Nanoscale SilkâHydroxyapatite Hydrogels for Injectable Bone

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Nanoscaled Silk-Hydroxyapatite Hydrogels for Injectable Bone Biomaterials Zhaozhao Ding, Hongyan Han, Zhihai Fan, Haijun Lu, Yonghuan Sang, Yuling Yao, Qingqing Cheng, Qiang Lu, and David L Kaplan ACS Appl. Mater. Interfaces, Just Accepted Manuscript • Publication Date (Web): 04 May 2017 Downloaded from http://pubs.acs.org on May 5, 2017

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ACS Applied Materials & Interfaces

Nanoscaled Silk-Hydroxyapatite Hydrogels for Injectable Bone Biomaterials Zhaozhao Dinga,b, #, Hongyan Hana,#, Zhihai Fanc, #, Haijun Luc, Yonghuan Sangb, Yuling Yaoa, Qingqing Chengb, Qiang Lua,b,*, David L Kapland a

School of Biology and Basic Medical Sciences & Collaborative Innovation Center of Suzhou Nano Science and Technology, Soochow University, Suzhou 215123, People’s Republic of China b

National Engineering Laboratory for Modern Silk, Soochow University, Suzhou 215123, People’s Republic of China

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Department of Orthopedics, The Second Affiliated Hospital of Soochow University, Suzhou 215000, People’s Republic of China

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Department of Biomedical Engineering, Tufts University, Medford, Massachusetts 02155, United States

#

The authors have contributed equally to the first author

*Address corresponding to [email protected]

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ABSTRACT Injectable hydrogel systems are important bone substitutes for regeneration due to handling properties and the ability to fill irregular defects. Silk-hydroxyapatite composite materials with silk nanofibers in hydrogels were prepared and used as biomaterials for osteogenesis. These thixotropic silk nanofiber hydrogels and waterdispersible silk-HA nanoparticles were blended to form injectable nanoscale systems with a homogeneous distribution of a high HA content (60 % w/w) to imitate bone niche. A modulus of about 21 kPa was also achieved following the addition of HA in the systems, providing physical cues to induce osteodifferentiation. The composite hydrogels supported improved osteogenesis compared to silk nanofiber hydrogels. The newly formed bone tissue and bone defect healing were detected after implantation of the silk-HA composite hydrogels, suggesting utility for the regeneration of irregular bone defects. KEYWORDS: injectability, hydroxyapatite, silk, bone regeneration, biomimetic.

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1. INTRODUCTION Large calvarial bone defects from trauma, abnormalities, nonunion fractures, infections or tumor resection remain a challenge in terms of repair.1-4 The implantation of autologous or allogeneic bone grafts has been widely used to repair such bone defects and achieved impressive functional restoration. However, the success rate for these approaches is suppressed by several inherent disadvantages, including donor site morbidity and infection.5-7 Therefore, alternatives such as graft substitutes as scaffolds, gels, cements and powders derived from synthetic and natural biomaterials have been considered to replace or repair damaged bone sites.8-11

Bone substitutes including titanium alloys, magnesium alloys and stainless steel have been widely used to replace damaged bones due to their excellent mechanical properties.12-15 However, the mechanical mismatch between these substitutes and the surrounding bones usually result in the impairment of the repair toward normal bone. The lack of tissue adherence and the accumulation of corroded metal ions also remain as challenges for better functional recovery.16-18 Recently, bioactive artificial bones were developed to induce the regeneration of new bone at damaged sites.19,

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Although these artificial bones have inferior mechanical properties than normal bones, they provided cues and preferred microenvironments to accelerate bone repair and

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functional recovery. Significantly, better bone regeneration was achieved with these artificial bones by adjusting composition toward ECM-like nanofibrous structures with tunable delivery of growth factors.21-24 Besides improvements in structure and bioactivity, injectability features for clinical applicability to reduce damage to surrounding tissues, to shorten surgical operation time, and to accelerate the recovery of the damaged sites are also attractive options.25-27 However, improvement in injectability often compromises bioactivity,24 thus, remaining a challenge.

