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Bioinspired Composite Matrix containing HA-Silica Core–Shell Nanorods for Bone Tissue Engineering Anitha A, Deepthy Menon, Sivanarayanan TB, Manzoor Koyakutty, Chandini C Mohan, Shantikumar V Nair, and Manitha B Nair ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b07131 • Publication Date (Web): 25 Jul 2017 Downloaded from http://pubs.acs.org on July 26, 2017

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ACS Applied Materials & Interfaces

Bioinspired Composite Matrix containing HA-Silica Core–Shell Nanorods for Bone Tissue Engineering

Anitha A, Deepthy Menon, Sivanarayanan TB, Manzoor Koyakutty, Chandini C. Mohan Shantikumar V. Nair*, Manitha B. Nair* Center for Nanosciences and Molecular Medicine, Amrita University, Kochi, Kerala 682041, India

*Dr. Manitha Nair (Corresponding author), Center for Nanosciences and Molecular Medicine, Amrita University, Kochi, Kerala 682041, India. E-mail:

[email protected]

*Dr. Shantikumar Nair (Co-corresponding author), Center for Nanosciences and Molecular Medicine, Amrita University, Kochi, Kerala 682041, India. Kochi, Kerala 682041, India. E-mail:

[email protected]

ACS Paragon Plus Environment


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Keywords: Core-shell nanorods, silica, hydroxyapatite, vascularization, bone tissue engineering

Abstract Development of multifunctional bioinspired scaffolds that can stimulate vascularization and regeneration is necessary for application in bone tissue engineering. Herein, we report a composite matrix containing Hydroxyapatite (HA) - silica core shell nanorods with good biocompatibility, osteogenic differentiation, vascularization and bone regeneration potential. The biomaterial consist of crystalline, rod-shaped nanoHA core with uniform amorphous silica sheath (Si-nHA) that retains the characteristic phases of the individual components, confirmed by HR-TEM, XRD, XPS and FTIR. The nanorods were blended with gelatinous matrix to develop as a porous, composite scaffold. Viability and functionality of osteogenically induced mesenchymal stem cells as well as endothelial cells have been significantly improved through the incorporation of Si-nHA within the matrix. Studies in chicken chorioallantoic membrane and rat models demonstrated that the silica containing scaffolds not only exhibit good biocompatibility, but also enhance vascularization in comparison to the matrix devoid of silica. Finally, when tested in a critical sized femoral segmental defect in rat, the nanocomposite scaffolds enhanced new bone formation in par with biomaterial degradation. In conclusion, the newly developed composite biomimetic scaffold may perform as a promising candidate for bone tissue engineering applications.


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Introduction Advances in nanotechnology have ignited interest in nHA for bone tissue engineering, owing to its chemical composition and size matching to that of the native bone mineral crystals (Native bone contains HA nanorods having 30 to 200 nm length and 10 – 50 nm width).1 Porous nHA or its composites (nHA-gelatin / nHA-collagen) are osteoconductive and osteointegrative in nature. However, in critical sized bone defects, the survival and ingrowth of endogenous bone-forming cells on bioceramics is inadequate due to its lack of vascularization inducing potential.2 Most important task of blood vessels is oxygen and nutrient supply; guidance of osteogenic differentiation of mesenchymal stem cells (MSCs) and progenitor cells; the transport of phosphate and calcium for mineralization and further bone regeneration at the defect site.3 Several approaches have reported in order to enhance vascularization capability of a nHA scaffold4, of which cost effective, cell and growth factor free process through incorporation of trace elements like Sr, Zn, Fe, and Si in HA has gained much interest. 5

Amongst the trace elements present in bone extracellular matrix (ECM), the special role of silicon (Si) in angiogenesis and the early development of bone have been studied. Supplementation of Si in diet was found to hasten growth and skeletal development in chicks6 and improve bone mineral density in animals and humans.7,8Subsequently, Si substituted HA [phosphorus substitution by silicon in HA, Ca10(PO4)6-x(SiO4)x(OH)2-x) – SiHA] has been fabricated9 and is available commercially under the trade name Actifuse™, Baxter, UK. SiHA appears to have a better ability to absorb proteins and, thereby improved bioactivity when compared to HA.10Likewise, SiHA stimulated the proliferation of human umbilical vein endothelial cells (HUVEC), up-regulated the expression of pro-angiogenic factors (vascular endothelial growth factor - VEGF, basic fibroblast growth factor – bFGF), and activated endothelial nitric oxide synthase expression.11,12 However, crystalline structure and sintering properties of SiHA are influenced by Si content. Phase of Si– HA with low Si percentage is consistent with that of HA, but higher Si doping decreases crystallinity of HA and introduces impurity phase to the final product.13

