Demineralized Bone Scaffolds with Tunable Matrix Stiffness for

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Biological and Medical Applications of Materials and Interfaces

Demineralized Bone Scaffolds with Tunable Matrix Stiffness for Efficient Bone Integration Qingxia Hu, Mengying Liu, Guobao Chen, Zhiling Xu, and Yonggang Lv ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.8b08668 • Publication Date (Web): 31 Jul 2018 Downloaded from http://pubs.acs.org on August 1, 2018

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ACS Applied Materials & Interfaces

Demineralized Bone Scaffolds with Tunable Matrix Stiffness for Efficient Bone Integration Qingxia Hu†, ‡, #, Mengying Liu†, ‡, #, Guobao Chen†, ‡, Zhiling Xu†, ‡, Yonggang Lv*, †, ‡ †

Key Laboratory of Biorheological Science and Technology (Chongqing University), Ministry of

Education, Bioengineering College, Chongqing University, Chongqing, 400044, PR China ‡

Mechanobiology and Regenerative Medicine Laboratory, Bioengineering College, Chongqing

University, Chongqing, 400044, PR China KEYWORDS: demineralized bone matrix scaffold, matrix stiffness, mesenchymal stem cells, rabbit femoral condyle defect model, bone integration

ABSTRACT: As a biophysical cue, matrix stiffness can decide the stem cell fate. However, most methods to construct three dimensional (3D) scaffolds may change 3D microstructure while altering their mechanical properties. In this study, demineralized bone matrix scaffolds with different compressive modulus (66.06 ± 27.83 MPa (High), 26.90 ± 13.16 MPa (Medium) and 0.67 ± 0.14 MPa (Low)) were constructed by controlling the decalcification duration (1 h, 12 h and 5 d), respectively. The pore size and porosity has no significant difference between scaffolds before and after decalcification. Cell experiments indicated that the Low scaffolds could promote the osteogenic differentiation of bone marrow mesenchymal stem cells (MSCs) in vitro. Rat subcutaneous implantation experiments further demonstrated that the Low scaffolds could

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efficiently improve the cell infiltration, deposition of collagen fibers and positive osteocalcin and osteopontin expression of endogenous cells as well as angiogenesis. Finally, rabbit femoral condylar defect experiments proved that the Low scaffolds could significantly promote the bone repair and integration and stromal cell derived factor-1α/CXC chemokine receptor signal pathway was essential for the stiffness-mediated bone repair. These investigations provided a novel method for fabricating 3D bone grafts with different stiffness, which is also of great significance for studying the effect of stiffness on the biological behavior of MSCs in 3D.

1. INTRODUCTION In native tissues, the complex biophysical and biochemical signals from extracellular matrix (ECM) can affect stem cell behaviors.1,2 It is a promising strategy to fabricate ECM matrix scaffolds with optimal biophysical and biochemical properties by mimicking stem cell niche for bone repair. Matrix stiffness, as a biophysical cue of ECM, has been considered as an important role in deciding stem cell fate, which not only influences the cell morphology, cell phenotype, cell proliferation and focal adhesions, but also defines the lineage differentiation of stem cells.3-5 Many researchers fabricated substrates or scaffolds with different stiffness to study the influence of different stiffness on the stem cell fate. Engler et al.6 firstly demonstrated that human bone marrow mesenchymal stem cells (MSCs) can differentiate into neurons, myoblasts, or osteoblasts on collagen-coated polyacrylamide (PA) hydrogel with stiffness similar to that of the brain, muscle, or collagenous bone. However, majority of previous work linking matrix stiffness to cell differentiation focused on two-dimensional (2D) or quasi-three-dimensional (quasi-3D) model systems. Studies of 3D scaffolds which can represent the real situation in vivo are still limited.


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In order to meet this challenge, researchers have made efforts to find methods of generating 3D scaffolds with different stiffness, such as changing the matrix density of natural proteins in different positions,7 using synthetic polymers with different cross-linking densities8 or different concentrations,5 and changing the ratios of mixed matrices with varied elasticity.9,10 For example, Tse and Engler10 used PA hydrogels to form matrix containing a physiological gradient of 1.0 ± 0.1 kPa/mm by photopolymerization. However, these methods change the structural properties of 3D scaffolds (e.g., morphology, porosity and pore size) while changing the stiffness. The structural properties of 3D scaffolds can also affect the differentiation of stem cells.11 Therefore, decoupling the effects of matrix stiffness and structural properties of 3D scaffold was indispensable and may provide insight into how the individual matrix mechanics regulates the biological behavior of stem cells. Recently, our group fabricated scaffolds with different stiffness but same microstructure by coating decellularized cancellous bone with collagen/ hydroxyapatite (HA) mixture in different collagen ratios.12-14 However, it is still needed to figure out the role of the small amount of collagen in matrix stiffness-mediated osteogenic differentiation of stem cells, as collagen is one of the major components of bone matrix and can promote stem cell adhesion. Without changing collagen content or adding any ingredient, 3D demineralized bone matrix (DBM) scaffold with tunable mechanical properties while maintaining the same 3D microstructure was further manufactured in this study. Early studies have shown that DBM scaffold is an appropriate material for bone repair and regeneration, because of its suitable osteoinductive, osteoinductive and osteogenic effects.15,16 The natural porous structure of the DBM scaffold provides sufficient surface area and internal space for ECM secretion. The DBM has good biocompatibility and biodegradability, low antigenicity, and no cytotoxicity.16 Decalcification which can remove HA, the crystalline