Silk fibroin, a natural protein extracted from Bombyx mori silkworms, has been extensively applied in biomedical fields due to its biocompatibility, adjustable biodegradability, and ease of processability.28-30 Silk-based bone scaffolds have been optimized by tuning composition, mechanical properties and delivery of growth factors. Although silk-based materials have been considered promising as matrices for bone regeneration, there has been less focus on injectable bone materials. Silk-based materials to fill irregular bone defects through minimally invasive surgery could address a significant gap in the current needs. Recently, injectable silk fibroin nanofiber hydrogels were prepared in our group and used as carriers to release anticancer drugs, suggesting a possibility of designing silk-based injectable bioactive artificial bone biomaterials.31 Meanwhile, hydroxyapatite (HA), main mineral phase

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of bone, has outstanding osteoconductivity and has been blended with silk fibroin to form composite scaffolds with porous structure and improved osteoinduction.32 Therefore, injectable artificial bone biomaterials were designed with silk fibroin nanofibers and HA nanoparticles. Water dispersible HA nanoparticles were blended with silk fibroin nanofibers to form silk fibroin-HA composite hydrogels. The ratio of HA nanoparticles and silk fibroin nanofibers was tuned to achieve mechanical properties and organic-inorganic compositions similar to natural bone, while retaining the ability to be injected.

Significant improvement in osteogenic capability was

demonstrated in vitro and in vivo using these hydrogels.

2. MATERIALS AND METHODS 2.1. Preparation of aqueous silk solutions. Silk fibroin solutions were prepared via a traditional dissolution process as described previously.33 Briefly, raw silk was boiled for 20 min in an aqueous solution of 0.02 M Na2CO3, and then rinsed thoroughly with distilled water to extract the sericin proteins. The extracted silk fibers were dissolved in 9.3 M LiBr solution at 60 oC for 4 h, and dialyzed against distilled water with dialysis tubes (MWCO 3,500) for 72 h to remove the salt. After centrifuging at 9,000 rpm for 20 min at 4 oC to remove silk aggregates formed during the process, the optically clear silk fibroin solution

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with a concentration of about 6 % (w/v) was prepared, which was determined by weighing the remaining solid after drying. 2.2. Silk fibroin-coated HA nanoparticle formation. Silk fibroin-coated HA nanoparticles were prepared via an aqueous precipitation reaction using silk fibroin as template and surface stabilizer.32 Aqueous solutions of Ca(OH)2 (0.02 M) and H3PO4 (0.06 M) were prepared by dissolving the above chemicals in distilled water separately. The fresh silk fibroin solution (6 wt%) was incubated at 60 oC for 24 h to generate homogeneous nanoparticles. The 20 mL H3PO4 solutions were first mixed with 20 mL of silk fibroin solution to form silk fibroin-H3PO4 mixed solution. Then the mixed solution was added drop-wise to 100 mL Ca(OH)2 suspensions at 90 mL h-1 along with vigorous stirring. The solution was heated in a water bath at 70 oC, following the adjustment of the solution pH to about 9 with 0.1 M NaOH. Finally, the emulsions were centrifuged at 9,000 rpm for 20 min and washed gently with distilled water to recover the silk fibroin-coated HA nanoparticles. 2.3. Silk fibroin nanofiber formation. The silk fibroin nanofibers were assembled through the recently reported concentration-dilution process.34,

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The solution (6 wt%) was slowly concentrated to about 20 wt% over 24 h at 60 o

C to form metastable nanoparticles, and then diluted to 2 wt% with distilled

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water. The diluted silk fibroin solution was incubated over 24 h at 60 oC to induce nanofiber formation. The silk fibroin nanofiber was termed SF. 2.4. Preparation of SF-HA composite hydrogels. The SF-HA composite hydrogels were fabricated by mixed SF and HA homogeneously under stirring sequentially. The hydrogels containing 20 %, 40 % and 60 % HA particles were termed SF-HA20, SF-HA40 and SF-HA60, respectively. 2.5. SF-HA hydrogel characterization. The microstructures of SF, HA and the composite hydrogels were characterized by scanning electron microscopy (SEM, S4800, Hitachi, Tokyo, Japan) at 3 kV. Samples were mounted on a copper plate and sputter-coated with gold prior to imaging.28 Fourier transform infrared spectroscopy (FTIR) analysis of samples was performed with a Nicolet FTIR 5700 spectrometer (Thermo Scientific, FL, USA). The infrared spectrum was recorded from 4000 to 500 cm-1. X-ray diffraction (XRD, Nano ZS90, Malvern instruments, Malvern, UK) analysis was carried out between 2θ values of 10 o to 55 o using monochromated Cu Kα radiation (30 mA, 40 kV) with a scanning speed of 6 °/min to confirm characteristic diffraction peaks of SF and HA in the composite hydrogels. 36 2.6. Rheological studies of the hydrogels. Rheological behavior of the hydrogels was measured on a Rheometer (AR2000, TA Instruments, New Castle, USA) fitted with a 20 mm cone plate (Ti, 20/1°). Frequency sweeps were collected continuously