The biological responses primarily start on the surface of the scaffolds and therefore, the surface modification of nHA by silica coating without altering its phase composition is proposed to elicit enhanced tissue regeneration. Previous studies have demonstrated the usefulness of nHA as tem-



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plating material for fabricating hollow silica nanostructures for drug or growth factor delivery applications.14-15 In those studies, nHA@SiO2 core-shell nanostructures were prepared by sol-gel method (Stober method) and subjected to etching processes to selectively remove core nHA and to develop hollow silica nanoparticles. The synthesis conditions can be tailored to produce silica nanoparticles of different sizes, shapes, crystallinity or porosity.

In this paper, the Stober method has been modiﬁed for preparing crystalline, rod-shaped nHA core with amorphous silica sheath (Si-nHA). These nanorods were dispersed in a gelatinous matrix (gelatin is the denatured form of native collagen) to develop a bio-inspired composite scaffold for bone tissue engineering. A detailed in vitro analysis of the viability, proliferation and functionality of osteogenically induced mesenchymal stem cells and endothelial cells was carried out. Finally, the ability of this composite scaffold to support vascularization and bone regeneration was analyzed in vivo. Although

2. Materials and Methods 2.1 Synthesis of core-shell nanorods HA nanorods were synthesized by a wet chemical synthesis route, following the previous reports by our own group.16 Briefly, 0.5M calcium chloride solution (Merck, USA) was added to 0.3M diammonium hydrogen phosphate solution (Merck, USA) at 90ºC and in a pH range of 10-12 under magnetic stirring. The precipitated HA was washed with distilled water by centrifugation at 3500 rpm and dried. Coating of silica over HA nanorods was carried out as reported elsewhere, with certain modifications.15 Briefly, HA powder was dispersed in absolute ethanol: water at a ratio of 1:0.4, to which ammonium hydroxide was added followed by mixing of varying concentrations of teraethylorthosilicate (TEOS) (0.1 mM, 0.4 mM and 0.8 mM). The solution was stirred for 3 h; centrifuged at 8000 rpm, washed with ethanol and water, dried and calcined at 550oC for 6 h. 2.2 Physicochemical characterization of core-shell nanorods The prepared powder was characterized for its morphology by Scanning Electron Microscope (SEM - JEOL, JSM-6490LA, Japan) and size / surface coating by Transmission Electron microscope (TEM - FEI Tecnai F20, JEOL, Japan) at an operating voltage of 15 kV and 300 kV respectively. Selected area electron diffraction (SAED) patterns of the nanorods were obtained using the


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ED attachment of TEM (microprobe mode). Fourier Transform Infrared (FTIR) Spectroscopy (Perkin Elmer Spectrum Rx-1, USA) in the wave number range of 4000–400 cm−1 confirmed the functional groups in the synthesized material. Crystallinity of the nanorods was verified by X-ray diffractometer (XRD) (PANanalytical X’Pert PRO, Germany), fitted with a CuKα source. XRD spectra were recorded in the range 5.0° to 80.0° at 0.015° intervals and phase identification was carried out with the help of the standard JCPDS database (JCPDS No: 09-0432). Surface composition analysis was performed by X-ray photoelectron spectroscopy (XPS) (Axis Ultra, Kratos Analytical, Shimadzu, Japan) with a focused monochromatic Al Kα source (1486.7 eV). XPS spectra in the binding energy range of 0-1200 eV were evaluated. Surface charge of the nanorods was evaluated by zeta potential analysis (Malvern Zetasizer nanoseries, USA). 2.3 Development and characterization of composite scaffolds The study comprises of two types of scaffolds (a) composite matrix with Si-nHA (CS) and (b) composite matrix with nHA (CH). The nanorods were blended with an aqueous solution of 15% wt/v of gelatin (Source: porcine skin Type-A gelatin, HiMedia, India) at a weight ratio of 65:35, under continuous stirring at 40°C for 1 h. The slurry was poured into suitable molds, which was lyophilized; cross-linked with 0.5% glutaraldehyde (Sigma Aldrich, USA) solution for 24 h at 37ºC; washed with distilled water 5-6 times and lyophilized again. The morphology and pore size distribution of composite scaffolds were analyzed by SEM (JEOL, JSM-6490LA, Japan) at an operating voltage of 15 kV. The elemental analysis was performed using the energy dispersive spectrometer (EDS) attachment of SEM (n=3). To determine the release of silicate species into the medium, the scaffolds (n=3) were immersed in phosphate-buffered saline (PBS - pH 7.4); incubated at 37ºC for different periods (day 1, 3, 7, 14 and 28) and the supernatant was analysed by Inductively coupled plasma atomic emission spectrometer (Thermo Electron IRIS INTREPID II XSP DUO). 2.4 In vitro cytocompatibility of composite scaffolds Rat adipose derived mesenchymal stem cells (ADMSCs) and HUVEC in passage 2-3 were used for cell culture studies after obtaining approval from the Institutional Animal and Human Ethical Committee. For isolating ADMSCs, the adipose tissue was digested with 0.5% type I collagenase in PBS for 30 min and centrifuged at 2500 rpm for 10 min. The cell pellet was cultured in αminimal essential medium (α-MEM, Invitrogen, India) with 10% fetal bovine serum (FBS, Invitrogen, India) and 1% antibiotic–antimycotic solution (Invitrogen, India) at 37°C in 5% CO2 in-



