Self-Assembled Injectable Nanocomposite Hydrogels Coordinated by

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Biological and Medical Applications of Materials and Interfaces

Self-Assembled Injectable Nanocomposite Hydrogels Coordinated by in situ Generated CaP Nanoparticles for Bone Regeneration Lijun Kuang, Xiaoyu Ma, Yifan Ma, Yuan Yao, Muhammad Tariq, Yuan Yuan, and Changsheng Liu ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.9b03173 • Publication Date (Web): 22 Apr 2019 Downloaded from http://pubs.acs.org on April 22, 2019

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ACS Applied Materials & Interfaces

Self-Assembled Injectable Nanocomposite Hydrogels Coordinated by in situ Generated CaP Nanoparticles for Bone Regeneration Lijun Kuang1,2, Xiaoyu Ma1,2, Yifan Ma1,2, Yuan Yao2, Muhammad Tariq1, Yuan Yuan1,2, 3*, Changsheng Liu1,2,3*

1.

Key Laboratory for Ultrafine Materials of Ministry of Education, East China University of Science and Technology, Shanghai 200237, PR China

2.School

of Materials Science and Engineering, East China University of Science and Technology, Shanghai 200237, PR China

3Engineering

Research Center for Biomaterials of Ministry of Education, East China

University of Science and Technology, Shanghai 200237, PR China Corresponding Authors *E-mail: [email protected]. *E-mail: [email protected]. ORCID Yuan Yuan: 0000-0001-7877-3175

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ABSTRACT Due to great similarity to the natural extracellular matrix (ECM) and minimally invasive surgeries, injectable hydrogels are appealing biomaterials in cartilage and bone tissue engineering. Nevertheless, undesirable mechanical properties and bioactivity greatly hamper its availability in clinic application. Here, we developed an injectable nanocomposite hydrogel by in situ growth of CaP nanoparticles (ICPNs) during the free radical polymerization of DMAEMA and HEMA matrix (PDH) for bone regeneration. The ICPNs are self-assembled by incorporation of Poly-L-glutamic acid (PGA) with abundant carboxyl functional groups during the formation of carboxyl-Ca2+ coordination and further CaP precipitation. Furthermore, the carboxyl groups of PGA could interact with the tertiary amines of DMAEMA fragments and thus improved the mechanical strength of hydrogel. Upon mixing solutions of DMAEMA and HEMA bearing PGA, Ca2+ and PO43-, this effective and dynamic coordination led to the rapid self-assembly of CaP NPs and PDH nanocomposite hydrogel (PDH/mICPN). The obtained optimal nanocomposite hydrogels exhibited suitable injectable time, enhanced tensile strength of 321.1 kPa and fracture energy of 29.0 kJ/m2, and dramatically facilitated cell adhesion and up-regulated osteodifferentation compared to hydrogels prepared by blending ex situ prefabricated CaP NPs (ECPNs). In vivo experiments confirmed the promoted osteogenesis, which brings striking contrast to pure PDH hydrogel. Additionally, the methacrylate groups on the monomers could easily functioned with aptamers and thereby facilitated recognition and capturing of BMSCs both in vitro and in vivo, and strengthened the bone regeneration. We believe that our conducted research about in situ self-assembled CaP nanoparticles-coordinated hydrogels will open a new avenue for bone regeneration in the future endeavors.

KEYWORDS: In situ generated CaP nanoparticles, Injectable nanocomposite hydrogels, Self-assembly, Apt-functionalized, Bone regeneration


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1. INTRODUCTION Due to the severe clinical challenges in the localized bone defects caused by trauma, infection or congenital diseases, recent years have witnessed rapid development of biomaterials designed from various perspectives for bone regeneration1-2. Among them, injectable biomaterials, especially bone-mimetic cement and polymeric hydrogels have attracted great attention of biomaterials researches. These biomaterials can replace implantation surgery with a minimally invasive injection method and can form any desired shape, to match irregular defects3-4. Unfortunately, majority of injectable materials currently available for clinical use, such as polymethylmethacrylate, calcium sulfate cement and calcium phosphate cement, suffer from poor biodegradation, overlong gelation time and high-polymerization isothermal phenomenon5. Therefore, a great deal of effort is being made to address these challenges towards the traditional injectable biomaterials. Injectable hydrogels, a free-flowing fluid before injection, followed by the spontaneously transformation into a semi-solid hydrogel once reaction initiated, have emerged as a promising platform for minimally invasive treatment in tissue regeneration6-7. This system possesses attractive superiorities, such as high water contents, similarity to the natural extracellular matrix (ECM) and porous framework for cell proliferation and differentiation8. In order to improve its mechanical strength and bioactivity, functional nanoparticles (NPs) or nanostructures, such as clay9-10, graphene oxide11-12, boron nitride13 and nanocellulose crystals14, have been incorporated into the polymer network of the hydrogel in previous researches15-16. Specifically, the prefabricated calcium phosphate (CaP) NPs (ex situ prepared) along with favorable osteoconductivity, resorbability and biocompatibility were considered to be the ideal fillers to fabricate nanocomposite hydrogels used for bone repairing17, while the in situ growth of calcium phosphate NPs (ICPNs) in hydrogels matrices fostered by soaking as-prepared hydrogels in CaCl2/NaH2PO4 solution were further developed to enhance the mechanical properties and efficient osteointegration of the hydrogels18. The CaP NPs-based system could improve the mechanical properties and osteogenic-related activities of hydrogels notwithstanding, whether the prefabricated CaP NPs or the in-situ strategy failed to meet homogeneous distribution and well combination with the polymer matrices, which were the prerequisites for the



mechanical stability. Moreover, the development of nanocomposite hydrogels by the abovementioned strategy often require a lengthy duration along with no-biocompatible pH, which are unsuitable for fabrication of injectable hydrogels or loading the bioactive cytokines19. Conclusively from the above discussion, developing an injectable hydrogel with in situ formed and well-distributed CaP NPs by one-pot method can be considered a more favorable solution for bone regeneration. Due to the protonation of the tertiary amines, the materials composed of dimethylaminoethyl methacrylate (DMAEMA) could swell in the physiological environment. It is considered as beneficial in nutrients delivery and enhancement of the local cell density20. Poly 2-hydroxyethyl methacrylate (poly-HEMA), has been considered as one of the tiny minority of synthetic biomaterials with attractive osteoinductive potential21. In current study, we have endeavored to fabricate a selfassembled, DMAEMA-HEMA based nanocomposite injectable hydrogel coordinated by in situ precipitation of CaP NPs (PDH/ICPN) via a facile, one-pot method to achieve successful bone repairing (Scheme 1). To achieve the in situ formation and uniform dispersion of CaP NPs in the matrix, Poly-L-glutamic acid (PGA), a biodegradable biomacromolecule with abundant carboxyl functional groups, was introduced to bind and serve as nucleation sites of calcium ion (Ca2+). Moreover, the carboxyl groups of PGA can also combine with the tertiary amines of DMAEMA fragments via electrostatic interaction to form a tight interface with the polymeric matrix. Furthermore, an aptamer specific to BMSCs named Apt19S were anchored on the hydrogel via covalent conjugation to endow it with the capacity to capture BMSCs22, as it is well known that stem cell recruitment and osteogenic differentiation are the key steps in the natural bone healing process17. The effects of different mass ratio of ICPNs on the surface topography, gelation time, swell behaviors and mechanical properties, as well as cell attachment, proliferation, osteogenic differentiation on the hydrogels were investigated thoroughly. Meanwhile, the function of the aptamers on recruitment of BMSCs in vitro and in vivo were also studied respectively. Finally, the nanocomposite hydrogels with the optimal NPs and aptamers amount were selected to investigate the in vivo osteogenesis in a rat femoral defect model.