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over a wide frequency range from 100 to 1 rad s-1 at 25 °C. The stability of the SF and SF-HA composite hydrogels was assessed by measuring the rheological behavior under a constant frequency from 22 oC to 40 oC. 37 2.7. Viscosity and injectability of the hydrogels. The viscosities of SF and SF-HA composite hydrogels were plotted against varying shear rates of 10-1 to 10-2 s-1 to obtain flow curves at room temperature. Different hydrogels were suctioned into 1 ml syringes and pushed out without a needle or through a needle (27G) under 1.5 N to assess injectability at 37 oC. Then inversion study was used to evaluate the stability of the hydrogels. A 5 ml aliquot of the hydrogels placed in cylindrical bottles with flat base, and the bottles were inverted to stand on their caps. The flow of the hydrogels was observed at 10 min, 2 h, 60 h and 100 h, respectively. 37 2.8. In vitro biocompatibility of the hydrogels. Male Sprague-Dawley rats (SD rats) (~40g) were euthanized and used to isolate bone mesenchymal stem cells (rBMSCs) under sterile condition. The use of rats was carried out following approval and granted by the animal ethics committee of Soochow University. The rBMSCs were seeded in cell culture dishes in DMEM containing 10 % FBS and 100 U/ml-1 penicillin-streptomycinand (Gibco, Grand Island, CA, USA), and continuously passaged to achieve the third generation. To study cell proliferation, rBMSCs were seeded in 6 well plates at a density of 5×105 cells per well containing complete

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DMEM. After the cells became confluent, the DMEM was removed. Then SF and SF-HA composite hydrogels sterilized with γ-irradiation at dose of 25 kGy were tiled on the wells, followed by adding complete DMEM and incubating at 37 oC. At the indicated time points (1, 3, 5 and 7 days), all samples were digested with proteinase K overnight at 56 oC.32 Then DNA content was determined using the PicoGreenTM DNA assay following the protocol of the manufacturer (Invitrogen, Carlsbad, CA, USA). 2.9. Cell differentiation. To study cell differentiation, passage three BMSCs were seeded on glass slides in the 24 well plates with DMEM containing 10 % FBS and 100 U/ml-1 penicillin-streptomycin (Gibco, Grand Island, CA, USA) and cultured for 24 h. Then the medium was removed and sterile SF and SF-HA composite hydrogels were tiled on the glass slides. The medium was exchanged with osteogenic media (low glucose-DMEM, 10 % FBS, 1 % streptomycin-penicillin, 10 nM dexamethasone, 10 mM sodium-β-glycerophosphate, and 0.05 mM ascorbic acid-2-phosphate) for 7, 14 and 21 days, respectively.

Alkaline phosphatase (ALP) is a glycoprotein located on the cell surface, and is the widely recognized marker of osteoblast differentiation.38 Other different bone-related gene expression markers, Runx2 (an early marker of osteogenic maturation),

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osteocalcin and osteopontin (later markers of osteogenic differentiation and mineralization),42,

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were measured to assess the differentiation states of BMSCs

using quantitative real-time PCR. All the measurements were processed according to the protocols that have been reported in our recent studies.36, 44