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component of bone, represents one of the most important and indispensable steps in histotechnology.17 Furthermore, decalcification removes the dense mineral components which wrap around bone morphogenesis related factors, consequently these factors can be released smoothly and exert osteogenic activity.18 The ability of DBM to enhance osteogenesis of MSCs is believed to be due to the interaction of osteoprogenitors with these matrix incorporated osteoinductive factors, which can induce MSCs into osteoblasts. However, previous studies always neglected that the mechanical property of DBM is also changed during this process, which may also affect on the osteogenic differentiation of MSCs. Guo et al.19 indicated that decalcification can decrease the mechanical property of bone with the processing time increase. In order to form different stiffness matrices and investigate the role of mechanical factor alone, this study constructed novel 3D DBM scaffolds with different matrix stiffness via treatment of ethylenediaminetetraacetic acid disodium salt (EDTA-2Na) for different time. Moreover, in order to get rid of the effect of exposed osteoinductive factors due to decalcification, the DBM scaffolds were immersed in guanidine hydrochloride (Gdn·HCl) to get successfully inactivated.20 Decellularized cancellous bone scaffold that has been proved to have poor osteoinductive capability was used as the control group.21 After seeded with rat MSCs and cultured for 3 weeks, the cell compatibility and the ability of osteogenic differentiation of the constructs were analyzed in vitro. The effects of DBM scaffolds with different stiffness on the cell infiltration, collagen deposition, and osteogenic differentiation of endogenous cells were also evaluated by subcutaneous test. Rabbit femoral condylar bone defect repaired with these scaffolds was further performed to validate our hypothesis that the scaffold with optimal matrix stiffness could recruit endogenous stem cells to the defect site and improve bone repair and regeneration. In addition, the relationship between the stromal cell derived factor-1α (SDF-1α)/CXC chemokine receptor


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(CXCR4) signal axis and the recruitment of endogenous stem cells and inflammation reaction in rabbit bone repair were discussed. The strategy proposed in this study provides a safer, more reliable and economical alternative implant to bone healing.

2. EXPERIMENTAL SECTION 2.1. Rat MSCs isolation and culture. Primary rat bone marrow MSCs were isolated from the tibias and femurs of 4-week-old Sprague-Dawley (SD) rats weighing 130 ± 10 g (Animal Center, Daping Hospital, Army Medical University, Chongqing, China) as described previously.13,22 MSCs were isolated by Percoll density gradient centrifugation (1.073 g/L) and grown at 37°C and 5% CO2 in low glucose Dulbecco’s modified Eagle’s medium (Gibco, USA) supplemented with 10% fetal bovine serum (Gibco, USA) and penicillin/streptomycin. MSCs were selected on the basis of adhesion and proliferation on tissue culture plastic substrate. Medium exchange was performed every three days. When the cells grew to near confluency, they were washed with phosphate-buffered saline (PBS) and detached from flasks using 0.25% trypsin/0.01% EDTA2Na. These cells were subcultured similarly as first passage cells (P1) until reached 75%-85% confluency and then trypsinized similarly to yield second passage cells (P2). MSCs in passages 3-4 were used for the following experiments. 2.2. Preparation of DBM scaffold with different stiffness. The decellularized bone scaffolds were prepared by a porcine cancellous bone as previously described.13 Scaffold with 10 mm in diameter and 5 mm in thickness was prepared for cell and subcutaneous experiments. Scaffold with 5 mm in diameter and 5 mm in thickness was built for animal experiment. The decellularized bone scaffolds with high density (0.688 ± 0.076 mg/mm3) scaffolds were selected for experiments.


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The selected high density scaffolds were decalcificated in 12% EDTA-2Na (8 mL per scaffold, Sangon Biotech, China) for 1 h, 12 h and 5 d, respectively. All scaffolds were inactivated according to the method of Sampath and Reddi’s study.23 Firstly, samples were extracted (30 ml/g of matrix) with Gdn·HCl (4 M Gdn·HCl, 50 mM Tris, pH 7.4) at 4°C for 16 h, and then extensively rinsed in distilled water. The inactivated DBM scaffolds were stored at 4°C after sterilized with 60Co γ irradiation (25 kGy) (Allanace, Southwestern Radiation Research Center, Army Medical University, Chongqing, China). 2.3. Characterization of 3D DBM scaffold. 2.3.1. Calcium content quantification. Scaffolds decalcified at different time points were solubilized in 6 M HCl. Residual calcium content was determined following the manufacturer’s protocol (Calcium Assay Kit, Nanjing Jiancheng Bioengineering Institute, China). Briefly, same amount of deionized water, calcium standard solution and test solution were respectively added to three 10 mL test tubes. Then 1 mL of methyl thymol blue reagent and 2 mL alkaline solution were added to the three test tubes. At the wavelength of 610 nm, the absorbance of each tube was determined by UV-Vis spectrophotometer. Distilled water was used as a blank to zero the spectrophotometer. Conversion formula of calcium ion concentration: Calcium content (mmol/g) = (measured OD value-blank OD value)/(standard OD value-blank OD value) × standard concentration (2.5 mm)/protein concentration (g/L) 2.3.2. Mechanical testing of scaffolds. Mechanical properties of scaffolds decalcificated for 0 h (Control),1 h (High), 12 h (Medium) and 5 d (Low) were measured by using a tabletop uniaxial compression testing instrument (Instron-E1000, USA). All samples were soaked into phosphate buffer saline (PBS) overnight. Prior to analysis, excess fluids on the surface of scaffolds were removed. Compressive testing was performed at a compressive speed of 2 mm/min at room