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2. MATERIALS AND METHODS 2.1. Chemical Reagents. 2-(dimethylamino) ethyl methacrylate (DMAEMA, 2867-47-2, purity: ≥ 99.5%), 2-Hydroxyethyl methacrylate (HEMA, 868-77-9, purity: ≥ 99.5%), di (ethylene glycol) dimethacrylate (DEGDMA, 2358-84-1, purity: ≥ 99.0%), ammonium persulfate (APS, 7727-54-0, purity: ≥ 99.5%), N, N, N, Ntetramethylenediamine (TEMED, 110-18-9, purity: ≥ 99.5%) and Poly (glutamic acid) (PGA, 25513-46-6) were obtained from Aladdin Industrial Co., Ltd., (Shanghai, China). (NH4)2HPO4 (7783-28-0) and Ca(NO3)2·4H2O (10124-37-5, purity: 99.0%) were obtained from General-Reagent Co., Ltd., (Shanghai, China). Nucleic acid oligonucleotides with FITC and acrydite modified named as Apt19S: 5'- acryditeAGGTCAGATGAGGAGGGGGACTTAGGACTGGGTTTATGACCTATGCGTGFITC-3' were produced by Sangon Biotech Co., Ltd., (China) and directly used without further purification. Fetal bovine serum (FBS, 10%) and MEM Alpha Modification (1X) were obtained from Life Technologies (California, USA). 3-(4,5-dimethylthiazol-2-yl)2,5-diphenyltetrazolium bromide (MTT) were bought from Sigma-Aldrich (CA, USA) 2.2. Synthesis of PDH Hydrogel. The copolymeric hydrogels of DMAEMA and HEMA were prepared in aqueous solution at room temperature with DEGDMA as the crosslink agent. 5.992 mL of DMAEMA/HEMA solution (5/95, 25/75, 50/50 mol/mol; Aladdin) with 0.0125 mol% Di (ethylene glycol) dimethacrylate (DEGDMA, Aladdin), was added into a three-necked round bottom flask and then bubbled with nitrogen gas for 20 min. Pre-prepared APS (5mM) with equal volume of TEMED (2.5mM) were used as redox initiator system to initiate the polymerization. After thoroughly mixed, the solution was poured into several polypropylene syringes and polymerized at 37oC for 48 h. 2.3. Synthesis of PDH/mICPN Hydrogel. The composite hydrogel was synthesized by in situ precipitation method. The composite hydrogels with different CaP NPs concentration were prepared by varying the Ca(NO3)2 and (NH4)2HPO4 content from 5 to 25 wt% with a constant Ca/P ratio (1.67/1). PGA, Ca(NO3)2·4H2O and (NH4)2HPO4 were dissolved in distilled water separately, followed by the addition to the pre-mixed solution of the monomers as described in the synthesis of PDH. Then the solution was bubbled with nitrogen gas for 20 min. After the addition of APS (5



mM) and TEMED (2.5 mM) at the same volume, the solution was poured into syringes and polymerized at 37 oC for 48h. 2.4. Chemical Characterization. Fourier transform infrared (FT-IR) spectra (Nicolet 5700, Thermo) of the hydrogels were detected in the range of 800-4000 cm-1. The components of the nanocomposite hydrogels were also analyzed by solid state nuclear magnetic resonance (SSNMR) (Bruker 600M, Bruker Corporation). The data obtained were analyzed using MestReNova NMR software. Thermogravimetric analysis (TGA, STA449F3) was used for studying the thermal properties of PDH/mICPN hydrogels. The surface topology and the microstructure of the hydrogels were observed by scanning electron microscope (SEM, Hitachi, JPN). 2.5. Gelation Time. The gelation time of the hydrogels were measured by vial inverting method. 2 mL of the solution was poured into a vial and then incubated in a water bath at 37oC. When the vial is inverted vertically, no visible flow within 60 s is considered a criterion for gel formation. Three replicates were performed for all experiments. 2.6. Equilibrium Swelling Studies. Equilibrium swelling of the composite hydrogels were performed in phosphate buffered saline with different values of pH, such as 5.5, 6.5, 7.4 and 8.0. Hydrogel sheets (10 × 10 × 3.0 mm) were placed into a centrifuge tube containing 50 mL of buffer solution on a shaker (80 ± 1 rpm) at 37 oC for 48 hours. After removing the swollen hydrogels from PBS, water on the hydrogels was absorbed using filter paper followed weighing (Wswollen). Next, the samples were lyophilized to a constant weight (Wdry). The equilibrium swelling ratio was calculated using the following equation: 𝐸𝑞𝑢𝑖𝑙𝑖𝑏𝑟𝑖𝑢𝑚 𝑠𝑤𝑒𝑙𝑙𝑖𝑛𝑔 𝑟𝑎𝑡𝑖𝑜(%) =

𝑊𝑠𝑤𝑜𝑙𝑙𝑒𝑛 ― 𝑊𝑑𝑟𝑦 × 100 𝑊𝑑𝑟𝑦

Eq. 1. Calculation method of equilibrium swelling ratio. 2.7. Rheological Measurement. A rotational rheometer (Thermo Hakke, US) with a steel parallel plate of 10 mm radius was used to determine the rheological properties of the hydrogel. All rheological tests were carried out at 37 ° C unless otherwise stated.


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The time scan test was performed for 1 hour to study gel formation kinetics. The changes in storage modulus (G') and the loss modulus (G") over time were recorded at a constant frequency of 1 Hz and a strain of 1%. For the frequency sweep test, the hydrogels sheets (d = 20mm, h = 1mm) were placed onto the parallel plate, and the strain was maintained at 1.0%. 2.8. Differential Scanning Calorimetry (DSC) Measurement. The glasstransition temperature (Tg) of PDH/12.5ICPN and PDH/12.5ECPN were determined by DSC (modulated DSC2910, 1090B, USA). All samples were immersed in neutral PBS at 37℃ and swelled for 48 h to reach the equilibrium swelling. These swollen samples were placed in specialized crucibles and then sealed. The thermal analysis of the swollen samples was carried out at a heating rate of 2℃/min from room temperature to 90℃. 2.9. Tensile Mechanical Testing. Uniaxial tensile test of dumbbell-shaped samples standardized to the JIS-K6251-7 size (x-axis (length) = 12.0 mm, y-axis (width) = 2.0 mm, and 2.4 mm, z-axis (thickness) = 2.4 mm were carried out using a commercial tensile machine (SANS CMT2503). The sample was drawn along the x-axis at a constant speed of 100 mm min-1 measured according to JIS-K6251. 2.10 rBMSCs culture and seeding. The rBMSCs were extracted from the femurs and tibias bone marrow of SD rats and cultured with α-MEM medium containing 10% FBS and 1% penicillin/streptomycin (Gibco) at 37°C, 5% CO2/95% air, and saturated humidity. After autoclaving at 120 °C, the hydrogels were placed in a 24-well plate (1 sample per well) under sterile conditions. After pre-soaking with α-MEM for 12 hours, the cell suspension was lifted from the top of the stent (1 mL/well). 2.11. In vitro Cellular Biocompatibility of the Composite Hydrogels. Cytotoxicity and cell proliferation on the hydrogels were assessed by the MTT assay. At every point in time, 100 mL of MTT solution (5 mg/mL) was added and incubated at 37 ℃ for 4 h to form formazan crystals, which were later dissolved in DMSO for 15 min. The optical density (OD) value at 492 nm was detected by a microplate reader (SPECTRA max 384, Molecular Devices, USA). These experiments were performed in triplicate.