The expression of bone specific proteins such as ALP, osteocalcin (OCN), and osteopontin (OPN) were analyzed by immunofluoresence. The rBMSCs were cultured on glass slides in 24 well plates with osteogenic media followed by cell differentiation. The expression of ALP was observed by labeling anti-ALP antibody at 7 days. The expression of OCN and OPN was detected by labeling anti-OCN (Abcam, Cambridge, MA, USA) and anti-OPN antibody (Boster, Pleasanton, CA, USA) at 21 days. Briefly, the samples were fixed in 4 % paraformaldehyde solution (Sigma-Aldrich, St. Louis, MO, USA) for 30 min, permeabilized with 0.1 % Triton x-100 for 10 min, washed three times with PBS, and blocked in 3 % bovine serum albumin (BSA) for 1 h. Cells were incubated with the primary antibodies as per manufacturers recommended after dilution overnight. Secondary antibodies were goat anti-mouse IgG (Abcam, Cambridge, MA, USA) and goat anti-rabbit IgG (Abcam, Cambridge, MA, USA) conjugated with fluorescein isothiocyanate (FITC, Thermofisher, Waltham, MA, USA), respectively. All the samples were counterstained for F-actin using

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tetramethylrhodamine (TRITC, Thermofisher, Waltham, MA, USA) conjugated phalloidin, and finally stained with DAPI for the nucleus and analyzed using laser confocal microscopy. 37 2.10. In vivo animal studies. The in vivo bone-forming ability of the SF-HA hydrogels was examined in a rat calvarium defect model, as previously described.9-11, 45

Thirty-six 8-week-old male Sprague-Dawley rats (SD rats) (~300g) were used and

randomly divided into four groups: SF, SF-HA20, SF-HA40 and SF-HA60. The SD rats were anesthetized by intraperitoneal injection of 4% chloral hydrate. A linear sagittal midline skin incision was made over the calvarium, and a full-thickness flap of the skin and periosteum was elevated. Critical-sized full-thickness bone defects (5 mm diameter) were created at the center of each parietal bone in the calvarium using a dental drill with rinsing by 0.9 % saline. The bone defects were filled with the hydrogels. After surgery, the rats were individually caged and fed. At weeks 8, 12 and 16 the rats were sacrificed to obtain the bony specimens. All procedures were carried out following approvals and granted by the animal ethics committee of Soochow University.

At each time point, the calvaria specimens were harvested and fixed in 10 % buffered neutralized formalin for 24 h, then the calvaria specimens were scanned by

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in vivo X-ray microcomputed tomography (µCT). Following fixation, 3D images were taken using a µCT scanner (SkyScan 1176, SkyScan, Aartselaar, Belgium) at a resolution of 18 µm, a voltage of 65 kV and a current of 381 µA to analyze the new bone formation. The collected 3D scans were reconstructed and analyzed using CTan (SkyScan).

After µCT imaging and analysis, the harvested samples were prepared for histological analysis. The fixed specimens were decalcified with 10 % EDTA at pH 7.4 for one month, dehydrated through a series of ethanol solutions, and embedded in paraffin. Sections were used for haematoxylin and eosin (HE) staining and immunohistochemistry targeting OCN (Abcam, ab13420, 1:200), osteopontin (OPN, Abcam, ab8448, 1:500) as previously described.29,

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2.11. Statistical analysis. The statistical analysis for multiple comparisons was performed using SPSS v.16.0 software. Comparison of the mean values of the data sets was performed using one-way AVOVA. Measures are presented as means ± standard deviations, unless otherwise specified. P≤ 0.05 was considered significant.

3. RESULTS AND DISCUSSION 3.1. hydrogel structure. The injectability of silk nanofiber hydrogels suggested the

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possibility of designing injectable biomimetic silk-HA artificial bone matrices.31 HA nanoparticles were added to the silk nanofiber hydrogels to prepare injectable matrices. Although higher HA content (above 50 wt%) was preferred based on the inorganic/organic ratio of natural bone, a challenge remains to avoid the aggregation of HA and also preserve injectability of the composite hydrogels at higher HA content. Recently, water-dispersible HA nanoparticles were prepared with silk as template and stabilizer and used to design silk-HA scaffolds.29,