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temperature. The elastic modulus was obtained in the strain range from 2% to 5% of the stressstrain curve. The experiment was performed for three samples in each group independently. 2.3.3. Structural characterization of scaffolds. Micro-computed tomography (µ-CT) was performed on scaffolds before decalcification (Control) and after decalcification for 5 d (Low) using a modification of a previously used protocol.13 After wetted with PBS overnight, the scaffolds were scanned at a resolution of 19 µm by vivaCT 40 system (Scanco medical AG vivaCT40, Switzerland). Special software µ-CT v6.1 of Scanco medical AG was used to visualize the 2D X-ray section images of the layered scaffolds and obtain 3D images. The microstructural parameters, such as porosity and pore size were used for comparison. 2.4. Cell seeding and cultivation in scaffolds. Scaffolds were seeded with bone marrow MSCs using a modification of method previously described.13 In brief, after resuspension and centrifugation, 20 µL (containing 4×105 cells) of MSCs suspension were seeded into the scaffolds for 3 times, each with 15 min intervals. Scaffolds were flipped every 15 min for 1 h to achieve uniform cell distribution. After adding 6 mL medium, seeded scaffolds were maintained in an incubator at 37°C, 5% CO2, and 95% humidity. 2.5. Live/dead assay. After harvest at day 3, day 7, and day 21, a half of one scaffold in each group was washed in PBS and stained with 2 µM calcein AM (staining for live cells) and 4 µM ethidium homodimer-1 (staining for dead cells) working solution for 30 min in dark, as indicated by the manufacturer’s instructions (Live/dead Viability/Cytotoxicity Kit, Molecular Probes, USA). The images were taken with a fluorescence microscope (Olympus IX71, Japan). 2.6. Alkaline phospatase (ALP) assay. After cultured 14 and 21 days in the medium, the constructs formed by co-culture of scaffolds and MSCs were washed with PBS and fixed with 4% paraformaldehyde. ALP production was then identified by staining with the BCIP/NBT ALP


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color development kit (Beyotime, Shanghai, China) for 1 h. The samples were washed thoroughly with PBS, followed by taking of images. 2.7. Ethics statement. Animal experiments were performed at the experimental animal center of Daping Hospital, Army Medical University. All animal handling were conducted in compliance with the principles of national community guidelines for the care and use of laboratory animals. The animal experiments and surgical procedures were submitted to and approved by the Animal Care and Use Committee of Chongqing University. 2.8. Subcutaneous implantation. Half of one scaffold in each group was subcutaneously implanted into SD rats (male, body weight 230 ± 5 g), respectively. The animals were anaesthetized with 7% chloral hydrate (0.5 mL/100 g). Scaffolds were inserted into the subcutaneous pockets at the left corner and right corner, respectively. After 28 days, animals were sacrificed by cervical dislocation and the implants were retrieved. 2.9. Rabbit femoral condyle defect model. 30 adult New Zealand white rabbits (weight 2.773 ± 0.042 kg) were randomly divided into six groups for assessment at 1, 2 and 3 months time points: (A) defect repair with pure decellularized bone scaffold (Control); (B) defect repair with low stiffness scaffold (Low); (C) defect repair with medium stiffness scaffold (Medium); (D) defect repair with high stiffness scaffold (High); (E) self repair group (Self repair); (F) normal bone group (Normal). Before surgery, all rabbits were anesthetized with pentobarbital sodium (30 mg/kg) by ear vein injection. After fixing, a longitudinal incision with 2-3 cm was made in the left distal femur via the knee joint. The skin layer, subcutaneous tissue and sarcolemma were peeled off layer by layer and the femoral condyle was exposed. A defect (5 mm in diameter and 5 mm in depth) was created at the lateral femoral condyle via the cortical and trabecular bone with a dental drill. The