Cell adhesion assay: each sheet was added 105 BMSCs were added into each sheet. After culturing for 24 h, the hydrogel was fixed with 4% paraformaldehyde for 10 min at 4oC. Next, FITC-Phalloidin (5 mg/mL) was added to the fixed cells and incubated at 37oC for 45 min and then incubated with DAPI (5 mg/mL) for 10 min at room temperature, for cytoskeleton and nuclear staining, respectively. Cell morphologies were observed by laser scanning confocal microscopy (LSCM, Leica TCS SP8 STED 3X, Germany) and the relative cell area was quantitatively analyzed using Image Pro 5.0 software. For studying cell morphology, 105 cells were seeded on the surface of each sheet. After 24 h incubation, the samples were washed twice with PBS and fixed by the addition of 2.5% glutaraldehyde. After washed twice with PBS and dehydrated at increasing concentrations of ethanol, samples were observed using SEM (S-3400, Hitachi, JPN). 2.12. Cell Osteogenic Differentiation and Gene Expression. For analysis of ALP activity, 2 x 104 cells were seeded on each scaffold and co-cultured with osteogenic induction medium in 24-well cell-culture plates. After 7d and 14d of culture, the cell culture medium was aspirated and washed twice with PBS, 1 mL of 1% NP-40 lysate was added to each well, and the cells were lysed by shaking at 37 ° C for 90 minutes. The total protein amount of the cell lysate was determined by BCA method. Transfer 50 μL of cell lysate per well to a 96-well plate, add 200 μL of ALP working solution for 15 minutes, and measure the absorbance at 405 nm (OD value). The 405 nm OD value/ (total protein amount * incubation time) of each group was calculated in 405 nm OD value / min / mg protein, and the obtained data were normalized to the blank control data to obtain relative ALP activity data of each group. All experiments were repeated three times. ALP staining: rBMSCs were inoculated on the hydrogels at a density of 2×104 cells/well, and BCIP/NBT (BCIP/NBT Alkaline Phosphatase Color Development Kit) alkaline phosphatase was used after 14 days of culture. Under the catalysis of alkaline phosphatase, BCIP will be hydrolyzed to produce a highly reactive product that will react with NBT to form insoluble dark blue to bluish-purple NBT-formazan. The bluer product, the higher the activity of alkaline phosphatase, and vice versa.


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The expressions of osteogenic gene were determined by a real-time quantitative polymerase chain reaction (RT-qPCR) system (Bio-Rad, Hercules, USA). The gene expressions of Runx2 (runt-related transcription factor 2), OCN, OPN and Col I were measured after 7d and 14d of culture, and blank plates containing no hydrogels were employed as control groups. All experiments were performed in quadruplicates. 2.13. Preparation and Fluorescence Imaging of PDH/12.5ICPN-nApt Hydrogels. Typically, after mixing the monomers, Ca(NO3)2 and (NH4)2HPO4 solutions and the aptamer solutions were added into the mixture at different concentration (0.2, 2.0, 5.0 nmol/mL) in total and stirred at room temperature for 20 min. Hydrogels either functionalized with FAM-labeled acrydite-modified Apt19S aptamers, immersed in FAM-labeled acrydite-modified Apt19S aptamers solutions or without aptamers were imaged through Confocal Microscope with a 10 × objective lens and a 488 nm argon laser. 2.14. In vitro Migration Assay of BMSCs. The ability of recruiting BMSCs in PDH/12.5ICPNaP-nApt group was assessed using a transwell system. Briefly, 104 cells re-suspended in 100 μL of serum-free DMEM were placed in the upper chamber, and the hydrogels incorporated with different concentration of aptamers with 200 μL of DMEM containing 10% FBS were placed in the lower chamber for 24 h. Then the cells were fixed with 4% paraformaldehyde and stained with crystal violet. The results were exhibited as mean number of cells per filed ± standard deviation. 2.15. Implantation of the Hydrogels in Rat Femur Defect. Animal experiment was approved by the Animal Research Committee of Sixth People’s Hospital, Shanghai Jiao Tong University School of Medicine. 7-week-old Male Sprague−Dawley (SD) rats (National Tissue Engineering Center, Shanghai, China) with average weight about 300g were used. The rats were randomized into three groups: PDH, PDH/12.5ICPN hydrogel, PDH/12.5ICPN-5Apt hydrogel. Prior to surgery, rats were anesthetized with pentobarbital (Nembutal 3.5 mg/100 g) and then the femurs of the rat were shaved (Fig. S5e). Two critical size defects were made by using a 3 mm diameter trephine (Surgident, Korea). Subsequently, the hydrogels were injected into the femurs defect, and the skin was closed. 2.16. In vivo Stem Cell Recruitment Flow Cytometry. Immunofluorescence staining and flow cytometry were used to evaluate the recruitment of stem cells in each



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group. The percentage of BMSCs was calculated by immunofluorescence staining of three cell surface antigens (CD29, CD44, CD45) and flow cytometry as previously reported23-24. Briefly, the hydrogels injected in the bone were transferred into 1.5 mL centrifuge tubes and further incubated for 10 min. After that, the pancreatic enzymes were used for digestion followed by adding 1 mL of PBS to wash the suspended cells. Each group of the cell suspension was divided into two equal volumes (one for immunostaining and the other as unstained control), and stained with monoclonal antiCD29, anti-CD44, and anti-CD45 antibodies (Biolegend, San Diego, USA) on ice for 30 min and finally obtained 500 μL of cell suspension for flow cytometry. When the data were processed, the unstained control group was used as the negative screening standard. CD29 and CD44 double positive cells were selected first, and then CD45 negative cells were selected from these cells, and the percentage of these cells in the total number of cells was calculated. 2.17. Micro-computed Tomography (micro-CT) Measurements. Samples taken at 2, 4, 8 and 12 weeks after surgery were immersed in 4% neutral paraformaldehyde solution (PFA) and dried at 37 °C for 72 hours. Micro-CT measurements at Shanghai Synchrotron Radiation Facility (SSRF) BL13W line station was applied for 3D rotation scanning of the femur samples25, using a monochromatic beam with an energy of 26 keV with a sample-to-detector distance of 0.8 m and 9 µm of scanning accuracy. 2.18. Hard Tissue Section Staining. The femurs samples of week 4, 8, 12 time points were dehydrated stepwise using gradient alcohols (from 75% to 100%) and embedded in polymethylmethacrylate (PMMA). The section for each specimen was cut, and then manually polished to a thickness of about 40 µm. At last, the sections were stained with van Gieson’s picrofuchsin for observation of new bone formation. 3. RESULTS and DISCUSSION 3.1. In Situ Formation and Characterization of Injectable PDH/mICPN Nanocomposite

Hydrogels.

Previous

studies

have

developed

CaP-based

nanocomposite hydrogels driven by the physicochemical reactions between polymer chains and NPs26-27. In this study, we developed a novel strategy to prepare injectable nanocomposite hydrogels by one-pot in situ self-assembled NPs with improved


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mechanical strength and bioactivity. Poly (DMAEMA-HEMA) (PDH), as a hydrogel proved to be highly pH-sensitive and potentially osteoinductive was chosen as the polymeric matrix. As shown in Fig. S1, the hydrogel was synthesized by APS/TEMEDinduced free-radical polymerization with DMAEMA, HEMA and DEGDMA served as monomers and cross-linker respectively. Three hydrogels PDH-x were obtained with the mole ratio of DMAEMA/HEMA: 5/95, 25/75 and 50/50. With the increase of DMAEMA ratio in PDH, the PDH-50/50 tended to be macroscopically heterogeneous (Fig. S2a) due to the disproportional monomer reactivity28. The successful synthesis of PDH-5/95 and PDH-25/75 were characterized by FT-IR and 13C NMR spectroscopy. As shown in Fig. S2b, wide absorption peaks for the -OH group in HEMA at around 3351 cm-1 and for the -CN group in DMAEMA at around 1244 cm-1 could be observed, which demonstrated the copolymerization of PDH. 13C-NMR spectroscopy results (Fig. S2c) showed that peaks at 𝛿 43.75, 75.85 and 181.64 ppm were corresponding to methyl, methylene and ester of DMAEMA moiety in the polymer backbone, while peaks around 55.48-71.34 ppm could be served for methylene of HEMA moiety. The ratio of DMAEMA and HEMA was calculated from