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The HA nanoparticles were

dispersed homogeneously inside the silk substrates at a HA content of 40 %, implying that the nanoparticles might be suitable candidates for designing injectable biomimetic silk-HA hydrogels. Therefore, various amounts of the HA nanoparticles were blended with silk nanofiber hydrogels to assess the feasibility of forming composite hydrogels. The HA particles easily dispersed in the silk nanofiber hydrogels via a simple physical process until the HA content was 60 %, while significant aggregation appeared in the hydrogels with an HA content of 80 %. The results indicated that homogeneous silk-HA composite hydrogels containing HA contents analogous to natural bone could be prepared using the SF and HA nanoparticles, which is superior to other silk-HA composite materials reported previously.32, 46 SEM images showed good microporous structures without structural deterioration after the addition of HA nanoparticles (Figure 1A). High magnification

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images of the porous walls further indicated ECM-like nanofibrous structures and homogeneous distribution of HA particles at nanometer scale, even when the HA content was 60 wt%. Both FTIR and XRD analyses revealed typical silk beta-sheet conformation and HA crystal structure in the composite hydrogels, which implied the stability in water (Figure 1B and 1C). Unlike previous silk-HA composite materials that only contained limited HA content, and lacked ECM-mimetic structures,

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bone-like ECM was achieved in the present system, providing an improved microenvironment for bone repair.

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Figure 1. SEM images (A), FTIR spectra (B) and XRD curves (C) of the SF and SFHA composite hydrogels. (A): (a-d) shows the microporous structures of the SF, SFHA20, SF-HA40 and SF-HA60 while (a1-d1) displays the distribution of silk nanofiber and HA nanoparticles in the hydrogels.

3.2. Hydrogel properties. The changes in injectability of the hydrogels containing HA nanoparticles were assessed. The SF nanofibers have high beta-sheet content and also high negative charge density, endowing them hydrophobic property and higher electrostatic repulsion simultaneously. The hydrophobicity of SF nanofibers results in gel formation while the electrostatic repulsion is the main reason of shear-thinning behavior.

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Similar to SF hydrogels,31,

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the SF-HA composite hydrogels

showed shear-thinning behavior, suggesting injectability. All the SF-HA composite hydrogels could be easily pushed out from syringes with and without needles, confirming injectable capacity (Figure 2A). Once the pushing force was removed, the hydrogels stabilized quickly and remained in the solid state under gravitational conditions more than 100 h (Figure 2B). The composite hydrogels could be injected into bone defect sites, making them superior to previous silk-HA materials. Besides injectability, the modulus of the composite hydrogels increased significantly to ten times higher than SF hydrogels when the HA content was 60% (Figure 2C). The

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composite hydrogels showed constant modulus in the temperature range of 22-40 oC, which implied stability in vivo at body temperature. Different studies revealed the impact of matrix elasticity on the differentiation of stem cells, suggesting that an elastic modulus of above 18 kPa facilitated differentiation of stem cells into osteoblast cells.37, 38 Although the modulus (21.2 kPa) was inferior to some artificial bone materials (above 80 MPa),

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the composite hydrogels (SF-HA60) had

achieved higher stiffness than most injectable bone hydrogels derived from natural biomaterials and also had potential to induce the differentiation of stem cells into osteoblast cells, providing physical cues to promotion bone formation.37, 38

Figure 2. Characterization of SF-HA hydrogels: (A) Injectability of the SF-HA

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hydrogels through a syringe without needle (a, c, e, g) and with needle (b, d, f, h). (a, b), (c, d), (e, f), and (g, h) samples were SF, SF-HA20, SF-HA40 and SF-HA60, respectively; (B) inversion test of the hydrogels at different time points. (C) rheological behavior of hydrogels: (a) frequency sweep, (b) elastic modulus, (c) complex modulus (G*) vs temperature, and (d) flow curves of the hydrogels, ***p≤0.001.

3.3. In vitro biocompatibility and osteodifferentiation of stem cells. rBMSC proliferation was used to evaluate cytocompatibility on the hydrogels. Similar cell proliferation appeared on the hydrogels in which cell numbers increased up to 7 days without reaching a plateau. Higher cell numbers were achieved on the SF-HA60 hydrogels, suggesting improved cytocompatibility due to the physical properties (Figure 3A). The cell culture results indicated that the addition of HA particles had no negative effect on cytocompatibility. Possibly due to silk degradation, cell migration behavior inward the hydrogels appeared when cells were cultured on the hydrogels for above 28 days (data not shown), which will be further investigated in our following study.

