified images of H&E and Masson trichrome staining (red boxes represent the magnified areas, scale bar = 100 μm; arrowheads indicate the new bone formation) forming hydrogel just simply by intercalation chemistry (Supple- 1.5% w/v, to exhibit the strongest conductive ability in BMSCs, mentary Fig. 4). We confirmed the existence of pores in the wet confirmed with ALP and mineralization staining and their cor- state of hydrogel with low-vacuum SEM (Fig. 2) and also with responding quantification (Fig. 5), as well as specific gene frozen sections (Fig. 4) to avoid the ambiguity as the evaporation expression (Fig. 6). We tempted to culture BMSCs only in the of water in the hydrogel leading to the formation of pores. basal medium in the MeGC-MMT hydrogels, and osteogenic Further in vitro characterization of the combination of markers (ALP and Runx2) were upregulated on day 7 (Supple- MeGC and MMT showed that the optimal weight of MMT was mentary Fig. 6), indicating potential osteoinductive ability of this NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications 7 Blank MeGC MeGC + 1.5% MMT Blank MeGC MeGC + 1.5% MMT Blank MeGC MeGC + 1.5% MMT MT H&E Relative bone growth surface area % BV/TV % –1 Tb. N. (mm ) ARTICLE NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 composite. Further experiments of ectopic bone formation should fetal bovine serum (FBS, MT35010CV) was supplied by Mediatech Inc. (Manassas, VA, USA). The live/dead staining solution (L3224) and cDNA transcription kit be carried out to confirm this speculation. The most appealing (18080-051) were purchased from Invitrogen (Carlsbad, CA, USA). Trizol reagent character of this nanocomposite hydrogel is their ability in vivo to (15596018) and RNeasy Mini Plant Kit (74106) were supplied by Qiagen (Valencia, recruit native cells for de novo bone formation (Fig. 7). Satis- CA, USA). Pierce BCA Protein Assay Kit was obtained from Thermo Scientific factory results were obtained in the mouse calvarial defect model (23235, Rockford, IL, USA). The nude mice were purchased from Charles River Laboratories (Wilmington, MA, USA). All solvents and products were used as only by applying the material itself in the absence of cells or received. growth factors. This feature is of importance for translation to the clinical application, as cells were not necessary to be harvested Preparation of MeGC and MeGC-MMT nanocomposite hydrogels. In all, 2% and re-implanted back to the patients or expensive growth fac- (w/v) GC and GMA were mixed at 1:1 molar ratio of GMA to the amino groups in tors, such as BMP-2, to be applied. In addition, we have treated GC in Milli-Q water. The solution was adjusted to pH 9.0, and placed on a shaker the mouse calvarial defect with commercially available collagen at room temperature. After 40 h, pH was readjusted to 7.0, and the solution was 39 40 dialyzed with 50 kDa tubes against Milli-Q water for 16 h. After lyophilization and sponge and demineralized bone matrix (DBX) for 6 weeks as rehydration with phosphate-buffered saline (PBS), a 4% (w/v) MeGC solution was positive controls (Supplementary Fig. 7). The defects treated with −1 obtained. MMT was dispersed in Milli-Q water and yield 100 mg mL stock 1.5% MeGC-MMT hydrogels, collagen sponge, and DBX were dispersion. Various amount of MMT (w/v) was mixed with 4% (w/v) MeGC covered by new bone at 69, 23, and 26%, respectively. The BV/TV solution, and the hydrogels were cured under visible blue light with riboflavin as −1 and Tb. N. rose up to 46% and 6.2 mm for the 1.5% MeGC- the photoinitiator (final concentration 6 µM). The final concentration of MeGC is 2% (w/v) with various amount of MMT. MMT group, considerably higher compared with that of the collagen sponge group (6.4% and 2.7) or the DBX group (13.1% Characterization of MeGC-MMT nanocomposite hydrogels. A 400 μLof and 4.3). Compared with the commercially available materials, hydrogel mixture was cured for 80 s in a 48-well plate, and compressive modulus our newly developed MeGC-MMT hydrogel presents great was measured using a flat-ended indenter (1.6 mm in diameter) on an Instron potential for bone healing. Electro-Mechanical Testing Machins (Instron, Model 5564, Norwood, MA, USA). The surface charges and the exchange ability of MMT were The Young’s modulus was determined from the slope of linear portion of the 25,41,42 11 obtained stress-strain curve using a Poissons’ ratio of 0.25 . well studied for delivery purpose . Taking advantage of this, Hydrogels were equilibrated in PBS for 24 h and lyophilized for 16 h to obtain further development of this chitosan-based hydrogel system dry gels. EWC was calculated using the following equation specifically for bone tissue engineering is indispensable. The M  M combination and optimization of PS–RGD–MeGC or poly- w d EWC ¼ ; ð1Þ sulfonate MeGC and MMT hold