JOURNAL OF TISSUE ENGINEERING AND REGENERATIVE MEDICINE RESEARCH J Tissue Eng Regen Med (2015) Published online in Wiley Online Library (wileyonlinelibrary.com) DOI: 10.1002/term.2023

ARTICLE

Nanostructured gellan and xanthan hydrogel depot integrated within a baghdadite scaffold augments bone regeneration Rekha R. Sehgal1, S. I. Roohani-Esfahani2, Hala Zreiqat2 and Rinti Banerjee1* 1

Department of Biosciences and Bioengineering, Indian Institute of Technology Bombay, Mumbai, India Biomaterials and Tissue Engineering Research Unit, School of Aerospace Mechanical and Mechatronic Engineering, University of Sydney, Australia 2

Abstract Controlled delivery of biological cues through synthetic scaffolds to enhance the healing capacity of bone defects is yet to be realized clinically. The purpose of this study was development of a bioactive tissue-engineered scaffold providing the sustained delivery of an osteoinductive drug, dexamethasone disodium phosphate (DXP), encapsulated within chitosan nanoparticles (CN). Porous baghdadite (BD; Ca3ZrSi2O9) scaffolds, a zirconia-modified calcium silicate ceramic, was coated with DXPencapsulated CN nanoparticles (DXP–CN) using nanostructured gellan and xanthan hydrogel (GX). Crosslinker and GX polymer concentrations were optimized to achieve a homogeneous distribution of hydrogel coating within BD scaffolds. Dynamic laser scattering indicated an average size of 521 ± 21 nm for the DXP–CN nanoparticles. In vitro drug-release studies demonstrated that the developed DXP–CN–GX hydrogel-coated BD scaffolds (DXP–CN–GX–BD) resulted in a sustained delivery of DXP over the 5 days (78 ± 6% of drug release) compared with burst release over 1 h, seen from free DXP loaded in uncoated BD scaffolds (92 ± 8% release in 1 h). To estimate the influence of controlled delivery of DXP from the developed scaffolds, the effect on MG 63 cells was evaluated using various bone differentiation assays. Cell culture within DXP–CN–GX–BD scaffolds demonstrated a significant increase in the expression of early and late osteogenic markers of alkaline phosphatase activity, collagen type 1 and osteocalcin, compared to the uncoated BD scaffold. The results suggest that the DXPreleasing nanostructured hydrogel integrated within the BD scaffold caused sustained release of DXP, improving the potential for osteogenic differentiation. Copyright © 2015 John Wiley & Sons, Ltd. Received 22 August 2014; Revised 20 January 2015; Accepted 23 February 2015

Keywords bone tissue engineering; chitosan nanoparticle; gellan; xanthan; hydrogel; baghdadite; dexamethasone; controlled drug delivery

1. Introduction Bone tissue engineering is emerging as a promising alternative to autografts and allografts, providing a biomimetic three-dimensional (3D) construct for bone healing (Meyer et al., 2004; Son et al., 2011; Salgado et al, 2004). While many conventional approaches, such as fibre bonding, solvent casting, particulate leaching, membrane lamination and melt moulding, are available for the synthesis of 3D

*Correspondence to: Rinti Banerjee, Department of Biosciences and Bioengineering, Indian Institute of Technology Bombay, Powai, Mumbai 400067, India. E-mail: [email protected] Copyright © 2015 John Wiley & Sons, Ltd.

porous scaffolds, the development of an ideal tissueengineered construct that provides the necessary bioactive signalling cues and can co-deliver stem cells for efficient differentiation is yet to be realized (Meyer et al., 2004; Hutmacher et al., 2007; Yang et al., 2011). Different strategies have been explored to improve cellular engrafting and to control the delivery of signalling molecules (Shi et al., 2010; Culpepper et al., 2010; Vo et al., 2012; Mooney and Vandenburgh, 2008) but still only few have reached clinical translation and they also show disadvantages in terms of poor cellular engraftment and insufficient control over the release profile of the signalling molecules (Bessa et al, 2008; Vo et al., 2012; Mooney and Vandenburgh, 2008). In this context, the use of a nanocarrier as a