13C

NMR data suggesting an effective

control over the polymer synthesis process via comparing the theoretical estimation and experimental analysis. Furthermore, a plate-to-plate rheometer experiment revealed a low storage modulus (G’) of both PDH-5/95 and PDH-25/75 at the initial stage, following with an accelerated gelling reaction. Consistent with the results (Fig. S2e) obtained by vial inverting method, Fig. S2d showed the gelation point of PDH-5/95 and PDH-25/75 were about 450 s and 300 s respectively. The gelation time became shorter with the increasing of DMAEMA to HEMA ratio because of the higher reactivity and weaker intermolecular hydrogen bonding of DMAEMA28. Considering the uniform phase composition (Fig. S2a) as well as a more rapid gelation behavior, PDH-25/75 hydrogel was chosen to composite with NPs and defined as PDH hydrogel hereafter. The nanocomposite hydrogels with the optimized ratio of DMAEMA/HEMA were then prepared with incorporating Ca(NO3)2 and APS/TEMED initiators solution into the precursor solution composed by the monomers, PGA and (NH4)2HPO4. As for the fabrication process of PDH/ICPN nanocomposite hydrogels illustrated in Fig. 1a, Ca2+ complexed with the carboxyl functional groups on PGA chains by electrostatic attraction and further combined with PO43- to form CaP NPs in situ during the



copolymerization of the DMAEMA-co-HEMA hydrogels. For comparison purpose, ex situ formed CaP NPs (ECPNs) (Ca/P = 1.67) were prepared by chemical precipitation and mixed into monomers solution to form the hydrogel with the mass ratio of CaP NPs controlled as 12.5% (referred to as “PDH/12.5ECPN” hereafter) (Fig. 1b). Incorporation of NPs in both hydrogels could be confirmed by the new peak at 1128 cm-1 corresponding to P-O stretching in PO43- in Fourier transform infrared (FTIR) spectra (Fig. 1c). The TGA profiles of the hydrogels (Fig. S3) displayed two distinct weight losses from room temperature to 300oC and 300oC to 500oC, which were associated with the amino groups in DMAEMA29 and lateral chains decomposition of both DMAEMA and HEMA. Notably, PDH/12.5ICPN exhibited a higher decomposition temperature than pure PDH and PDH/12.5ECPN, indicating a stronger electrostatic interaction between ICPNs and the polymer network. Moreover, as shown in Fig. 1h, the thermal behaviors of PDH/ICPN and PDH/ECPN hydrogels by DSC indicated that PDH/ICPN displayed a single glass transition temperature (Tg) at 60 oC, while PDH/ECPN exhibited a dual Tg at 45oC and 75oC respectively, which revealed a phase separation of PDH matrix in PDH/ECPN

28.

Similar heterogeneity, primarily

irregular pattern could also be observed in the macroscopic illustration of PDH/ECPN (Fig 1i). In further analysis of microstructural level, the incorporation of CaP via different strategies were found to exert distinct effect on the surface morphology of PDH hydrogel. As shown in Fig 1j, CaP NPs with about 50 nm in diameter were homogeneously distributed throughout the observed areas of PDH/ICPN, whereas severe agglomeration of CaP NPs was significant in PDH/ECPN system, suggesting an essential phase separation in accordance with previous results. In order to further demonstrate the distribution of CaP NPs in the hydrogel, we analyzed the crosssectional morphology of PDH, PDH/12.5ICPN and PDH/12.5ECPN by SEM-EDS as shown in Fig. 1k and Fig. S4. These findings confirmed that PDH/ICPN possessed well phase stability of hydrogel matrix and improved distribution of CaP in the nanohydrogel.


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Fig. 1. (a) Formation of PDH/mICPN by in situ PGA-mediated precipitation process. (purple arrows: positive charges of DMAEMA) (b) Formation of PDH/12.5ECPN by ex situ as-fabricated method. (c) FT-IR spectra of pure PDH and nanocomposite hydrogels. (d) Time-dependent (at strain of 1%) oscillatory shear rheology and (e) Storage modulus as a function of time of various hydrogels. (f) Gelation time of the different hydrogels. PDH/12.5ECPN had the longest gelation time about 360s,



whereas the gelation time of PDH/12.5ICPN was much shorter about136s. (g) A demonstration of the injectability and moldability of PDH/12.5ICPN. (“*”: p < 0.05, indicating a significant difference between the two groups). (h) DSC curves of PDH/12.5ICPN and PDH/12.5ECPN. (i) Microscopic observation of all the hydrogels. (j) SEM images of the lyophilized hydrogels (scale bars represent 20 and 2 µm for the upper and the lower rows respectively). (k) SEM morphology and EDS analysis of the cross section of PDH/12.5ICPN (scale bar: 2 µm).

In the system of ICPNs reinforced injectable hydrogels, three features could be characterized: the copolymerization of the hydrogel matrix, the in-situ growth of nanoparticles on the polymer chain, and intertwining ICPNs formation and gelation profiles throughout the process. In order to explore the effect of NP on the hydrogel formation, we first investigated the changes in dynamic rheological properties. As shown in Fig. 1d-e, the complex viscosity (𝜂*) and storage modules (G’) of both PDH/ICPN and PDH/ECPN as a function of time at 37oC, were low before 120 s and started to rise with a prolonged time sweep. In this early stage for PDH/ICPN, since the particle size of ICPNs was small and the polymerization had not yet started, the system exhibited a relatively low viscosity and resultant high liquidity, which endowed the hydrogel a desirable injectability for operation. As the reaction proceeded, polymer cross-linking network started to generate, and the nanoparticles in PDH/ICPN gradually grew up and tightly combined with the DMAEMA fragments in the polymer network, thereby leading to a rapid increase of 𝜂* and G’. In contrast, because of weak interaction between the ECPNs and the polymer network and poor dispersion of the ECPNs in the matrix, there was sluggish increase of the 𝜂* and G’ for PDH/12.5ECPN. More remarkable differences lied in the gelation points (identified as the time points when 𝜂* started dramatically increasing) and setting rates (directly proportional to the slopes of the curves30) of the hydrogels. It was found that incorporation of NPs by in situ method dramatically accelerated the gelling reaction. Specifically, the gelation point tested was about 2.5 min for PDH/12.5ICPN, while PDH/12.5ECPN and PDH hydrogel were 5.3 min and 5.0 min respectively. Meanwhile, PDH/12.5ICPN exhibited a quick rise of 𝜂* and reached 1.5 × 103 Pa·s after 7 min setting, but 𝜂* of PDH/12.5ECPN and pure PDH, reached about 0.26 × 103 Pa·s only. Furthermore, the gelation time of the hydrogels determined via inverting method was shown in Fig. 1f, where the gelation time of PDH/12.5ICPN was about 136 s, while PDH/12.5ECPN had a significantly longer


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gelation time: about 360 s. These results were well agreed with the previous results obtained by the 𝜂*-t and G’-t curves. Based on the aforementioned rheological properties, PDH/12.5ICPN could be facilely injected, quickly adapt to the desired shape (Fig. 1g) and render to a rapid remolding. Such rheological behaviors and injectable properties made the PDH/ICPN hydrogels of potential interest for bone repair via minimal surgery. The PDH/ICPN nanocomposite hydrogels with various content of CaP NPs ingrowth were further investigated. By varying the mass of PGA, Ca(NO3)2 and (NH4)2HPO4 in precursors, we obtained three formulas of nanocomposite hydrogels (defined as “PDH/mICPN”, m referred to the CaP content (5.0, 12.5 and 25.0 wt%) in PDH) along with the ratio of Ca/P and PGA/CaP kept at 1.67 at 0.25 respectively. By the TGA analysis shown in Fig. 2a and Table S1, the actual CaP contents for PDH/5ICPN, PDH/12.5ICPN and PDH/25ICPN composite hydrogels were determined to be 6.4%, 16.5% and 22.5%, respectively. In macroscopic observations of PDH/mICPN (Fig 2e), all groups showed homogeneous dispersion of ICPNs in the gel matrices. As illustrated in the SEM images (Fig. 2f), with the increase of CaP NPs, PDH/mICPN still kept spongy structure as pure PDH. Well-distributed ICPNs could be further observed in higher magnification of SEM (Fig. 2f, lower rows), and PDH/12.5ICPN showed larger size but less density of NPs than PDH/5ICPN. Notably, when the CaP NPs amount was over 25%, introduced CaP tended to form larger clusters, which might give rise to severe aggregation.



Fig. 2. (a) TGA profiles of pure PDH and PDH/mICPN. Compositing amount of CaP (5.0%, 12.5% and 25.0%) were in accordance with the weight loss of composite hydrogels. (b) Time-dependent (at strain of 1%) oscillatory shear rheology and (c) Storage modulus as a function of time of various hydrogels. (d) Gelation time of PDH and PDH/mICPN. (e) Microscopic observation of the hydrogels. (f) SEM images of the lyophilized hydrogels (scale bars represent 20 and 2 µm for the upper and the lower rows respectively).