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cubator. These cells (5×104 cells) were seeded on ethylene oxide (ETO) sterilized CH and CS scaffolds (3 mm × 9 mm discs) and maintained in basal medium. After 24 h, the medium was supplemented with osteogenic supplements (10 mM β-glycerophosphate, 10−8 M dexamethasone, and 0.05 mg mL−1, L-ascorbic acid (Sigma Chemical Co., India) to induce the differentiation of stem cells into osteogenic lineage (MSC-OS). Endothelial cells were isolated from human umbilical cord as reported earlier.17 HUVECs were cultured on 2% gelatin (Sigma Aldrich, USA) coated culture flasks in Iscove's Modified Dulbecco's Medium (GIBCO, Invitrogen, USA) with 1% antibiotic–antimycotic solution and 150 µg / mL endothelial cell growth factor in a humidified CO2 incubator at 37°C and when it reached confluency, the cells were seeded onto the composite scaffolds as described earlier. Both MSC-OS and HUVEC were evaluated for viability, proliferation and functionality by following studies. 2.4.1 Cell viability After 24 h, the cells on scaffolds were washed with phosphate buffered saline (PBS), stained with calcein AM (4 mM in DMSO) and ethidium bromide [2 mM in DMSO/H2O 1:4 (v/v)] for 5 min and imaged by confocal fluorescence microscope (Leica TCS SP5 II) under green (480 nm / 520 nm) and red (605 nm / 635 nm) filters. For cell adhesion studies, the same scaffolds were washed with PBS, fixed with 1% glutaraldehyde in Sorenson phosphate buffer and dehydrated with increasing concentration of ethanol (50, 70 and 80% for 10 min twice, and 90, 100% for 15 min each). The samples were then air dried, sputter-coated with gold and viewed under SEM (JEOL, JSM-6490LA, Japan). 2.4.2 Cell proliferation The cell proliferation was determined using a picoGreen® dsDNA quantitation kit (Invitrogen, USA) according to the manufacturer’s instructions (n=6) (Day 1, 7, 14 and 21 for MSC-OS and Day 1 and 7 for HUVEC). For this, the cells were lysed with 1% Triton X-100 for 50 min with sonication for 10 min. An aliquot of each cell lysate (50 µL) was mixed with picogreen in TrisEDTA buffer (200 mM Tris-HCl, 20 mM EDTA, pH 7.5). The intensity of fluorescence was measured with a multifunction microplate reader (Beckman Coulter, DTX 880 multimode detector) at an excitation and emission wavelengths of 485 and 535 nm respectively. Relative fluorescence units were correlated with cell number using a calibration line constructed with increasing concentrations of cells.