R. R. Sehgal et al.

controlled drug delivery vehicle within a biocompatible nanostructured hydrogel could be a possible solution to control the release of bioactive agents. Encapsulation of cells and bioactive molecules inside the 3D nanostructured hydrogel matrices may provide a viable micro-environment for cell growth, similar to the extracellular matrix of many tissues (Geckil et al., 2010). In this study we applied a tissueengineering strategy in which dexamethasone disodium phosphate (DXP)-encapsulated CN (DXP–CN) nanoparticles and MG 63 osteoblast-like cells were entrapped within the porous matrix of a gellan and xanthan hydrogel. Finally, this nanoparticulate hydrogel was coated within a porous baghdadite (BD) scaffold to provide controlled delivery of cells and DXP. The BD scaffold (Ca3ZrSi2O9) was used as a framework to provide the essential stiffness and bioactivity to the scaffold. BD scaffold is well known for its in vitro and in vivo biocompatibility and bioactivity (Ramaswamy et al., 2008; Roohani-Esfahani et al., 2012). We hypothesized that coating the BD scaffold with DXP–CN–GX biopolymeric hydrogel could further improve the bioactivity of the BD scaffold through the controlled localized delivery of drugs. To the best of our knowledge, there are no reports in the published literature where simultaneous delivery of cells and osteoinductive drugs has been achieved within composite scaffolds using a nanoparticulated hydrogel. Chitosan as a biopolymer has been used in many tissueengineering applications because of its excellent biodegradability, good biocompatibility and antimicrobial properties (Levengood and Zhang, 2014; Oliveira et al., 2006; Custodio et al., 2010). Also, CN nanoparticles can improve the release kinetics of the drug through charge interactions, as well as cell adhesion and differentiation, through better control over the DXP dosage and release kinetics (Rajam et al., 2011; Guzmán-Morales et al., 2009). Similarly, dexamethasone (DEXA), a synthetic corticosteroid, can be a useful signalling molecule because of its anti-inflammatory properties and the modulation of mesenchymal stem cells (MSCs) differentiation at a nanomolar dose (GuzmánMorales et al., 2009). The existing literature has mainly focused on poly(lactic-co-glycolic acid)-based scaffolds for the controlled delivery of DEXA, involving the use of toxic organic solvents to solubilize the hydrophobic drug and initial higher drug doses to compensate for the amount of drug precipitated, hampering their clinical translation (Son et al., 2011; Yang et al., 2011; Shi et al., 2010). In this context, we have focused our present work on fabricating a biocompatible scaffold using a solvent-free method of entrapping the hydrophilic drug DXP within CN nanoparticles without the involvement of any harsh chemical, making it a feasible approach for clinical translation. In the present study, varying concentrations of gellan and xanthan gums with different molarities of calcium chloride as a crosslinker were optimized to fine-tune the gelation time, rigidity and viscoelasticity of the hydrogel, such that the incorporation of cells and DXP–CN nanoparticles was feasible at 37°C within gellan and xanthan solution (sol state), subsequently leading to the formation of homogeneously distributed hydrogel inside the BD porous matrix Copyright © 2015 John Wiley & Sons, Ltd.

within a few min. Gellan gum is an anionic polysaccharide that forms a thermo-reversible hydrogel with a setting point of 30–50°C, depending on the polymer concentration and the degree of acylation. Further, the addition of divalent, monovalent ions and positively charged crosslinkers such as CN nanoparticles allows modulation of the hydrogel’s mechanical strength (Grasdalen and Smidsrod, 1987; Oliveira et al., 2010). Xanthan gum is a non-gelling, anionic polysaccharide that forms a viscous solution at low shear rates (Rodriguez-Hernandez and Tecante, 1999). Hence, xanthan incorporation within gellan polymer can improve the viscoelastic behaviour of gellan hydrogels. The coating of such combinations of gellan, xanthan and chitosan nanoparticles, which can mimic the natural extracellular microenvironment, encapsulate cells and drug-loaded nanoparticles within its 3D network and can be fine-tuned for gelation and viscoelasticity by varying the crosslinker concentrations, provide new insights in the field of tissue engineering. Previous studies also have demonstrated the ability of gellan and xanthan gum to support the growth and differentiation of osteoblasts and stem cells for tissue-engineering applications (Dyondi et al., 2013; Silva-Correia et al., 2011; Carvalho et al., 2014).