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hole was flushed with sterile saline to remove the bone fragments and the corresponding scaffold was implanted. Finally, the incision was sutured carefully with absorbable sutures. 50 million units penicillin was immediately injected intramuscularly for 3 consecutive days after operation. 2.10. X-ray analysis. To evaluate the osseointegration between implants and host bone, X-ray examination of femoral condyle specimens was done at the same condition and distance using a specimen radiography system (Faxitron X-ray, MX-20, USA) at the energy value of 26 kV for 10 s. 2.11. µ-CT analysis. In order to assess the microstructure of the implanted scaffold in the defect site and the bone integration, µ-CT analysis was executed using a Scanco Medical AG vivaCT40 (Switzerland) with 70 kVP X-ray source and 114 µA. Samples were scanned through a 360º rotation angle at a resolution of 19 µm. 3D µ-CT images were reconstructed using the µ-CT v6.1 of Scanco medical AG software (Switzerland) to evaluate the repair progress. 2.12. Histological and immunohistochemical processing. All samples were fully decalcified in 12% EDTA-2Na, dehydrated by gradient alcohol and embedded in paraffin. Serial sections with 6 µm thick were prepared by a microtome (Leica, Germay). Following dewaxing, part of sections was stained with H-E (Beyotime, China) for extracellular matrix and Masson’s trichrome stain (Nanjing Jiancheng Biotechnology Institute, China) for collagen fibers and muscle fibers as previously described.13 Additionally, immunohistochemistry staining of different markers was performed as previously described.13 Among these, rabbit anti-rat osteocalcin (OC) polyclonal antibody (Bioss, China; diluted 1:50), rabbit anti-rat osteopontin (OPN) polyclonal antibody (Bioss, China; diluted 1:300), the primary antibody anti-SDF-1α (Boster, China; diluted 1:200) or antiCXCR4 (Boster, China; diluted 1:200) was used, respectively. Three samples for each group


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were observed and analyzed by an inverted light microscopy (Olympus IX71, Japan). Semiquantitative analysis (capsule thickness, peri-implant gap, and mean density for OPN and OC) of the histological and immunohistochemical staining was conducted using the Image J (National Institutes of Health, USA) and Image-Pro Plus software (Media Cybernetics, USA), respectively. Using Image-Pro Plus software, the sum of total area and integral optical density (IOD) of the selected pictures were calculated separately, and their ratio is the average IOD per area (IOD/area). 2.13. Statistics. Each experiment was repeated at least three times. All data were shown as means ± standard deviation (SD). One-way analysis of variance (ANOVA) was used to analyze and compare the data between groups with Origin 8.0 software. p < 0.05 and p < 0.01 indicate significant difference and extremely significant difference, respectively.

3. RESULTS 3.1. Characterization of scaffold. 3.1.1. Decalcification rate of scaffold. Decellularized bone scaffolds were decalcified for different time. At 0 h, 1 h, 6 h, 12 h, 1 d, 2 d, and 3 d, the calcium content of scaffolds were 136.72 ± 8.22 mg/g, 106.10 ± 8.53 mg/g, 97.01 ± 8.34 mg/g, 84.83 ± 10.84 mg/g, 70.37 ± 6.68 mg/g, 24.38 ± 2.83 mg/g, and 12.15 ± 7.23 mg/g, respectively (Figure 1A). This indicates that the calcium content of scaffolds decreased and the decalcification rate increased with the decalcification time increased. After decalcification for 5 days, the calcium content of scaffolds were 0.23 ± 0.4 mg/g and regarded as fully decalcification.


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Figure 1. Mechanical and structural characterization of scaffold. Calcium content and decalcification rate of the scaffolds at different decalcification time points (A). Mechanical properties of the scaffolds at different decalcification rate (B). µ-CT images of scaffold (Low) before and after decalcification (C). Pore size and porosity (D) of scaffolds (Low) before and after decalcification. The scale bar indicates 1 mm. Date are shown as the means ± SD, n ≥ 3.

3.1.2. Mechanical properties of scaffolds with different decalcification rate. The compressive elastic modulus of scaffolds at different decalcification rates was measured by compression testing. When the decalcification rate was 2.20 ± 2.70%, 22.39 ± 6.24%, 37.96 ± 7.93%, 48.53 ± 4.88%, 82.16 ± 2.07%, 91.12 ± 5.29%, 99.83 ± 0.29%, the compressive elastic modulus of scaffold was 230.93 ± 72.65 MPa (Control), 66.06 ± 27.83 MPa (High), 26.90 ± 13.16 MPa (Medium), 6.99 ± 3.85 MPa, 0.72 ± 0.25 MPa, 0.83 ± 0.42 MPa, 0.67 ± 0.14 MPa (Low). The compression modulus value has no significant difference among last three groups (decalcification rates 82.16 ± 2.07%, 91.12 ± 5.29% and 99.83 ± 0.29% group) (Figure 1B).


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3.1.3. Structural characteristics of scaffolds. The architecture of scaffolds was evaluated by µCT before and after decalcification. The 3D reconstruction image clearly demonstrated the pore size and interconnectivity of the porous structure (Figure 1C). After demineralization, the mean pore size of the Low scaffold was 380.64 ± 38.29 µm, and mean porosity was 75.33 ± 3.6% (Figure 1D). Before decalcification, the mean pore size of the Control scaffold was 357.90 ± 10.43 µm, and mean porosity was 75.87 ± 2.12% (Figure 1D). As expected, the architecture parameter of the scaffolds has no obvious difference between them. 3.2. Cell compatibility of the scaffolds with different stiffness. Live/dead assay of cultured scaffolds revealed uniform cell distribution after culture for 3 days, with thin cell layers lining the scaffold surfaces. Over the first week of culture, the cell number increased and the majority of the internal spaces were filled with MSCs. After 3 weeks of culture, cell density was relatively constant in all groups and most of the MSCs were viable (Figure 2). Live/dead assay showed good cell compatibility over the culture period. In all groups, areas with the greatest cell density were observed close to the scaffold surface.