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2.4.3 Osteogenic differentiation The osteogenic differentiation of MSC-OS was evaluated by alkaline phosphatase (ALP) activity and osteocalcin release in the medium. ALP activity (Day 7, 14, 21) was determined based on the hydrolysis of p-nitrophenyl phosphate (Sigma) to p-nitrophenol (n=6). For this, the cell cultured materials were lysed with 1% Triton-X-100. The lysate was added to 50 µL ALP reaction buffer (pH 9.8) (Sigma); incubated at 37°C for 30 min; added with 3N sodium hydroxide to stop the enzymatic reaction and the absorbance was read at 405 nm (Biotek Powerwave XS). A calibration line was plotted using different concentrations of ALP enzyme. The concentration of osteocalcin (ng / mL) in the medium in which cells were grown on scaffolds (n=3) was determined using ELISA technique (QnDsystems™-Rat Osteocalcin ELISA Kit). For this, the medium (Day 14, 21, 28) was collected and the level of osteocalcin was analyzed as per manufacturer’s instructions. The absorbance was recorded at 405 nm against a reference filter of 630 nm. A calibration curve was plotted and the osteocalcin concentration was determined with respect to calibration curve. 2.4.4 Endothelial differentiation The functionality of HUVEC was evaluated by nitric oxide (NO) and vascular endothelial growth factors (VEGF) assay (Day 1 and 7). The amount of NO in the medium, in which cells were grown on the scaffolds, was determined using Griess reagent (Sigma, USA) as per the manufacturer’s instructions. The absorbance was read at 540 nm (Biotek Powerwave XS). A calibration curve was plotted and the NO concentration in sample was analyzed with respect to calibration curve. Similarly, the concentration of VEGF secreted by the cells into the medium was determined using ELISA technique as per the manufacturer’s instructions (Raybio). The absorbance was read at 405 nm against a reference filter of 630 nm (Biotek Powerwave XS). A calibration curve was plotted and the VEGF concentration in sample was determined with respect to calibration curve. 2.5. Vasculariziation potential of composite scaffold 2.5.1 Chicken chorioallantoic membrane (CAM) assay The angiogenic potential of scaffold was analyzed by CAM assay based on protocol reported earlier.18 Fertilized white leghorn chicken eggs were obtained from Regional poultry farm, Kuruppampady, Cochin, India. Eggs were incubated at 37°C with 60% humidity. Albumin (2-3 mL) was aspirated using 21-gauge needle from a four day-old embryonated eggs and a small window



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was made in the shell under aseptic conditions. On day eight, ETO sterilized scaffolds (CH and CS) were placed onto the CAM of the eggs without disturbing the embryo. The window was resealed with adhesive tape and the eggs were returned to the incubator until day 12 of chick embryo development. Following the incubation period, the eggs were photographed and evaluated for angiogenesis under a stereoscopic microscope (Leica MZ 10 R, Germany). Eggs subjected to identical experimental conditions, but without scaffolds were served as the control. The length of vessels branched towards the scaffolds were quantified using the image analysis software Image J 1.47v (n=6).18 2.5.2 Subcutaneous implantation studies The care and management of animals were conducted as per the guidelines of the Institutional Animal Ethical Committee of Amrita Institute of Medical Sciences and Research Centre, Kochi, India. Wistar rats with an average body weight of 300 - 350 g and four defects per animal were divided into two groups: CH and CS (5mm dia X 1.5mm thickness) (n=4 in each group). The rats were anesthetized by giving an intramuscular injection of ketamine hydrochloride at 75 mg / kg body weight and xylazine hydrochloride at 5 mg / kg body weight. The skin was shaved and disinfected with betadine-povidone iodine. The quadriceps was exposed; the materials were implanted subcutaneously and the wound was closed with 4.0 mer silk sutures. The animals were administered with meloxicam (5 mg / kg body weight) analgesic and ceftriaxone antibiotic (20 mg / kg body weight) after the surgery for 3 days. The animals were kept in cages and fed with standard diet and water ad libitum. After 2 weeks, the animals were euthanized; the scaffold was explanted and fixed in 10% neutral buffered formalin in Sorensen phosphate buffer and analyzed for neovascularization. For this, the tissues were dehydrated in graded ascending series of acetone; embedded in paraffin and thin sections (5 µm) were made using a microtome (Leica, Germany). The sections after deparaffinising in xylene were rehydrated in descending grades of ethanol series and stained with hematoxylin and eosin Y. In addition, presence of CD31, which is an endothelial cell specific surface marker, was evaluated by immunohistochemical staining. Briefly, the deparaffinised sections were subjected to antigen retrieval; treated with CD31 monoclonal antibody (1:100, Novus, USA) for 1 h followed by HRP conjugated secondary antibody (1:150, Novus, USA) for 30 min at room temperature and with 100 µL substrate (DAB) reagent. The sections were dehydrated, mounted and imaged under light microscope (Olympus BX51). The blood vessel lumen area of CD31 positive samples was quantified using the image analysis software Image


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J (n=3). For this, CD31 positive areas (brown colour) in the blood vessel lumen were marked and the pixel intensities were quantified and compared between CS and CH groups.