2. Experimental 2.1. Materials The drug DXP, USP grade, was provided by Kirit Works Tablets (Mumbai, India). Gellan gum, in its deacetylated form (food grade, KELCOGEL), and xanthan gum (food grade) were bought from CP Kelco US (Chicago, IL, USA) and Loba Chemie Pvt. Ltd (India), respectively. Chitosan powder (>80% deacetylation) and sodium tripolyphosphate (TPP), used in the preparation of CN nanoparticles, were obtained from Marine Chemicals (India) and CDH Laboratories (New Delhi, India). Dialysis membrane (molecular weight cutoff 5000–10 000), Dulbecco’s modified Eagle’s medium (DMEM), fetal bovine serum (FBS), antibiotic–antimycotic solution, sodium dodecyl sulphate (SDS), Dulbecco’s phosphate-buffered saline (DPBS), trypsin–EDTA solution and alizarin red dye were purchased from Himedia Laboratories Pvt. Ltd (Mumbai, India). 3-(4,5-Dimethylthiazol-2yl)-2,5-diphenyltetrazolium bromide (MTT), calcein AM and propidium iodide, used to test cell viability and growth, were purchased from Sigma Aldrich (India). Alkaline phosphatase (ALP) assay kit and antibodies for osteocalcin and collagen type 1 marker were bought from Genxbio Health Sciences Pvt. Ltd and Allied Scientific, Kolkata (India), respectively. Unless stated otherwise, all other items were purchased from Sigma Aldrich (USA) and used as received.

2.2. Synthesis of DXP–CN nanoparticles and nanocomposite DXP–CN–GX–BD scaffolds Blank CN nanoparticles and DXP–CN (polymer:drug 4:1w/w) nanoparticles were synthesized by an ionic J Tissue Eng Regen Med (2015) DOI: 10.1002/term

Gellan–xanthan–baghdadite scaffold augments bone regeneration

crosslinking method (Calvo et al., 1997). BD scaffolds (Ca3ZrSi2O9; dimensions 10 × 10 × 3 mm) were prepared using the sponge template method, as previously described (Roohani-Esfahani et al., 2012). Detailed preparation methods for DXP–CN nanoparticles and BD scaffolds can be found in the supporting information (see Supplementary methods). Mixtures of gellan and xanthan (GX) hydrogels were prepared by an ionic crosslinking method (Oliveira et al., 2010). Briefly, gellan and xanthan (gellan:xanthan 9:1 w/w) polymers with different concentrations (0.5%, 0.7% and 0.9% w/v) were added into hot Milli-Q water (temperature 90°C) and stirred until the formation of a homogeneous dispersion. Subsequently, the solutions were allowed to cool at room temperature until a temperature of 37°C achieved. These solutions were encapsulated with DXP–CN nanoparticles (0.4%) and calcium chloride was added as a crosslinker in varying concentrations (1, 2 and 3 mM) at 37°C. Immediately after the addition of the crosslinker, 130 μl of the optimized polymer solution was poured inside the porous BD scaffold with the help of a pipette (Scheme 1).

2.3. Characterization of DXP–CN nanoparticles DXP–CN and CN nanoparticles were characterized for mean diameter and size distribution by photon correlation spectroscopy (PCS), using a laser particle analyser (BI 200SM, Brookhaven Instruments Corp., USA). The surface charge of the nanoparticles was determined using a zeta potential analyser (ZetaPALS, Brookhaven Instruments Corp.). Transmission electron microscopy (TEM; Philips CM200 instrument, operating at 200 kV) was used to visualize the morphology of the nanoparticles, using a negative staining method. The morphology of DXP–CN nanoparticles and hydrogel formation in the presence of optimized