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Figure 2. Live/dead assay of the scaffolds cultured for 3 days, 14 days, and 21 days. Green indicates live cells, and red indicates dead cells. The scale bar indicates 50 µm.

3.3. In vitro osteogenic differentiation of MSCs on scaffolds with different stiffness. To assess the effects of 3D DBM scaffolds with different stiffness on the osteogenic differention of MSCs, the production of ALP was tested using ALP assay method at 14, and 21 days postseeding (Figure 3A, B). As shown in Figure 3A, the High, Medium, and Low groups exhibited much more ALP production than the Control group after seeding, and the most ALP production


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was observed in the Low group. The results of in vitro composite culture for 21 days are consistent with 14 days. The results of in vitro culture for 14 days and 21 days together demonstrated that the low stiffness group significantly increased ALP activity and promoted early osteogenic differentiation of MSCs compared with other stiffness groups.

Figure 3. Effects of different stiffness on the osteogenic differentiation of MSCs in scaffolds in vitro. ALP staining on scaffolds with different stiffness after culture for 14 days (A) and 21 days (B). H-E staining and positive immunohistochemical of OPN and OC on scaffolds with different stiffness after culture for 14 days (C) and 21 days (D). Semi-quantitative data for OPN and OC on scaffolds with different stiffness after culture for 14 days (E) and 21 days (F). The scale bar indicates 1 mm in (A) and (B) while 100 µm in (C) and (D). Date are shown as the means ± SD, n ≥ 3; * p < 0.05.


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The expression of specific protein markers (OPN and OC) were further investigated by immunohistochemistry after culture 14 and 21 days in vitro (Figure 3C-F). After culture for 14 days, except the positive expression of OC in the Low group is significantly lower than that in the Middle group, there is no significant difference among other groups (Figure 3E). The OPN expression in the experimental groups is higher than that in the Control group (Figure 3E). After culture 21 days, the Low stiffness group exhibited higher expression of OC than other groups (Figure 3F). The positive expression of OPN in the Low group is significantly higher than that in the Middle group (Figure 3F). The above results indicate that the Low group promote the expression of OC and OPN after culture for a certain period of time. 3.4. Subcutaneous implantation study. After implantation for 28 days, histological sections were examined by H-E staining to detect cell infiltration (Figure 4A) and by Masson’s trichrome staining to detect accumulation of collagen fibers (Figure 4B). H-E staining showed that cell density in the Control group was relatively lower than those in the High, Medium, Low groups. As the decrease of stiffness, almost all of the pores of the scaffold were filled with infiltrated cells and some invaded micro-vessels also could be observed. Masson’s trichrome staining indicated that the accumulation of collagen fibers in the experimental groups was more than that in the Control group. Among them, the blue collagen deposition in the Low group was significantly higher than those in the Middle and High groups. Meanwhile, more capillary invaded into the dense collagen fibers areas in the Low group. The histological results indicate that vessel could invade into the experimental scaffolds and cells could well infiltrate into the scaffolds.


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Figure 4. Histological characterization of scaffolds after subcutaneous implantation. H-E staining (A) and Masson’s trichrome staining (B) of scaffolds with different stiffness after subcutaneous implantation for 28 days. S, implanted scaffold. V, vessel. The scale bar indicates 100 µm.

Immunohistochemical staining of bone protein OC and OPN were used to evaluate the osteogenic differentiation of host cells in the scaffolds with different stiffness (Figure 5). After implantation for 28 days, the positive expressions of OC and OPN in the Low group was the highest, while the expressions of them in the Control group is very low (Figure 5C). This result suggests that proper matrix stiffness may induce differentiation of endogenous cells into the osteogenic direction.


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Figure 5. Characterization of MSCs osteogenic differentiation on scaffolds with different stiffness postimplantation for 28 days. Positive immunohistochemical staining of OPN (A) and OC (B) on scaffolds. Semiquantitative data for OPN and OC on scaffolds (C). The scale bar indicates 100 µm. Date are shown as the means ± SD, n ≥ 3; * p < 0.05.

3.5. In vivo osseointegration of DBM scaffold with different stiffness. 3.5.1. Radiographic assessment. X-ray analysis of osseointegration between implants and host bone at 1, 2 and 3


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months were shown in Figure 6A. At 1 month post implantation, the defect area in the Self repair group was very remarkable. Fusion phenomenon between the edge of the scaffold and host bone was clear in the Low and Medium groups. The scaffold implantation area was clearly visible and the fusion phenomenon could not be seen in the High group. Because the stiffness of the Control group is close to that of the cancellous bone of the host bone, only non-transparent white area with greater bone density could be observed. Therefore, the specific repair and fusion effects in the Control group cannot be analyzed.