Various studies clarify that biomaterials with higher stiffness and higher HA content

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usually possess better osteogenic differentiation capacity.46,

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These composite

hydrogels were subsequently used in cellular differentiation studies with SF hydrogels as a control. The level of ALP activity and Runx2 protein expression, two early markers of osteogenesis, were investigated to evaluate the osteogenic differentiation of rBMSC on the hydrogels (Figure 3B, 3C). The cells on the composite hydrogels showed significantly higher ALP activity than those on the pure SF hydrogels. ALP activity was further enhanced following the increase of HA content, achieving the highest value in the rBMSCs cultured on the SF-HA60 hydrogels. Similarly, Runx2 expression achieved highest expression in the BMSCs on the SF-HA60 hydrogels. These results indicated that the increased stiffness and HA in the SF-HA60 hydrogels provided stronger physical cues to improve the osteogenic differentiation of the BMSCs. Similar to previous studies, both markers showed a trend that peaked on day 7 and decreased on days 14 and 21, suggesting subsequent stages of bone regeneration after 7 days.32, 37 The expression of the other two corresponding markers in rBMSCs, osteocalcin and osteopontin, were studied and exhibited significant and sustained increases from 7 to 21 days for all the hydrogels (Figure 3D and 3E). The highest osteocalcin and osteopontin expression was also observed on the SF-HA60 hydrogels, confirming the best osteogenic differentiation capacity. The osteogenic differentiation of rBMSCs was further clarified by immunostaining for various bone specific proteins

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(ALP, OCN and OPN) at the different time points (Figure 4A-4C). Based on the differentiation results, the expression of ALP was analyzed at day 7 while OCN and OPN were detected at day 21. The expression of the three markers was significantly improved in the cells cultured on the composite hydrogels and showed optimized enhancement in the cells cultured on SF-HA60 hydrogels, which supported the quantitative data. Both cell growth and differentiation results indicated that the optimized osteogenic differentiation capacity could be achieved through tuning different physical cues such as stiffness and compositions of the hydrogels. Various signals such as the HA content, nanostructure and mechanical response have been tuned to accelerate bone regeneration. However, few studies achieved the design where several signals were optimized simultaneously in same biomaterial system for bone repair. Here, silk hydrogels with ECM mimetic nanofibrous structures were blended with water-dispersible HA nanoparticles to possess ECMbiomimetic structure, similar to the organic-inorganic composition and with suitable mechanical cues. The multiple simulations of the milieu of bone resulted in enhanced osteogenic differentiation capacity of these hydrogels. The injectability provides further benefits to the system compared to other silk-HA bone matrices, especially for repairing irregular bone defects.

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Figure 3. The proliferation and differentiation of rBMSCs on different hydrogels: (A) Cell proliferation from day 1 to 7; and (B-E) ALP activity, mRNA levels of RUNX2related transcription factor, osteocalcin and osteopontin quantified by RT-PCR. *p≤0.05, **p≤0.01, ***p≤0.001.

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Figure 4. Immunofluorescence staining indicating ALP expression (A) after 7 days in cells exposed to different hydrogels, and OCN (B) and OPN (C) expression after 21 days in cells exposed to different hydrogels. The scar bars were 150 µm,

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3.4. Orthotopic bone formation in the composite hydrogels. A cranial bone defect model was used to evaluate bone regeneration capacity of the injectable SF-HA hydrogels. The 2 wt% SF nanofiber hydrogels containing various amounts of HA, or controls without HA, were injected into the circular defect of a rat calvarium. At 8, 12 and 16 weeks post-injection, rats were euthanized and defect sites were collected to assess the in vivo bone reconstruction.