2.6 Bone regeneration potential of composite scaffold 2.6.1 Animal surgical procedure 6-7 month old Wistar rats (n=4), weighing approximately 300 - 350 g, were used for the study. Method and procedures adopted for the management, surgery and care of the animals were approved by the Institutional Animal Ethical Committee of Amrita Institute of Medical Sciences and Research Centre, Kochi, India. The animals were maintained under anesthesia via intramuscular injection of ketamine hydrochloride at 75 mg / kg body weight and xylazine hydrochloride at 5 mg / kg body weight. Rats were placed in right lateral recumbent position, and a long transverse skin incision of approximately 2.5 cm running the full length of femur was created, exposing muscles, and facia. After removing periosteum, a 5 mm segmental defect was created in the mid diaphysis using a dental burr and the defect was stabilized with stainless steel (SS) mini plates with Bar (four holes) and SS mini screws (1.5 mm) (Ortho Max Mfg. Co. Pvt. Ltd, India) The scaffolds (CH / CS) were placed at the defect site, after which the defect was closed using vicryl sutures of size 4-0. Sham (defect without scaffolds) was used as the control. Meloxicam (5 mg / kg body weight) analgesic and ceftriaxone antibiotic (20 mg / kg body weight) was administered subcutaneously and intramuscularly respectively after the surgery for 5 days. The animals were euthanized after 4 months of implantation through an overdose of CO2 gas in a euthanasia chamber. Femoral tissue was harvested and fixed in 10% neutral buffered formalin. Radiographs of the defect site were taken immediately after surgical implantation (day 0) and at the end of 4 months using an in vivo imaging system (KODAK in-vivo Imaging Systems, Carestream, USA). 2.6.2 Histological and histomorphometric analysis For histological analysis, the post implanted samples were dehydrated in graded ascending series of alcohol, infiltrated, and embedded in polymethyl methacrylate (Merck). Longitudinal sections (100–120 µm) were made from each block using a high-speed precision saw (Isomet 5000 Precision Saw; Buehler, USA) and polished to a thickness of 50-70 µm. The sections were stained with Stevenal’s blue and van Gieson’s picrofuchsin and viewed under light microscope (10X, Olympus BX51). Histomorphometric analysis were done from calibrated stereomicrographs (Leica MZ 10 R, Germany) using the image analysis software Image J. The percentage area of the newly



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formed bone (orange-red by van Gieson’s picrofuchsin staining) and the material remained (bluish black) at the defect site were quantified by marking the total area of the defect (5 mm) as 100%.

2.6 Statistics Statistical analysis was carried out by one-way analysis of variance (ANOVA). All results are expressed as mean of all values ± standard deviation, and p-values less than 0.05 were considered significant for all analysis. 3. Results 3.1 Physiochemical characterization of core-shell nanorods Hydroxyapatite nanorods were prepared in high yield (50 g from 1L precursors) through a wet chemical synthesis procedure at 80ºC and alkaline pH conditions. Scanning electron micrographs demonstrated rod shaped nHA, but these structures formed sub-micron sized aggregates (Fig. 1A). The size of the nanorods was approximately 100 - 200 nm in length and 20 - 50 nm in width (Fig. 1B). Lattice spacing measured from the HR-TEM images (n=3) was around 0.28 nm, which correspond to the (211) plane of HA (Fig. 1C).19Electron diffraction pattern of nHA depicted in Fig. 1D shows the crystalline nature of the material with two distinct diffraction rings from the (211) and (002) lattice planes of HA. Elemental analysis of the synthesized nanorods confirmed through EDAX (inset of Fig. 1A, (n=6) displayed peaks relating to Ca, P and O, with Ca/P ratio as 1.6.