great promise for bone regen- w eration. The critical defect fully filled with regenerated bone where M and M refer to the weight of wet and dry hydrogels, respectively. w d within experimental time will soon be achieved. The underlying The degradation of hydrogels was carried out for 42 days. Hydrogels were incubated at 37 °C in cell growth medium (DMEM, 10% FBS, and 1% P/S), which mechanism should be further studied to elucidate the strong was replaced every 7 days. At a pre-set time, hydrogels were lyophilized for weight osteoconductivity of this biomaterial and the type of tissue that measurement. The present residual weight of hydrogels was calculated using the replaces the hydrogel site with time progression needs to be following equation closely monitored as well in our future work. Our novel nanocomposite hydrogel system can be utilized not Residual dry gel weight% ¼ ´ 100%; ð2Þ only as 3D scaffolds for non-load-bearing injury site such as 0 cranial defects but also synthetic biological carriers for stem cells, where M and M refer to the weight of hydrogels at time 0 (hydrogels did not 0 t bioactive agents, or currently available tissue grafts for regen- undergo degradation) and t, respectively. Hydrogels were imaged in low vacuum using scanning electron microscopy erative medicine. We did not evaluate mechanical properties of with X-ray microanalysis (SEM/EDS, FEI Nova NanoSEM 230, Hillsboro, OR, the regenerated bone in this study, because the calvarial defect is USA) to observe the microstructure and chemical compositions. For Transmission small and the healing site is not load bearing. However, the Electron Microscope (TEM) analysis, hydrogels were embedded in sucrose-PVP further study of biomechanical properties will be needed in other and mounted on Cu EM grid after sectioning with glass 45 knife with 100 -nm fracture models such as large segmental defects. It is possible that thickness (Leica UC6/F6 −100 C), then imaged with JEM1200EX. The thermogravimetric analysis (TGA) was performed with the lyophilized hydrogels the hydrogel may not possess sufficient mechanical properties in o o o −1 from 50 C to 600 Cat 30 C min speed. The Powder X-ray Diffraction (XRD, the load-bearing bone defect sites. A promising alternative is to Bruker Corporation, Germany) pattern was collected from 2 to 10 (2θ range) with combine the hydrogel with solid particles or bone graft materials. ̊ a diffractometer using Ni-filtered CuKα X-ray radiation (λ = 1.5418 A). In conclusion, we have demonstrated that with the help of intercalation chemistry, proportional MMT mixing with MeGC 3D cell culture in MeGC-MMT nanocomposite hydrogels. BMSCs at a density 6 −1 leads to the formation of interconnected microporous structure, of 2 × 10 cells mL were mixed in MeGC-MMT dispersion. The hydrogel was cured by exposing 40 μL of the suspension to visible blue light (400–500 nm, which can promote native cell infiltration, proliferation, and −2 500–600 mW cm , Bisco Inc., Schaumburg, IL) in the presence of riboflavin as a in situ differentiation in the absence of any growth factors, small photoinitiator, (final concentration 6 µM). The resulting hydrogels were incubated molecular drugs, or genes. This biocompatible, bioactive, and in 1 mL of media accordingly. injectable nanocomposite material shows great promise being applied in a wide range of tissue regeneration. Cytotoxicity. Cytotoxicity of MMT in nanocomposite hydrogels was evaluated using alamar blue assay and live/dead staining (Invitrogen, Carlsbad, CA). Hydrogels with encapsulated BMSCs and MMT in various concentrations were Methods incubated at 37 C and 5% CO in fresh growth medium. Then the medium was Materials. Glycol chitosan (GC, ~100 kDa, 072-1581) was purchased from Wako replaced with 10% (v/v) alamar blue solution in growth medium at a pre- Chemicals USA, Inc. (Richmond, VA, USA). MMT (682659-500G), glycidyl determined time. After a 3-h incubation, the fluorescence intensity (F) of alamar methacrylate (GMA, 151238-100G), 1-ethyl-3-(3-dimethylaminopropyl)-carbodii- blue was measured at 585 nm with an excitation wavelength of 570 nm. For the mide (EDC, E1769-10G), tween-20 (p1379-500ML), p-nitrophenyl phosphate blank group, the 10% (v/v) alamar blue solution was added in an empty well and (N7653-100ML), β-glycerophosphate (G5422-100G), -ascorbic acid (A5960-25G), L incubated together with other samples. The relative cell viability (%) was calculated dexamethasone (D4902-25MG), Nitro Blue Tetrazolium (NBT, N5514-25TAB), 5- using the following equation bromo-4-chloro-3-indoxylphosphate (BCIP, B6149-50MG), alizarin red S (A5533), and ethylenediaminetetraacetic acid (EDTA, EDS-1KG) were supplied by Sigma- F  F s b Relative cell viability ¼ ´ 100%; ð3Þ Aldrich (St. Louis, MO, USA). The mouse bone marrow stromal cell line (BMSCs, F  F c b D1 ORL UVA [D1], D1 cell, CRL-12424) was obtained from American Type Culture Collection (ATCC, Manassas, VA, USA). High glucose Dulbecco’s Mod- where F , F , and F refer to the fluorescence intensity of the sample after incu- s c b ified Eagle’s Medium (DMEM, 11995-065), penicillin/streptomycin (100 U/mL, bation for 24 h, intensity of the corresponding sample without treatment, and the 15140122) were purchased from Life Technologies (Grand Island, NY, USA), and blank, respectively. 