As a pH-sensitive hydrogel composed of DMAEMA, swelling behaviors could be controlled by the local physiological environment to assist cargo delivery and cell concentrating. Hence, swelling ratios of hydrogels were measured as a function of composition and pH, and physiological pH range from 5.5 to 8.0 was chosen in view of the extracellular pH level during wound healing31. As shown in Fig. S4, the swelling ratio of nanocomposite hydrogels at pH 5.5 reduced to some extent when compared to


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pure PDH, nevertheless, all groups of PDH/mICPN showed significant pH-responsive swelling once bringing down pH level, which revealed that the in-situ formation of ICPNs would not impair the swelling behaviors of PDH hydrogels. Next, the dynamic rheological properties of PDH/mICPN during the hydrogel formation were analysis. As shown in Fig. 2b-c, the initial viscosity (𝜂*) and storage modules (G’) of all groups in PDH/mICPN remained low at initial stage for some time and started to rise with a prolonged time sweep as before mentioned. Obviously, PDH/5ICPN, the content of ICPNs was too low to affect the 𝜂*. The gelation time of the hydrogels further determined via inverting method were shown in Fig. 2d where gelation time dramatically decreased with the incorporation of NPs by in situ method. When the ratio of ICPNs reached 12.5% and 25.0%, the gelation time was all about 136 s without distinct difference. We believed such gradient addition of CaP could immensely affect a series of mechanical and biological properties of PDH/mICPN nanocomposite hydrogels. 3.2. In Situ Formation of ICPNs Improve the Mechanical Properties of Nanocomposite Hydrogels. As shown in Fig. 3b, pure PDH was weak as expected and failed to maintain its original outline after compression. In contrast, for PDH/12.5ICPN, the weaker bonds between PGA-CaP NPs or between PGA-CaP NPs and polymer chains formed by electrostatic interaction are broken firstly. This process dissipates a large quantity of energy results in that the crosslinked network of the polymer was not destroyed during the deformation process and exhibited significantly improved stability and could almost fully recover after excessive compression. To verify the effect of the in-situ generated CaP NPs on the hydrogels’ mechanical properties, the rheological behaviors measured by storage modulus (G’) and the loss modulus (G’’) were analyzed. Strain-dependent oscillatory rheology (Fig. 3c) demonstrated that the ICPNs could dramatically enhance the G’ and G”. Particularly, PDH/12.5ICPN exhibited a broad linear viscoelastic region and a great antishear ability, merely network failure at high strains over 100%, indicating a satisfactory toughness32. In contrast, PDH/12.5ECPN hydrogel was observed as discontinues at strain of about 20%, which implied a decrease toughness owing to the absence of tight combinations between PDH network and ECPNs. The dynamic oscillatory rheology of the hydrogels was also measured and shown in Fig. 3d. It confirmed the role of CaP NPs played on the G’. Specifically, G’



of PDH/12.5ICPN was 109.7 ± 12.6 kPa, nearly sevenfold and eightfold as that of PDH/12.5ECPN and pure PDH hydrogel. Also, the increased ICPNs mass fraction markedly enhanced the hydrogel mechanical property (Fig. 3e). Under the uniaxial tensile testing, the PDH/5ICPN, PDH/12.5ICPN and PDH/25ICPN reached the high tensile strengths of 152.4, 321.1 and 183.2 kPa, respectively, which were calculated as 1.9, 4.0 and 2.3 times higher than that of pure PDH (78.7 kPa) and 1.1, 2.4 and 1.4 times higher than that of PDH/12.5ECPN (132.4 kPa) (Fig. 3f). The fracture energy increased as 5.3-fold from 542.7 ± 0.2 J m-2 for pure PDH to 2897.6 ± 0.3 J m-2 for PDH/12.5ICPN (Fig. 3g). The dramatically mechanical improvement of in-situ forming PDH/12.5ICPN nanocomposite hydrogels were on account of the following factors. Firstly, as a typically inorganic fillers, ICPNs with more uniform distribution, can thereby enhance the strength of the hydrogels15. More importantly, compared to the preformed ECPNs, the electrostatic interaction between unreacted carboxyl groups of PGA chains and the tertiary amines of DMAEMA part resulted in a higher binding affinity between ICPNs and PDH polymer network. On the other hand, CaP NPs, as a filler, could inhibit the interaction and collisions of the reactant monomers, which would reduce the degree of cross-linking of the polymer chains to a certain extent, and thereby weaken the mechanical properties of the material. Therefore, from this viewpoint, PDH/12.5ICPN realized a balance between filler enhancer and crosslinking density of PDH, and thereby exhibited the improved optimum mechanical properties. But the PDH/25ICPN might incorporate too much ICPNs and exerted negative effect on the curing of hydrogels (Fig. 3a).


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Fig. 3. In situ formation of ICPN NPs enhanced the mechanical properties of the PDH/mICPN nanocomposite hydrogels. (a) Schematic illustration of different combination modes between NPs and polymer networks. (b) Optical images of the PDH and PDH/12.5ICPN nanocomposite hydrogels under excessive compression (scar bar: 1cm). (c) Strain-dependent (ω=10 rad s-1) (solid lines represent G’ and dashed line represent G”) and (d) frequency-dependent (at strain of 1%) oscillatory rheological analysis results. (e) Stress-strain curves of the nanocomposite hydrogels with different CaP content. (f) Fracture stress and (g) Fracture energy of PDH, PDH/mICPN and PDH/12.5ECPN hydrogels (“*”: p < 0.05, indicating significant differences).

3.3. In Vitro Cellular Adhesion and Osteogenic Differentiation of BMSCs Cultured on the Composite Hydrogels. Cell attachment is the first step of the interaction between cells and materials, which exerts direct regulations on the following cell behaviors33. SEM and fluorescence stain were used to detect cell adhesion on various hydrogels. As illustrated in Fig. 4a, BMSCs cultured on the pure PDH and PDH/12.5ECPN for 24 hours exhibited low cell density. Furthermore, the cells were seen as the fibroblastic in appearance, with an elongated and aligned morphology.



Obviously, an increase in BMSCs density was observed on PDH/mICPN hydrogels, especially on PDH/25ICPN. Interestingly, for PDH/12.5ICPN, cells were observed as more elongated and well-spread with significant outstretched filopodia extensions and lamellipodia protrusions as shown in the lower row of Fig. 4a, indicating that the adding of ICPNs was beneficial to cell attachment. As shown in Fig. 4b, the variation tendency of cell density observed in fluorescence pictures was similar to that of SEM images. Quantitatively, the cell density on PDH/12.5ICPN was about 165 cells/mm2 whereas the average density was about 60 cells/mm2 on PDH/12.5ECPN (Fig. 4d). In addition, there was no significant difference on the cell area among the samples (Fig. 4e). As shown in Fig. 4c, the cell viability remained over 95% even for 25.0% CaP, indicating a good biocompatibility of the nanocomposite hydrogels.

Fig. 4. Cell attachment, morphology and proliferation of BMSCs on the nanocomposite hydrogels. (a) Cell morphology on the hydrogels after 24h. The upper and the lower rows show typical lowmagnification and high-magnification SEM micrographs, respectively. (scale bar: 50 µm and 10 µm,


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respectively). (b) Fluorescent images of cell morphology on the hydrogels by confocal microscope with FITC and DAPI staining. (scale bar: 100µm). (c) The initial cell adhesion on the hydrogels after culturing for 24 h was further determined by MTT assay. (d) ImageJ analysis of cell number and (e) cell area on PDH and PDH/mICPN hydrogels. The result showed an up-trend with increased CaP amount (“*”: p < 0.05, indicating a significant difference between the two groups).