To obtain core-shells nanorods, the prepared nHA were subjected to TEOS coating through hydrolysis and subsequent polycondensation at alkaline conditions. A uniform silica deposition over the HA nanorods yielding a core-shell structure was evident from the representative TEM images shown in Fig. 2A-I. Increase in the concentration of TEOS led to an increase in the silica shell thickness, but with a concurrent formation of redundant silica nanoparticles in the solution. The average shell thickness for 0.1, 0.4 and 0.8 mM TEOS coated nHA was 5.6 ± 0.54, 12.2 ± 2.5, and 22.3 ± 3.1 nm respectively (Fig. 2D-F). From Fig. 2A-C, it could be clearly observed that at lower TEOS concentration (0.1 mM), condensation of silica occurred only around nHA core, with no additional nucleation of silica nanoparticles. However, with increased TEOS concentrations (0.4 and 0.8 mM), free silica nanoparticles also formed in addition to Si-nHA. Elemental analysis


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of Si-nHA illustrated the characteristic Ca and P peaks of HA together with a new Si peak, the intensity of which increased concurrently with increase in TEOS concentrations (Fig. 2G-I). The mass% of Si in nHA treated with 0.1, 0.4 and 0.8 mM TEOS was10.5±1.2, 22.2±2.4 and 34±5.1% respectively. Thus, the optimal concentration of TEOS (0.1 mM) that yielded a uniform surface coverage of silica over HA nanorods was selected for further studies.

HR-TEM of Si-nHA (Treated with 0.1 mM TEOS) clearly displayed the crystalline HA core and amorphous silica shell around it (Fig. 3). The corresponding ED patterns revealed two diffraction rings of the core that correlated with the lattice planes of nHA (Fig. 1D), while that of the shell showed only a diffuse halo. The intensity of the core rings was reduced as a consequence of silica coating. FTIR analysis of nHA revealed its characteristic PO43- peaks at 560, 603, 950, 1024 cm1 20

. With deposition of silica over nHA, a decrease in the phosphate stretching band at 950 cm-1

was observed. Furthermore, additional peaks at 1250 cm-1 and 800 cm-1 corresponding to the asymmetric and symmetric Si-O-Si vibration modes of SiO44- groups respectively were noted in Si-nHA (Fig. S1A). The characteristic asymmetric stretching vibrations of PO43- groups overlapped with the asymmetric stretching vibration band of Si-O-Si at ~1080 cm-1. The XRD spectrum

of

nHA

corresponded

well

with

the

crystalline

apatite

phase

for

HA,

while for Si-nHA hybrid, the overall crystallinity was reduced, specifically in the 10-25° region (Fig. 4A). There were no secondary phases generated in Si-nHA groups. XPS analysis was done to further confirm the coating of silica on nHA (Fig. S1B). Representative high resolution spectra collected from uncoated HA revealed typical C (1s - 284.65 eV), O (1s - 530.85 eV), Ca (2p3/2 347.05 eV & 2p1/2 – 350.65 eV)and P (2p3/2 – 132.95 eV) peaks. Nevertheless, in Si-nHA samples, an additional peak of Si (2p3/2) at 103.25eV was observed (Fig. 4B). The intensity of peaks of calcium and phosphorous of Si-nHA were reduced in intensity compared to that of nHA. The zeta potential values of nHA and Si-nHA was found to be -13.5±1.3 mV and -26.9±0.60 respectively (data not shown).

3.2 In vitro cytocompatibility of composite scaffolds To develop a bio-inspired scaffold for bone tissue engineering applications, the core-shell nanorods were incorporated in gelatinous matrix. Figure 5A shows the scanning electron micrographs of CH and CS, revealing their porous geometry, with pore sizes ranging from 50 to 300 µm. No



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significant differences in the pore size were noted between two groups of scaffolds. Elemental analyses of the scaffolds confirmed through EDS that displayed peaks relating to Ca, P, Si and O. The mass% of Si in the composite scaffold was found to be 6.15±0.8% (Fig. 5B). When the scaffolds were immersed in PBS, silicate species were released and its concentration was increased overtime and reached about 40 ppm within 28 days (Fig. 5C).

The osteogenic and endothelial cells were cultured on nanocomposite scaffold for evaluating their cytocompatibility in vitro. The cells remained viable and morphologically intact with extended filopodial extensions on CH and CS scaffolds (Fig. 6A&B). Both cell types appeared to be rimming the pores of scaffold, without pore occlusion or pore bridging. Quantitatively, proliferation of MSC-OS and HUVEC on composite scaffold was determined by picogreen assay (Fig. 6C). The number of osteogenic and endothelial cells was significantly higher on CS than CH scaffolds (p

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