Figure 6. Imaging assessment of scaffolds in rabbit femoral condyle defect model. X-ray images (A) and 3D reconstructions (cross-sectional view and sagittal view) of µ-CT images (B) of scaffolds with different stiffness at 1, 2 and 3 months post-implantation. The red square indicates the defect repair area. The scale bar indicates 5 mm.

At 2 months post implantation, the defect area in the Self repair group was still clearly visible. There was a well fusion in the Low group and the bone density was higher than that in the


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Medium group. In the High group, a great quantity of hyperplasia irregular bone was visible in the surface of defect site, and the fusion phenomenon could not be seen in the implant-bone interface. Hyperplasia bone was also formed in the Control group. At 3 months post implantation, there was still a large vacancy and no distinct bone connection in the defect area in the Self repair group. The bone defect was almost completely repaired in the Low group, and the internal structure was closest to that in the Normal group. Implantation area was still clearly visible in the High group, indicating that there was no significant bone fusion. 3.5.2. µ-CT analysis. The 3D restructured from µ-CT scan were performed to evaluate the bone integration during healing process of the femoral condylar defected bone at 1, 2 and 3 months (Figure 6B). The evident bone binding or massive bone formation was not observed in the Self repair group until 3 months. There were poorer osseointegration effects between the host and scaffold in the Control and the High groups, compared to the Low and Medium groups at 3 months. Part of edges in the Low scaffold began to fuse with the host bone at 1 month. In contrast, a small gap could be observed in the Medium group. In addition, from the internal structures observed in the six images, it can be seen that only the Low group obtained almost the same internal structure as the Normal group. 3.6. Effect of scaffolds with different stiffness on the osteogenesis in bone defect animal experiment. Histological analysis (H-E staining (Figure 7A) and Masson trichrome staining (Figure 7B)) of retrieved specimens was performed after implantation for 1, 2 and 3 months. The Self repair group showed numerous red fibrous tissues lasted till 3 months. At 1 month, there were clear gap between the host and scaffolds in the Medium, High and Control groups. The gaps in the High and Control groups were larger than that in the Medium group, indicating that these two groups had inefficient osteointegration repair. However, there was no significant gap


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between the host and scaffolds in the Low group, just filled with few fibers. The Low group showed better osteointegration repair at the early stage of bone repair. With the extension of repair time, each group get a certain extent repaired, showing a decreased gap between the host bone and the graft. At 2 and 3 months, there were significant increase in deposition of ECM and mature bone distribution in the defect repair area in the Low group as compared to other groups. After 1 month of implantation, the fiber capsule thickness were 32.43 ± 27.64 µm (Low), 120.67 ± 6.88 µm (Medium), 202.58 ± 66.30 µm (High), 162.19 ± 11.55 µm (Control) (Figure 7C). The results showed that the fiber capsule thickness formed at the defect site of the Low group was the smallest, indicating that the in vivo inflammatory response was the lowest. Masson's trichrome staining showed a better osseointegration in the Low group compared to other experimental groups, in which only with a small gap filled with denser blue collagen fibers between the host bone and the graft. At 3 months, more mature bone and well fusion between host bone and graft could be observed in the Low group. The Self repair group did not show significant collagen deposition after 1 month. Semi-quantitative data for peri-implant gap from Masson’s trichrome staining images were shown in Figure 7D. After 1 month of implantation, the peri-implant gap in the Medium, High and Control groups were 3.77 ± 1.38, 5.05 ± 1.74 and 6.78 ± 1.20 times than that in the Low group, respectively. At 1, 2, and 3 months, the size of the peri-implant gap showed: the Low group < the Medium group < the High group < the Control group, and there was a significant difference between each two groups. In addition, the gap in each group decreased with the prolonged repair time. These findings suggested that the Low group with smallest peri-implant gap was more conducive to the bone integration and healing compared with the Control group.


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Figure 7. Histological characterization of scaffolds with different stiffness after implantation in rabbit femoral condylar bone defects for 1, 2 and 3 months. H-E staining (A) and Masson’s trichrome staining (B) of scaffolds. Semi-quantitative data for capsule thickness from H-E staining images at 1 month (C) and periimplant gap from Masson’s trichrome staining images at 1, 2 and 3 months (D). G, grafts. H, host bone. F, fibrous tissue. CF, collagen fibers. NB, new bone. MB, mature bone. CC, chondrocytes. The yellow arrows indicate the border between the implant and the host bone. The scale bar indicates 100 µm. Date are shown as the means ± SD, n ≥ 3; * p < 0.05, compared with the control group. # p < 0.05, compared with the low stiffness group. & p < 0.05, compared between the medium stiffness group and high stiffness group.

3.7. Effect of SDF-1α/CXCR4 pathway in bone repair mediated by matrix stiffness. The expression of SDF-1α and CXCR4 were detected in the defect repair area, indicating that SDF-


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1α and CXCR4 were involved in the bone repair at the early stage (Figure 8). Compared to the host bone, higher levels of SDF-1α and CXCR4 expression were detected in the DBM scaffolds. The expression of SDF-1α in the High and Control groups at the border of the host bone and implants was higher than that in the Low group. The Self repair group also showed a higher level of SDF-1α and CXCR4 compared to the Normal group. It can be concluded that the expression of SDF-1α and CXCR4 were correlated with the recruitment of endogenous osteoprogenitor cells and the inflammatory response mediated by the matrix stiffness of the implanted scaffold.