ALP activity, an early stage marker of stem cells osteogenic differentiation 34, was evaluated to investigate the effect of NPs on the osteogenic differentiation of BMSCs. It could be intuitively seen from ALP staining (Fig. 5a), that the addition of CaP obviously enhanced ALP activity, and the most intensive ALP staining could be observed in response to PDH/12.5ICPN. The results of quantitative assay (Fig. 5b) of 7 and 14d exhibited a light distinction as compared to trend of the framed ALP staining. When compared to the pure PDH hydrogel, PDH/5ICPN exhibited a slightly higher ALP activity, while PDH/12.5ICPN revealed a dramatically higher ALP activity than others. Notably, PDH/25ICPN showed slightly lower ALP activity. We further quantified the expression levels of the osteogenic genes such as Col, OCN, OPN and Runx2 was evaluated using RT-qPCR analysis to confirm the tendency in ALP activity. The results of Fig. 5c-f indicated that the PDH/12.5ICPN showed the most significant up-regulated expression in all four genes. The ALP and qPCR results together clearly indicated that PDH/12.5ICPN had a higher osteogenic potential than that of pure PDH or the hydrogel filled by ex situ CaP NPs.



Fig. 5. (a) Alkaline phosphatase activity was stained with BCIP/NBT (20×) at day 7. ALP activity of PDH and PDH/mICPN hydrogels were measured by soluble substrate p-nitrophenylphosphate at day 7 (b) and 14 (c). Osteogenic gene expressions of (d) Col I, (e) OCN, (f) OPN, and (g) Runx2 after 14 days’ culture on different hydrogels (“*”: p < 0.05, indicating a significant difference between the two groups).

It is well-accepted that many chemical cues as well as physical cues, such as matrix stiffness and surface topography exert dramatic impacts on cell behaviors35 and determine stem cell fate36. Appropriate nanodisplaced topography, nanoscale disorder surface, significantly increased cellular adhesion and osteoblastic differentiation rather than planar surface or highly random surface37. A fast stress relaxation matrix could promote cell shape changes, proliferation and differentiation, and thereby the bone matrix formation38. The work by Jiandong Ding et al.39 reported that stiff hydrogel with large space nanotopography surface promoted osteogenesis significantly. Furthermore, Dalby et al.37 found that cells on smooth and flat surfaces exhibited poor cell adhesion and osteoblastic differentiation, and the surfaces with a certain disorder nanodisplaced topography significantly increased osteospecific differentiation. Based on these previous studies, it can be hypothesized that the desirable adhesion and increased


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osteospecific differentiation of BMSCs on PDH/12.5ICPN should be attributable to a high Young’s modulus (109.7 ± 2.6 kPa), ideal nanotopography surface and viscoelastic property (Fig. 2e). 3.4. Preparation and Cell Recruitment of PDH/12.5ICPN-nApt hydrogels in Vitro and in Vivo. It is widely accepted that recruitment of BMSCs to the defect site is essential for bone repairing process40. Many recent studies mostly focus on loading bone morphogenetic proteins or chemotactic factors which have the functions of promoting BMSCs recruitment and bone regeneration with as-prepared scaffolds23. But the problems such as poor efficiency, side effects of high doses and expensive costs make it difficult to use them in clinic. In this study, in order to facilitate the in situ BMSCs recruitment, a BMSCs-specific aptamer Apt19S which is high efficient and stable was chosen and chemically functionalized with acrydite and FITC at its 5’-end and 3’-end for polymerization and labeled respectively (Fig. 6a). As shown in Fig. S5a, PDH/12.5ICPN-nApt hydrogels by chemical reaction exhibited a much stronger fluorescence intensity than other groups, indicating a successful incorporation of Apt into the hydrogel network 41. The addition of Apt exerted no effect on the macroscopic appearances and gelation time (Fig. S5b-c). Fig. S5d intuitively showed that PDH/12.5ICPN-5Apt could be formed in any mold with desired shape by the use of a syringe and the addition of Apt had no influence in the fluidity of the hydrogel. In vitro binding of BMSCs with PDH/12.5ICPN and PDH/12.5ICPN-nApt was examined via the transwell assay. As shown in Fig. 6b, the higher the concentration of the material-complexed aptamer, the greater the number of cells recruited. The statistical data (Fig. 6c) showed that there were 167 ± 5 BMSCs in the PDH/12.5ICPN5Apt hydrogel, while only 10 ± 3 BMSCs were in the hydrogels without aptamer. This experiment indicated that the aptamer-hydrogel did favor BMSCs capture in a typical concentration-dependent manner. Furthermore, we chose rat femoral platform implantation as in vivo model to determine the in vivo BMSCs recruitment capacity of the nanocomposite hydrogels by immunostaining and flow cytometry. Pure PDH, PDH/12.5ICPN and PDH/12.5ICPN5Apt were injected into the defect site respectively (Fig. S5e). As shown in Fig. 6e, after 4 days of implantation, the percentage of BMSCs recruited by the PDH/12.5ICPN5Apt group was 3.85 ± 0.13%, which was significantly higher than the percentage of



the PDH group (0.84 ± 0.20%) and the percentage of the PDH/12.5ICPN group (0.85 ± 0.34%). The results confirmed the positive effect of the complex aptamer on stem cell recruitment42.

Fig. 6. Synthetic and characterization of PDH/12.5ICPN-nApt hydrogels. (a) Reaction scheme of hydrogel synthesis. Apt19S-aptamer sequences were modified at their 3’ ends with fluorescence and 5’ ends with Acrydite and for conjugation during PDH polymerization. (b) Light microscopy image of the transwell assay of BMSCs toward PDH/12.5ICPN and PDH/12.5ICPN-nApt (scale bar: 500µm). (c) Statistical data of the transwell assay. (d) Flow cytometric profiles of sort gates of CD44-positive, CD29-positive and CD45-negtive cells (criteria defining BMSCs). (e) quantitative histogram of the percentage of recruited BMSCs. PDH/12.5ICPN-5Apt group exhibited the highest recruited BMSCs percentages.

3.5. Bone Regenerative Efficiency in A Rat Femoral Condyle Defect Model. Given the observed enhancement of osteoblastic differentiation of PDH/12.5ICPN hydrogels and excellent stem cell recruitment ability of PDH/12.5ICPN-5Apt hydrogel, the following experiments were conducted to evaluate their bone regenerative


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efficiency in vivo using a rat femoral condyle defect models. Micro-CT results (Fig. 7a) showed the difference of bone repair process in each group at 2, 4, 8 and 12 weeks intuitively. Obviously, the group of PDH had lower volume of bone in the defect area than that in other experimental groups after implanting 2 and 4 weeks, only a small amount of new bone tissues were formed in the area near the ulnar defect. PDH/12.5ICPN with ICPNs could prominently facilitate the osteogenesis. Indeed, PDH/12.5ICPN-5Apt could enhance the osteogenesis, especially in the initial 4 weeks. As the time passes, all the groups could finish bone regeneration. The quantitative data also proved the results as shown in Fig. 7b and Fig. 7c.

Fig. 7. (a) Micro-CT images at week 2, 4, 8 and 12, including 3D views, longitudinal section and cross sections (scale bar: 5mm). Micro-CT-quantified histograms of (b) bone volume fraction and (c) trabecular thickness (“*”: p < 0.05, indicating a significant difference between the two groups).

To further evaluate new bone formation and mineralization in different phases of bone defect healing, undecalcified specimens stained with van Gieson’s picro fuchsin were analyzed. Representative histological images were shown in Fig. 8. As anticipated, the newly formed bone in PDH/12.5ICPN group was observed more than that in pure



PDH at 4 weeks. At 8 weeks, there was more bone trabecula in three groups contrast with 4 weeks. At 12 weeks, most mature bone tissue could be observed in PDH/12.5ICPN and PDH/12.5ICPN-5Apt. Comparatively, PDH/12.5ICPN with 5Apt could strongly promote new bone formation.

Fig. 8. Histological images of undecalcified sections counter-stained with Van Gieson at week 4, 8 and 12. Green arrows: implanted hydrogels, yellow arrows: new bones, blue arrows: Haversian osteons (Scale bar = 1 mm).