8 NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications NATURE COMMUNICATIONS | https://doi.org/10.1038/s41467-019-11511-3 ARTICLE ALP and alizarin red S staining and quantification. Gels were incubated in Data availability osteogenic media (growth medium was supplemented with 10 mM β-glyceropho- All relevant data are available within the article and Supplementary Information. The −1 sphate, 50 μgmL L-ascorbic acid, and 100 nM dexamethasone). At a pre- source data underlying Figs. 1B–D, 2, 3, 5B–C, and 66 and Supplementary Figs. 1, 4, 5, 6, determined time, gels were fixed in 10% formalin for 20 min, rinsed with PBS, and and 7B are provided as a Source Data file. Requests for other materials should be incubated in a solution consisting of NBT and BCIP stock solutions in an ALP addressed to the corresponding author. buffer (100 mM Tris, 50 mM MgCl , 100 mM NaCl, pH 8.5) for 2 h, at room temperature. The stained samples were observed with an Olympus SZX16 Ste- Received: 20 October 2018 Accepted: 1 July 2019 reomicroscope (Olympus, Tokyo, Japan). ALP expression appeared in blue. For the ALP activity assay, gels were rinsed with PBS, incubated in a lysis buffer (0.1% Tween-20 in PBS) at 4 °C for 5 min. ALP activity was determined colorimetrically using p-nitrophenyl phosphate as a substrate and measured at 405 nm. 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Mayer’s hematoxylin. 22. Mieszawska, A. J. et al. Clay enriched silk biomaterials for bone formation. Acta Biomater. 7, 3036–3041 (2011). Statistical analysis. Three independent experiments at least, unless otherwise 23. Hsu, S. H., Wang, M. C. & Lin, J. J. Biocompatibility and antimicrobial stated, were performed and data were presented as mean ± standard deviation. evaluation of montmorillonite/chitosan nanocomposites. Appl. Clay Sci. 56, Multiple comparisons were assessed using one-way or two-way analysis of variance 53–62 (2012). (ANOVA). The analysis of variances followed by Tukey’s post hoc test was 24. Katti, K. S., Katti, D. R. & Dash, R. Synthesis and characterization of a novel employed in this work and p < 0.05 was considered statistically significant. chitosan/montmorillonite/hydroxyapatite nanocomposite for bone tissue engineering. Biomed. Mater. 3, 12 (2008). 25. Hu, J. L. et al. Visible light crosslinkable chitosan hydrogels for tissue Reporting summary. 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Effect of organic modifiers on dynamic and static nanomechanical properties and crystallinity of intercalated clay-polycaprolactam nanocomposites. J. Appl. Polym. Sci. 105, 790–802 Open Access This article is licensed under a Creative Commons (2007). Attribution 4.0 International License, which permits use, sharing, 38. Yu, S. Z., Zhao, J. H., Chen, G., Juay, Y. K. & Yong, M. S. The characteristics of adaptation, distribution and reproduction in any medium or format, as long as you give polyamide layered-silicate nanocomposites. J. Mater. Process Technol. 192, appropriate credit to the original author(s) and the source, provide a link to the Creative 410–414 (2007). Commons license, and indicate if changes were made. The images or other third party 39. Yang, H. S. et al. Comparison between heparin-conjugated fibrin and collagen material in this article are included in the article’s Creative Commons license, unless sponge as bone morphogenetic protein-2 carriers for bone regeneration. Exp. indicated otherwise in a credit line to the material. If material is not included in the Mol. Med. 44, 350–355 (2012). article’s Creative Commons license and your intended use is not permitted by statutory 40. Townsend, J. M. et al. Effects of tissue processing on bioactivity of cartilage regulation or exceeds the permitted use, you will need to obtain permission directly from matrix-based hydrogels encapsulating osteoconductive particles. Biomed. the copyright holder. To view a copy of this license, visit http://creativecommons.org/ Mater. 13, 034108 (2018). licenses/by/4.0/. 41. Amsden, B. G., Sukarto, A., Knight, D. K. & Shapka, S. N. Methacrylated glycol chitosan as a photopolymerizable biomaterial. Biomacromolecules 8, 3758–3766 (2007). © The Author(s) 2019 10 NATURE COMMUNICATIONS | (2019) 10:3523 | https://doi.org/10.1038/s41467-019-11511-3 | www.nature.com/naturecommunications

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