CaCl2 crosslinker (0.5% GX hydrogel with 2 mM CaCl2, as discussed in section 3.2) was determined by atomic force microscopy (AFM) in tapping mode (Veeco Instruments Inc., Plainsview, NY, USA). To determine the encapsulation efficiency of DXP, a suspension of DXP–CN nanoparticles was centrifuged at speed of 16000 × g and the amount of free DXP in the supernatant was determined by spectrophotometry (Calvo et al., 1997) (λmax 245 nm, UV Spectrophotometer, Perkin-Elmer). To determine the possible interactions involved in the formation of CN nanoparticles, Fourier transform infrared spectroscopy (FTIR) was performed. The lyophilized samples were mixed with potassium bromide to make pellets. The FTIR spectra were scanned in the range 400–4000 cm–1.

2.4. Gelation time, rheological analysis, swelling and FTIR analysis of DXP–CN–GX hydrogels Varying percentages of gellan and xanthan (with or without DXP–CN nanoparticles), using different crosslinker concentrations, were evaluated for gelation time and for their distribution within the porous BD scaffolds (Table 1). After the addition of a crosslinker and DXP–CN nanoparticles at 37°C, 1 ml polymer solution was withdrawn to a test tube and inverted at different time intervals to check the gelation time, which is defined as the time when the polymer flow has stopped completely after inverting the tube. Hydrogels which gave feasible gelation times (according to Table 1) and had infiltrated within the BD scaffolds (visual observation) were further selected for rheological analysis. Dynamic oscillatory shear tests were performed, using a cone-plate rheometer (Physica MCR 301, Anton Paar, Germany) to determine the effect of crosslinker, DXP–CN nanoparticles and GX polymer concentrations on the viscoelastic

Scheme 1. Schematic diagram showing (A) the preparation of nanostructured DXP–CN–GX–BD scaffold and (B) optical images of baghdadite scaffold before and after coating with DXP–CN–GX hydrogel Copyright © 2015 John Wiley & Sons, Ltd.

J Tissue Eng Regen Med (2015) DOI: 10.1002/term

R. R. Sehgal et al. Table 1. Assessment of hydrogels gelation time and coating within the baghdadite scaffold

GX polymer concentration (0.4% DXP–CN in each case) 0.9% 0.7% 0.5%

Concentration of CaCl2 crosslinker (mM)

Gelation time of GX hydrogel (min)

Gelation time of DXP–CN–GX hydrogel (min)

Assessment of hydrogel infiltration within the BD scaffold

1 2 3 1 2 3 1 2 3

5 ± 0.5 3 ± 0.5 3 ± 0.5 10 ± 1 6±1 3 ± 0.5 18 ± 2 8±1 4 ± 0.5

4 ± 0.5 2±1 2±1 8±1 5±1 2 ± 0.5 16 ± 2 7±1 3±1

Gel on surface Gel on surface Gel on surface Infiltrated within pores Infiltrated within pores Gel on surface No hydrogel formation Infiltrated within pores Gel on surface

(a) (b) (c) (d) (e) (f) (g) (h) (i)

behaviour of the DXP–CN–GX hydrogels. The viscoelastic nature of the samples was determined using frequency sweep experiments (range 1–10 Hz at 37°C, constant strain of 0.5%). Further, the swelling rates of the hydrogels were checked by soaking the lyophilized hydrogels in 2 ml phosphate-buffered saline (PBS) solution, taken within Tarson tubes according to the standard method (Pereira et al, 2011). All tubes were kept in a shaking waterbath for 24 h at 37°C and 100 rpm. After a fixed time, swollen hydrogels were removed from the PBS solutions and weighed after removing the extra water from the surfaces of the hydrogels using filter paper. The swelling rate was calculated using the following equation:

polymer (Mandal et al, 2013). Briefly, samples of dimensions 1 × 1 × 0.3 cm were immersed in a graduated cylinder containing an already set volume (V1) of hexane. The samples were kept in the immersed condition for 5 min and then brief evacuation–repressurization cycles were conducted, two or three times, to allow the liquid hexane to pass within the pores. The total volume of the immersed scaffold and the hexane was recorded as V2. After removal of the scaffold, the remaining volume of hexane was recorded as V3. The percentage open porosity of samples was calculated as:

Swelling rate ð%Þ ¼ ðW t –W i Þ100=W i

Compressive strength and modulus were calculated using a computer-controlled universal testing machine (H5KS Materials Testing Machine, Tinius Olsen) having a 10 kN load. Samples of size 10 × 10 × 3 mm were compressed at a crosshead speed of 1 mm/min. Compressive modulus was measured as the slope of the linear portion of the stress–strain curve; each experiment was repeated four times and the data were calculated as average ± standard deviation (SD). The in vitro bioactivity of the DXP–CN–GX–BD scaffold was evaluated in a simulated body fluid (SBF), according to a standard protocol (Ramaswamy et al., 2008). Samples of size 10 × 10 × 3 mm were incubated with 20 ml SBF at 37°C. After a predetermined time, the samples were withdrawn from the solutions, washed with deionized water and lyophilized for further study. Apatite deposition was determined by SEM, FTIR and XRD studies and was further confirmed by inductively coupled plasma atomic emission spectroscopy (ICP–AES) analysis.

where Wt is the wet weight of the swollen hydrogel and Wi is the initial dry weight of the hydrogel. To check the possible interactions and the mechanism of hydrogel formation after the addition of the crosslinker and DXP–CN nanoparticles within the GX polymer solution, FTIR spectroscopy was performed for the optimized formulation, as discussed in section 3.2 (0.5% GX with 0.4% DXP–CN nanoparticles and 2 mM CaCl2). All different samples (gellan polymer, xanthan polymer, DXP–CN nanoparticles, lyophilized GX hydrogel and lyophilized DXP–CN– GX hydrogel) were mixed with KBr pellets to crush the powder and finally spectra were taken in the frequency range 400–4000 cm–1.

2.5. Characterization of hydrogel-coated (DXP– CN–GX–BD) and uncoated BD scaffolds: phase, pore morphology, porosity, compressive strength and bioactivity analysis The formation of pure BD scaffolds and the homogeneous coating of hydrogels were confirmed by X-ray diffraction (XRD) analysis (Siemens D5000, Germany). Microstructure and pore morphology of the samples was determined by scanning electron microscopy (SEM; Hitachi S-3400N; Hitachi Ltd, Tokyo, Japan). All samples were frozen at – 80°C, lyophilized and sputter-coated with gold before imaging. The porosity of different scaffolds was calculated by a liquid displacement method, using hexane as the displacement liquid, since it acts as a non-solvent for the Copyright © 2015 John Wiley & Sons, Ltd.

Porosity ð%Þ ¼ ½ðV 1 –V 3 Þ=ðV 2 –V 3 Þ100

2.6. In vitro release study In vitro release studies of DXP from DXP–CN–GX–BD scaffolds were conducted using the dialysis bag method (Feng et al., 2010). Briefly, DXP loaded samples were dialysed against 10 ml PBS solution in a shaking water bath (37°C, 100 rpm); 1 ml aliquots were withdrawn from the release medium at predetermined time intervals and the amount of DXP released was quantified, using a UV spectrophotometer at λmax = 245 nm. To show the effect of different components on the release profile, releases were also performed from control scaffolds, such as DXP-loaded BD scaffolds J Tissue Eng Regen Med (2015) DOI: 10.1002/term

Gellan–xanthan–baghdadite scaffold augments bone regeneration

(DXP–BD), DXP-loaded GX hydrogels (DXP–GX), DXP–CN– GX hydrogels and free DXP–CN nanoparticles.

2.7. In vitro cell culture All cell culture studies were performed using the MG 63 human osteosarcoma cell line (obtained from National Centre of Cell Science, Pune, India). MG 63 cells represent typical features of an undifferentiated osteoblast cell, such as the expression of alkaline phosphatase activity, collagen type 1 and osteocalcin, and they have also been shown to respond to 1,25-dihydroxyvitamin (Clover and Gowen, 1994). MG 63 cells were grown to > 80% confluency in a complete cell culture medium (DMEM + 10% FBS + 1% antibiotic–antimycotic solution) in sterile tissue culture flasks (25T and 75T) at 37°C, in 5% CO2 in a humidified atmosphere.