As we know, in a typical bone regeneration process, a rapid initiation of efficient stem cell recruitment and adhesion, and ensuing osteodifferentiation are critical for final bone regeneration and bone remodeling. The injectable nanocomposite PDH/ICPN hydrogel developed in this study can be injected into the defect site via minimally invasive surgery, and quickly fill the irregular shape due to the good fluidity. PDH/ICPN, especially with Apt19S, could realize a higher efficiency in recruiting stem cells, and provided sufficient mechanical support, up-regulated osteoinductivity and synchronized degradation rate with bone formation. Therefore, PDH/12.5ICPN5Apt showed optimum osteogenic activity both in vitro and in vivo.

4. Conclusion


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In this study, we developed a self-assembled, injectable and tunable nanocomposite hydrogel coordinated by in situ generated CaP nanoparticles. The in situ self-assembled hydrogels was of the high strength and great toughness via one-step copolymerization of DMAEMA and HEMA and PGA-mediated in situ precipitation of ICPN NPs. The uniform distribution of ICPN NPs in the PDH matrix not only endowed excellent injectability, enhanced mechanical properties, but also rendered them more conductive to cell adhesion, proliferation and osteogenesis. Furthermore, the addition of the Apt19S to composite hydrogel could accelerate recruitment BMSCs after implantation into the defect site. We believe that our research findings provide the strong evidence an effective method has been developed to incorporate the CaP NPs into polymer networks. By designing the injectable hydrogels, it was expected to offer new thoughts on the establishment of biomaterials used for bone repairing in the future endeavors.

Scheme 1. Schematic illustration of the fabrication of the injectable PDH/ICPN hydrogel for bone regeneration. (a) Formation of the polymer network from HEMA and DMAEDA via Michael addition reaction (purple arrows: positive charges of DMAEMA). (b) The in situ self-assembly of



CaP NPs around the –COOH groups of the PGA via –COO- – Ca2+ coordination. (c) The interaction between the PGA and the polymer network via electrostatic attraction.

Acknowledgements: The authors wish to express their gratitude to the financial supports from the National Natural Science Foundation of China for Innovative Research Groups (No. 51621002), National Key R&D Program of China (2018YFC1105700) and the National Natural Science Foundation of China (No. 31771040) and Leading talents in Shanghai in 2017. Supporting Information. Fabrication and characterization of pure PDH hydrogel; TGA profiles of pure PDH, PDH/12.5ICPN and PDH/12.5ECPN; Mass swelling ratios of various hydrogels in pH 5.5, 6.5, 7.4 and 8.0 phosphate buffers; Characterization of PDH/12.5ICPN-nApt; Procedures of animal surgery and hydrogel implantation. REFERENCES (1) Seliktar, D. Designing cell-compatible hydrogels for biomedical applications. Science 2012, 336 (6085), 1124-1128. (2) Zhang, L.; Xia, K.; Lu, Z.; Li, G.; Chen, J.; Deng, Y.; Li, S.; Zhou, F.; He, N. Efficient and Facile Synthesis of Gold Nanorods with Finely Tunable Plasmonic Peaks from Visible to Near-IR Range. Chem. Mater. 2014, 26 (5), 1794–1798. (3) Burdick, J.; Mauck, R.; Gerecht, S. To Serve and Protect: Hydrogels to Improve Stem Cell-Based Therapies. Cell Stem Cell 2016, 18 (1), 13-15. (4) Nasim, A.; Ali, T.; Jorge Alfredo, U.; Mohsen, A.; Luiz E, B.; Chaenyung, C.; Gulden, C.-U.; Mehmet R, D.; Nicholas A, P.; Ali, K. 25th anniversary article: Rational design and applications of hydrogels in regenerative medicine. Adv. Mater. 2014, 26 (1), 85-124. (5) María, V. R.; Eduardo, R. H. Bioceramics: from bone regeneration to cancer nanomedicine. Adv. Mater. 2011, 23 (44), 5177-5218. (6) Liu, M.; Zeng, X.; Ma, C.; Yi, H.; Ali, Z.; Mou, X.; Li, S.; Deng, Y.; He, N. Injectable hydrogels for cartilage and bone tissue engineering. Bone Res. 2017, 5, 17014. (7) Zhang KY; Jia ZF; Yang BG; Feng Q; Xu X; Yuan WH; Li XF; Chen XY; Duan L; Wang DP; Bian LM; Adaptable Hydrogels Mediate Cofactor-Assisted Activation


Page 28 of 33

Page 29 of 33 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


of Biomarker-Responsive Drug Delivery via Positive Feedback for Enhanced Tissue Regeneration. Adv. Funct. Mater. 2018, 5, 1800875. (8) Burdick, J. A.; Prestwich, G. D. Hyaluronic acid hydrogels for biomedical applications. Adv. Mater. 2011, 23 (12), H41-H56. (9) Martin, J. E.; Patil, A. J.; Butler, M. F.; Mann, S. Guest‐Molecule‐Directed Assembly of Mesostructured Nanocomposite Polymer/Organoclay Hydrogels. Adv. Funct. Mater. 2011, 21 (4), 674–681. (10) Dawson, J. I.; Oreffo, R. O. C. Clay: new opportunities for tissue regeneration and biomaterial design. Adv. Mater. 2013, 25 (30), 4069-4086. (11) Shin, S. R.; Aghaei-Ghareh-Bolagh, B.; Dang, T. T.; Topkaya, S. N.; Gao, X.; Yang, S. Y.; Jung, S. M.; Oh, J. H.; Dokmeci, M. R.; Tang, X. S. Cell-laden microengineered and mechanically tunable hybrid hydrogels of gelatin and graphene oxide. Adv. Mater. 2013, 25 (44), 6385-6391. (12) Linlin, W.; Masataka, O.; Masaki, T.; Akinori, S.; Shu, S.; Yasuhiro, I.; Takuzo, A. Magnetically induced anisotropic orientation of graphene oxide locked by in situ hydrogelation. Acs Nano 2014, 8 (5), 4640. (13) Nagarajan, S.; Belaid, H.; Pochat-Bohatier, C.; Teyssier, C.; Bechelany, M. Design of Boron Nitride/Gelatin Electrospun Nanofibers for Bone Tissue Engineering. ACS Appl. Mater. Inter. 2017, 9 (39), 33695−33706. (14) Mckee, J. R.; Appel, E. A.; Seitsonen, J.; Kontturi, E.; Scherman, O. A.; Ikkala, O. Healable, Stable and Stiff Hydrogels: Combining Conflicting Properties Using Dynamic and Selective Three-Component Recognition with Reinforcing Cellulose Nanorods. Adv. Funct. Mater. 2014, 24 (18), 2706-2713. (15) Mehrali, M.; Thakur, A.; Pennisi, C. P.; Talebian, S.; Arpanaei, A.; Nikkhah, M.; Dolatshahi-Pirouz, A. Nanoreinforced Hydrogels for Tissue Engineering: Biomaterials that are Compatible with Load-Bearing and Electroactive Tissues. Adv. Mater. 2016, 29 (8), 1603612. (16) Zhang, K.; Lin, S.; Feng, Q.; Dong, C.; Yang, Y.; Li, G.; Bian, L. Nanocomposite hydrogels stabilized by self-assembled multivalent bisphosphonatemagnesium nanoparticles mediate sustained release of magnesium ion and promote in-situ bone regeneration. Acta Biomater. 2017, 64, 389-400. (17) Anderson, J. M.; Patterson, J. L.; Vines, J. B.; Amjad, J.; Gilbert, S. R.; HoWook, J. Biphasic peptide amphiphile nanomatrix embedded with hydroxyapatite