The 3D structure of the repaired skull was investigated by micro-CT (Figure 5A). For all the hydrogels, the contour of the defect decreased gradually from 8 weeks to 16 weeks, suggesting new bone formation. Compared to the pure SF hydrogel control group, bone regeneration was significantly faster in the SF-HA composite hydrogel groups and further improved at higher HA contents, achieving the best bone reconstruction in the SF-HA60 hydrogels. For example, at 16weeks post-injection, the SF-HA60 hydrogels has been almost occupied by new formed bones while significant cavity remained in the SF and SF-HA20 hydrogels. Further quantitative analysis of the bone formation is shown in Figure 5B-5E. The bone volume (Figure 5B), bone volume/total volume ratio (BV/TV) (Figure 5C), trabecular number (Tb.N) (Figure 5D) and trabecular thickness (Tb.Th) data (Figure 5E) increased gradually in all cases, but showed different rates for the groups with various HA content. The highest values

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were achieved for the SF-HA60 group, followed by decreased HA contents and the pure silk nanofiber hydrogel, respectively. At 8 weeks, the new bone volumes were about 7.10±0.09 mm3 , 7.32±0.09mm3 , 8.41±0.26 mm3 and 11.61±0.33 mm3 with SF, SF-HA20, SF-HA40 and SF-HA60 hydrogels, respectively, which increased to 10.82±0.07 mm3, 11.71±0.03 mm3, 14.95±0.07 mm3 and 18.21±0.05 mm3 at 16 weeks. The average trabecular number per mm as well as trabecular thickness of the new formed bone was also significantly higher in SF-HA composite hydrogels than those with SF hydrogels. After 16 weeks implantation, the new formed bone had 0.87 ±0.4 trabecula per mm with thickness of 0.77 mm in SF-HA60 hydrogels but only 0.65±0.2 trabecula with thickness of 0.61 mm in the SF hydrogels.

HE staining and immunohistochemical analysis of samples were applied to clarify bone reconstruction behavior inside the hdyrogels (Figure 6). HE staining after 12 and 16 weeks implantation showed newly formed osteoid tissue, confirming new bone formation inside the hydrogels (Figure 6 A). Both OCN and OPN, two typical markers of osteogenic differentiation and mineralization, were intensively stained in all the SF-HA composite hydrogel groups at 12 and 16 weeks, further suggesting new bone formation (Figure 6B and 6C). Although quantitative comparison of the immunohistochemical results between the various hydrogels is qualitative, the in vivo

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results support bone regeneration in the SF-HA hydrogels and better bone defect healing with the SF-HA60 hydrogels. A summary showed that the silk-HA hydrogels with high HA content (SF-HA60) were suitable injectable matrices for in vivo bone defect reconstruction.

Figure 5. Micro-CT image analysis on new bone formation in rat calvarium model with SF-HA hydrogels: (A) Micro-CT images, (B) bone volume, (C) bone volume/total volume ratio, (D) trabecular number, (E) trabecular thickness. *p≤0.05, **p≤0.01, ***p≤0.001. 24


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Figure 6. Histological and immunohistochemical staining analysis of bone formation at 12 and 16 weeks post-implantation by HE staining (A), and immunohistochemistry staining of OCN (B) and OPN(C). The red arrows point to blood vessels, green arrows point to osteoblast, black arrows indicate osteoclast, respectively. NB shows new bone formation, while the brown illustrates positive expression of osseous proteins of OCN and OPN. The scale bars were 100 µm.

4. CONCLUSIONS The simulation of ECM was achieved by introducing water-dispersible HA

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nanoparticles into injectable silk nanofiber hydrogels, in which, silk fibroin provided ECM-mimetic nanofibrous structures while the water-dispersed HA nanoparticles resulted in high HA loading. The injectable SF-HA hydrogels possessed suitable HA content, microstructure and mechanical properties, out-performing silk-HA composite materials reported previously. The optimized SF-HA hydrogels demonstrated good cytocompatibility, osteogenic differentiation capacity in vitro, and supported bone formation in vivo, suggesting promising applications in bone tissue engineering and regeneration.

ACKNOWLEDGMENTS The authors thank National Basic Research Program of China (973 Program, 2013CB934400), NSFC (81272106) and the NIH (R01 DE017207, P41 EB002520). We also thank the Natural Science Foundation of Jiangsu Province (Grants No BK20140397, BK20140401) and the second affiliated hospital of Soochow university preponderant clinic discipline group project funding (NO.XKQ2015010) for support of this work.

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Abstract Graphic

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Nanoscale SilkâHydroxyapatite Hydrogels for Injectable Bone

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