2.7.1. Cell encapsulation Before performing 3D cell encapsulation, the percentage cell viability of MG 63 cells with extracts of DXP–CN–GX– BD and BD scaffolds was calculated by MTT assay (for details, see supporting information, Supplementary methods, section 1.3). For 3D cell encapsulation, 80% confluent MG 63 cells were trypsinized and suspended in DXP–CN–GX polymer solution at 37°C before the addition of the crosslinker (105 cells/130 μl solution). Immediately after the addition of the crosslinker, DXP–CN–GX solution (130 μl) was poured in a homogeneous manner over all the surfaces of the BD scaffold (10 × 10 × 3 mm) for gelation, and finally transferred to 24 well plates for further experiments. DXP–CN nanoparticles doses within the hydrogels were adjusted (200× dilution) to ensure the required DXP concentration for osteoblastic differentiation (10–100 nM) and its toxicity level (1000 nM or > 400 ng/ml) (Kim et al., 2003). In all hydrogel controls [GX, DXP–GX, blank chitosan nanoparticles loaded with GX hydrogel (CN–GX) and DXP–CN–GX], 130 μl cell-loaded solution (with similar cell numbers; cell count 105 cells/130 μl) was directly poured into the wells of a 24-well plate. In the case of BD scaffolds only, 105 cells in 20 μl medium were seeded directly onto the porous matrix. To perform different experiments, all cell–scaffold constructs were incubated for 7, 14 and 21 days with 1 ml DMEM supplemented with 10% FBS and 1% antibiotic–antimycotic solution at 37°C and in 5% CO2 in a humidified atmosphere, with a change of medium every second or third day.

culture, according to a published protocol (Ferreira et al., 2007). Briefly, all samples were incubated for 4 h at 37°C with 200 μl 5 mg/ml MTT solution. Formazan was extracted by sonicating with 10% SDS solution (pH 3.7–4.1). The absorbance of the supernatant was measured using a plate reader (Thermo Electron Corp., Vantaa, Finland) at 560 nm, with 690 nm as the reference reading.

2.7.3. Osteoblast differentiation study The ALP activity of cultured MG 63 cells was determined using a SensoLyte® pNPP Alkaline Phosphatase Assay Kit (Anaspec, USA), according to the manufacturer’s protocol. A semi-quantification alizarin red-based assay (originally described by Gregory et al, 2004) was used to quantify the calcium secretion by cells. Briefly, scaffolds cultured with cells and control scaffolds without cells were incubated with 40 mM alizarin red solution, pH 4.1, for 20 min. After washing, alizarin red in the samples was extracted by treating with 10% acetic acid solution. The samples were neutralized with 10% ammonium hydroxide solution and colour absorbance was read at 405 nm using a plate reader (Thermo Electron Corp.). The absorbance of control scaffolds cultured without cells was subtracted from the absorbance of scaffolds cultured with cells and the results were normalized with respect to the total protein content. Immunohistochemistry on paraffin-embedded sections containing MG63 cell–scaffold constructs was performed at 21 days of culture. Briefly, all samples were fixed with 4% neutral-buffered paraformaldehyde solution. After fixation, BD-containing scaffolds (BD and DXP–CN–GX–BD) were decalcified in 20% ethylenediaminetetra-acetic acid (EDTA) for 48 h at 37°C. After the standard procedure of dehydration, paraffin embedding and sectioning, slides for immunohistochemical analysis were incubated with 4% bovine serum albumin, followed by primary antibodies for collagen type 1 and osteocalcin and their specific secondary antibodies, according to the standard procedure (Fan et al., 2010). All samples were observed under a confocal microscope at ×40 resolution. Details of methods for cell differentiation studies are given in Supplementary methods (see supporting information).

2.7.4. Statistical analysis All experiments were performed in triplicate and repeated at least three times. Results are expressed as mean ± SD. Statistical significance was assessed by performing oneway ANOVA; p < 0.05 was considered to be statistically significant in all experiments.