nanoparticles for stimulated osteoinductive response. Acs Nano 2011, 5 (12), 94639479. (18) Xu, B.; Zheng, P.; Gao, F.; Wang, W.; Zhang, H.; Zhang, X.; Feng, X.; Liu, W. A Mineralized High Strength and Tough Hydrogel for Skull Bone Regeneration. Adv. Funct. Mater. 2016, 27 (4), 1604327. (19) Zhang, K.; Feng, Q.; Xu, J.; Xu, X.; Tian, F.; Yeung, K. W. K.; Bian, L. Self‐ Assembled Injectable Nanocomposite Hydrogels Stabilized by Bisphosphonate‐ Magnesium (Mg2+) Coordination Regulates the Differentiation of Encapsulated Stem Cells via Dual Crosslinking. Advanced Functional Materials 2017, 27 (34), 1701642. (20) You, J.-O.; Rafat, M.; Almeda, D.; Maldonado, N.; Guo, P.; Nabzdyk, C. S.; Chun, M.; LoGerfo, F. W.; Hutchinson, J. W.; Pradhan-Nabzdyk, L. K.; Auguste, D. T. pH-responsive scaffolds generate a pro-healing response. Biomaterials 2015, 57, 22-32. (21) Winter, G. D.; Simpson, B. J. Heterotopic bone formed in a synthetic sponge in the skin of young pigs. Nature 1969, 223 (5201), 88-90. (22) Hu, X.; Wang, Y.; Tan, Y.; Wang, J.; Liu, H.; Wang, Y.; Yang, S.; Shi, M.; Zhao, S.; Zhang, Y.; Yuan, Q. A Difunctional Regeneration Scaffold for Knee Repair based on Aptamer-Directed Cell Recruitment. Adv. Mater. 2017, 29 (15), 1605235. (23) Lin, D.; Chai, Y.; Ma, Y.; Bing, D.; Yuan, Y.; Liu, C. Rapid initiation of guided bone regeneration driven by spatiotemporal delivery of IL-8 and BMP-2 from hierarchical MBG-based scaffold. Biomaterials 2019, 196, 122-137. (24) Yang, Z.; Xie, J.; Zhu, J.; Kang, C.; Chiang, C.; Wang, X.; Wang, X.; Kuang, T.; Chen, F.; Chen, Z.; Zhang, A.; Yu, B.; Lee, R. J.; Teng, L.; Lee, L. J. Functional exosome-mimic for delivery of siRNA to cancer: in vitro and in vivo evaluation. J Control Release 2016, 243, 160-171. (25) Ma, Y.; Zhang, W.; Wang, Z.; Zi, W.; Xie, Q.; Niu, H.; Han, G.; Yuan, Y.; Liu, C. PEGylated poly(glycerol sebacate)-modified calcium phosphate scaffolds with desirable mechanical behavior and enhanced osteogenic capacity. Acta Biomater. 2016, 44, 110-124. (26) Bing, X.; Zheng, P.; Fei, G.; Wei, W.; Zhang, H.; Zhang, X.; Feng, X.; Liu, W. A Mineralized High Strength and Tough Hydrogel for Skull Bone Regeneration. Adv. Funct. Mater. 2017, 27 (4), 1604327.


Page 30 of 33

Page 31 of 33 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


(27) Vo, T. N.; Tatara, A. M.; Santoro, M.; Jj, V. D. B.; Leeuwenburgh, S. C.; Jansen, J. A.; Mikos, A. G. Acellular Mineral Deposition within Injectable, Dual-Gelling Hydrogels for Bone Tissue Engineering. J. Biomed. Mater. Res. A 2017, 105 (1), 110. (28) Martín-Gomis, L.; Cuervo-Rodriguez, R.; Fernández-Monreal, M. C.; Madruga, E. L.; Fernández-García, M. Monomer reactivity ratios and glass‐transition temperatures of copolymers based on dimethyl amino ethyl methacrylate and two structural hydroxy‐functional acrylate isomers. J. Poly. Sci. A Poly. Chem. 2003, 41 (17), 2659-2666. (29) Roy, D.; Guthrie, J. T.; Perrier, S. Synthesis of natural–synthetic hybrid materials from cellulose via the RAFT process. Soft Matter 2007, 4 (1), 145-155. (30) Farrar, D. F.; Rose, J. Rheological properties of PMMA bone cements during curing. Biomaterials 2001, 22 (22), 3005-3013. (31) Young Cheol, W.; Soo Seog, P.; Subieta, A. R.; Brennan, T. J. Changes in tissue pH and temperature after incision indicate acidosis may contribute to postoperative pain. Anesthesiology 2004, 101 (2), 468. (32) Xing, R.; Liu, K.; Jiao, T.; Zhang, N.; Ma, K.; Zhang, R.; Zou, Q.; Ma, G.; Yan, X. An Injectable Self‐Assembling Collagen–Gold Hybrid Hydrogel for Combinatorial Antitumor Photothermal/Photodynamic Therapy. Adv. Mater. 2016, 28 (19), 36693676. (33) Kang, H.; Dsh, W.; Yan, X.; Jung, H. J.; Kim, S.; Lin, S.; Wei, K.; Li, G.; Dravid, V. P.; Bian, L. Remote Control of Multi-Modal Nanoscale Ligand Oscillations Regulates In Vivo Stem Cell Adhesion and Differentiation. Acs Nano 2017, 11 (10), 9636. (34) Xue, W.; Krishna, B. V.; Bandyopadhyay, A.; Bose, S. Processing and biocompatibility evaluation of laser processed porous titanium. Acta Biomater. 2007, 3 (6), 1007-1018. (35) Sean J, M.; Allan C, S. Stem cells and niches: mechanisms that promote stem cell maintenance throughout life. Cell 2008, 132 (4), 598-611. (36) Dalby, M. J.; Nikolaj, G.; Oreffo, R. O. C. Harnessing nanotopography and integrin-matrix interactions to influence stem cell fate. Nat. Mater. 2014, 13 (6), 558. (37) Dalby, M. J.; Gadegaard, N.; Tare, R.; Andar, A.; Riehle, M. O.; Herzyk, P.; Wilkinson, C. D.; Oreffo, R. O. The control of human mesenchymal cell



differentiation using nanoscale symmetry and disorder. Nat Mater 2007, 6 (12), 9971003. (38) Chaudhuri, O.; Gu, L.; Klumpers, D.; Darnell, M.; Bencherif, S. A.; Weaver, J. C.; Huebsch, N.; Lee, H. P.; Lippens, E.; Duda, G. N.; Mooney, D. J. Hydrogels with tunable stress relaxation regulate stem cell fate and activity. Nat. Mater. 2016, 15 (3), 326-34. (39) Ye, K.; Wang, X.; Cao, L.; Li, S.; Li, Z.; Yu, L.; Ding, J. Matrix Stiffness and Nanoscale Spatial Organization of Cell-Adhesive Ligands Direct Stem Cell Fate. Nano Letters 2015, 15 (7), 4720. (40) Undale, A. H.; Westendorf, J. J.; Yaszemski, M. J.; Khosla, S. Mesenchymal Stem Cells for Bone Repair and Metabolic Bone Diseases. Mayo Clin. Proc. 2009, 84 (10), 893-902. (41) Chen, N.; Zhang, Z.; Soontornworajit, B.; Zhou, J.; Wang, Y. Cell adhesion on an artificial extracellular matrix using aptamer-functionalized PEG hydrogels. Biomaterials 2012, 33 (5), 1353-1362. (42) Enam, S. F.; Krieger, J. R.; Saxena, T.; Watts, B. E.; Olingy, C. E.; Botchwey, E. A.; Bellamkonda, R. V. Enrichment of endogenous fractalkine and anti-inflammatory cells via aptamer-functionalized hydrogels. Biomaterials 2017, 142, 52-61.


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Graphical Abstract

We developed an PDH-based injectable nanocomposite hydrogel coordinated by in situ growth of CaP nanoparticles (ICPNs) during the free radical polymerization of DMAEMA and HEMA matrix (PDH) for bone regeneration. The ICPNs are self-assembled by incorporation of Poly-L-glutamic acid (PGA) with abundant carboxyl functional groups during the formation of carboxyl-Ca2+ coordination and further CaP precipitation. Furthermore, the carboxyl groups of PGA could interact with the tertiary amines of DMAEMA fragments and thus improved the mechanical strength of hydrogel. Upon mixing solutions of DMAEMA and HEMA bearing PGA, Ca2+ and PO43-, this effective and dynamic coordination led to the rapid self-assembly of CaP NPs and PDH nanocomposite hydrogel (PDH/ICPN). The nanocomposite PDH/ICPN hydrogel can be injected into the defect site via minimally invasive surgery, and quickly fill the irregular shape due to the good fluidity and shows optimum osteogenic activity both in vitro and in vivo.


Self-Assembled Injectable Nanocomposite Hydrogels Coordinated by

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