Figure 8. Immunohistochemistry analysis of scaffolds with different stiffness after implantation in rabbit femoral condyle defects for 1 month. Positive immunohistochemical staining of SDF-1α (A) and CXCR4 (B) on scaffolds. The red arrow indicates the border between the implant and the host bone. G, graft. H, host bone. The scale bar indicates 100 µm.


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4. DISCUSSION In this study, a novel method to create 3D scaffolds with controllable stiffness but the same 3D microstructure was developed by controlling the decalcification time. Subcutaneous transplantation and rabbit femoral defect were performed to assess various capabilities of 3D DBM scaffold with different stiffness. The pore size of the 3D scaffold is an important structure feature, which can directly determine cell survival and growth.24 In order to maintain the homogeneity of scaffold, the decellularized bone scaffolds were weighed and divided into three groups according to their quality: high-density group (0.688 ± 0.076 mg/mm3), medium-density group (0.535 ± 0.076 mg/mm3) and low-density group (0.382 ± 0.076 mg/mm3), which is consistent with the mass density reported by Marcos-Campos et al.25 µ-CT results showed that the average pore size and porosity of the high-density scaffolds were 393.6 ± 30.6 µm and 68.9% ± 3.1% (see Supporting Information Figure S1), which are very close to the reported optimal pore size (325 µm)26 and the best porosity (75%).27 Therefore, high-density scaffolds were selected and used for decalcification to prepare DBM scaffold. After fully decalcification, the average pore size was 380.64 ± 38.29 µm which not only is close to the optical pore size in bone tissue engineering but also have no significant difference between before and after decalcification. In the present study, 3D DBM scaffolds with different compressive elastic modulus ranging from 0.67 MPa to 230 MPa were prepared. Previous studies28 have demonstrated that the osteogenic differentiation of the MSCs occurred predominantly at 11-30 kPa, especially at 22 kPa. However, MSCs will actually also respond to stiffness beyond this range. For example, Guo et al.29 found that osteogenic differentiation and mineralization of MSCs increased with increasing substrate modulus in the MPa range. More interestingly, Maggi and his colleagues30


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constructed 3D nano-structured scaffolds with different stiffness similar to our decalcified scaffolds, specifically 0.69 ± 0.2 MPa, 16.8 ± 6.9 MPa and 60.2 ± 7.4 MPa, respectively. Their study found that the nanolattice with stiffness of 0.7 MPa had ~20% more intracellular F-actin than the stiffer nanolattice. They concluded that the cellular mineralization-inducing ability of 3D substrates is very sensitive to their structural rigidity and the best osteoblast function can be obtained on substrates at 0.7 to 3 MPa (similar to cartilage). These results coincided with our in vitro experiments. The ability of DBM scaffolds to support osteogenic differentiation of bone marrow MSCs in vitro was demonstrated by evaluation of several osteogenic markers. ALP assay showed that the Low group exhibited higher ALP production than other groups. The ALP production at 21 days is less than that at 14 days. Previous in vitro studies31 have demonstrated that the synthesis of ALP often decreases after 21 days because the synthesis of ALP occurs in early osteogenic differentiation, and less abundant cell population begins to proliferate and synthesize ALP during the third weeks. The OC expression on the Low scaffolds was much higher than those on other stiffness scaffolds, which suggests that the Low scaffolds could promote the osteogenic differentiation. Surprisingly, the OC expression in all groups has no significant difference at 14 days. This study speculates that the OC expression may not increase at 14 days, as OC is a late stage marker of osteogenic differentiation. This is consistent with the result of ALP production. In addition to inducing osteogenic differentiation of MSCs, angiogenesis is also one of the important indicators for evaluating the repair effect of scaffold in vivo. In subcutaneous experiments, a large number of micro-vessels invaded in the dense collagen fibers area in the Low group, indicating that scaffold with a stiffness of 0.67 ± 0.14 MPa not only has the potential to promote new bone formation but also significantly promote angiogenesis. In addition, efficient


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in vivo vascularization of different scaffolds can be found after subcutaneous implantation for 4 weeks (Figure S2). Although this study proposed a novel method to create 3D scaffolds with controllable stiffness but the same 3D microstructure (porosity and pore size), some limitations on mass production (e.g., relatively limited source, and long production cycle) of the proposed matrix can not be ignored. And these limitations need to be further improved in the future study. In order to ensure the consistency of the stiffness of DBM used in rabbit femoral condylar bone defect and in vitro cell experiments, the compressive modulus of the scaffolds with different shapes (Φ5 mm × 5 mm and Φ10 mm × 5 mm) were compared. The compressive elastic modulus of scaffold (Φ5 mm × 5 mm) was 0.68 ± 0.01 MPa (Low), 21.32 ± 1.74 MPa (Medium), 66.80 ± 21.60 MPa (High) and 233.92 ± 34.02 MPa (Control) (Figure S3). The results showed that the compression modulus has no significant difference between the scaffolds with different shapes after decalcification with the same time. This result indicates that the method of constructing 3D bone graft with different matrix stiffness by controlling the degree of decalcification is reproducible and feasible in our experiments. In the rabbit femoral condyle defect model, the methods of imaging and histology were used to assess the integration of DBM scaffolds with different stiffness. Moreover, analysis of the irregular hyperplastic bone in defect surface, fibrogenesis, covered smooth and transparent soft tissue, and the exposed area of implanted scaffold (Figure S4) of retrieved samples were also performed. All the findings indicated that the Low group had better osteointegration and bone healing in vivo. Analysis of angiogenesis of scaffolds with different stiffness in rabbit femoral condyle defects at 1 and 3 month post-implantation was also performed (Figure S5). At the early stage of bone defect repair, more vascular structures were formed in the middle region between


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the implant and host bone in the Low group, which was consistent with the results of subcutaneous experiments and indicated that the 3D DBM scaffold with a stiffness of 0.68 ± 0.01 MPa was favorable for angiogenesis at the defect site. The genuine morphology and function reconstruction of bone defects achieved by promoting recruitment of endogenous stem/progenitor cells in situ provides new insights into in situ tissue regeneration. It has been reported that the interaction between inflammatory environment,32 hypoxia,33 and chemokine14,34 affect the recruitment of endogenous MSCs and bone healing. Among them, the cross-talk between the chemokine SDF-1α (also referred to as CXCL12) and its receptor CXCR4 plays a major role in regulating endogenous stem cell homing and has been reported to be involved in repair and regeneration of various tissues and organs, especially in cartilage35 and bone.34 Partial results of immunohistochemistry in this study (Figure. 8) was consistent with our previous study,36 both revealed higher expression of SDF-1α in the graft region compared with the host bone, which confirmed that SDF-1α may be one of the important participants in bone repair process of bone marrow MSCs. More interesting, the group with best repair had lower SDF-1α expression level compared with the High and Control groups. The probable cause is that SDF-1α expression level was also closely related to the strength of inflammation in vivo.37 The High and Control groups both with poor defect repair result in stronger inflammatory response after implantation. This indirectly proved by the thickness of the fibrous capsule counted by H-E staining (Figure 7C). Studies have reported that the chemotaxis of SDF-1α to lymphocytes, monocytes, dendritic cells and other inflammatory cells resulted in significant high expression of SDF-1α under inflammatory conditions.38 Simultaneously, the CXCR4+ inflammatory cells accumulate, adhere, cross the endothelial cells, reach the site of inflammation, and are widely expressed on immune cells. Considering that the repair effect of


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the implants was not only affected by matrix stiffness mediated orthotopic osteogenic differentiation, but also related to the early inflammatory response in vivo induced by matrix stiffness, the correlation between scaffold stiffness and inflammatory response is of great worthy for our in-depth study in following research work.

5. CONCLUSIONS 3D DBM scaffold with tunable matrix stiffness but the same 3D microstructure (porosity and pore size) has been successfully constructed by adjusting the degree of decalcification. The influence of growth factors exposed by decalcification was excluded by inactivation of guanidine hydrochloride. This study demonstrated that the DBM scaffold with a stiffness of 0.67 ± 0.14 MPa could promote the osteogenic differentiation of MSCs, significantly improve new bone formation and enhance bone repair and integration compared to scaffolds with 26.90 ± 13.16 MPa, 66.06 ± 27.83 MPa and 230.93 ± 72.65 MPa. The roles of SDF-1α/CXCR4 pathway in endogenous stem/progenitor recruitment and osteogenic differentiation during bone repair were further confirmed. The novel method for controlling the matrix stiffness as a single variable proposed in present study provides additional references for the construction of new alternative implant and can expand the researches of the effects of 3D stiffness on the MSC behavior and repair of bone defects.

ASSOCIATED CONTENT Supporting Information. Additional experimental results are included in the support information, including the mean pore size and porosity of decellularied cancellous bone scaffolds with different mass density groups,


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gross observation of scaffolds after subcutaneous implantation, mechanical test of each group of scaffold, gross appearances in DBM scaffolds with different stiffness at after implantation in rabbit femoral condyle defects for 1, 2 and 3 months, analysis of angiogenesis of scaffolds with different stiffness in rabbit femoral condyle defects at 1 and 3 month post-implantation (PDF). These materials are available free of charge via the Internet at http://pubs.acs.org. AUTHOR INFORMATION Corresponding Author *Tel: 86-23-65102507. Fax: 86-23-65102507. E-mail: [email protected]. Author Contributions The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. Q.H. and M.L. contributed equally to this work. Notes The authors have no competing financial interests. ACKNOWLEDGMENT This work was supported in part by grants from the National Natural Science Foundation of China (11702043, 11672051) and the Fundamental Research Funds for the Central Universities (2018CDQYSG0015).

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Table of Contents/Abstract Graphic


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Demineralized Bone Scaffolds with Tunable Matrix Stiffness for

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