2.7.2. Cell viability and proliferation Live/dead staining assay was used to check the viability of cells inside the scaffolds over 21 days of cell culture, using calcein AM (1 μM) and propidium iodide (2 μM). Samples were visualized under a 3D confocal microscope (Olympus Fluoview, FV500, Tokyo, Japan) and an epifluorescent microscope (Axioskop2 MAT, Zeiss, Germany). Cell proliferation studies were performed by MTT assay over 14 days of Copyright © 2015 John Wiley & Sons, Ltd.

3. Results 3.1. Physicochemical characterization of nanoparticles DXP–CN nanoparticles showed a hydrodynamic diameter of 521 ± 21 nm (Figure 1C). TEM and AFM showed J Tissue Eng Regen Med (2015) DOI: 10.1002/term

R. R. Sehgal et al.

Figure 1. Physicochemical characterization of CN and DXP–CN nanoparticles and AFM micrographs of hydrogels. (A) TEM image of DXP–CN nanoparticles; scale bar = 400 nm. (B) Zeta potential of nanoparticles (mV); p < 0.05, n = 3. (C) Hydrodynamic diameter of nanoparticles (nm), as measured using DLS; p < 0.05, n = 3. (D) AFM micrograph of DXP–CN nanoparticles (2D) with (E) its section analysis; scale bar = 1 μm, height = 30 nm. (F, G) AFM micrographs (2D) of DXP–CN-loaded 0.5% GX hydrogel without crosslinker and 0.5% GX hydrogel with 2 mM CaCl2 crosslinker, respectively; scale bar = 1 μm

spherical, homogeneously distributed nanoparticles with a size range of 100–250 nm (Figure 1A, D). Nanoparticle size evaluated by TEM and AFM demonstrated that the DXP–CN nanoparticles size was smaller than the hydrodynamic diameter obtained by dynamic light scattering (DLS), due to swelling of the polymeric particles in the aqueous solution, as reported in earlier studies (Wassel et al., 2007). Both nanoparticles (DXP–CN and CN) showed high zeta potential (surface charge > 50 mV; Figure 1B), suggesting the stability of the nanoparticles. Different DXP:CN ratios were tested for optimizing the encapsulation efficiency, and DXP–CN nanoparticles (CN:DXP = 4:1 w/w) showed the highest encapsulation efficiency of 25 ± 3%. These nanoparticles were used for further experiments in this study. AFM images of CaCl2 crosslinked DXP–CN–GX hydrogel (Figure 1G) showed formation of a nanofibrous crosslinked network (0.5% GX, 0.4% DXP–CN with 2 mM CaCl2), while the uncrosslinked polymer without CaCl2 (Figure 1F) did not reveal any network formation. These results suggest that the GX polymer forms a nanostructured hydrogel (fibre diameter < 100 nm) by the addition of calcium ions, which provides the natural nanostructured environment Copyright © 2015 John Wiley & Sons, Ltd.

to encapsulated cells growing within their 3D network and may help in their growth. To check the interaction of chitosan with DXP and TPP in DXP–CN nanoparticles, FTIR spectroscopy was performed (see supporting information, Figure S4). The FTIR spectrum of chitosan gave a broad peak at 3100–3700 cm–1, which is due to the stretching vibration of –O–H and –NH2 groups. Also, it showed a broad peak at 1657 cm–1 that may reflect the bending vibrations of the primary –NH2 group and the carbonyl group of the amide bond. In the spectrum of blank CN nanoparticles, the broad peak of chitosan at 1657 cm–1 was split into two peaks at 1549 cm–1 (crosslinked NH2 bending vibrations) and 1651 cm–1 (amide bond), due to ionic interactions with the phosphate group of TPP. Also, it showed a new peak at 1752 cm–1 that could be due to the presence of protonated –NH+ 3 groups. In the FTIR spectrum of DXP–CN nanoparticles, the peak at 1549 cm–1 became more intense compared to 1651cm–1, which showed the possible interaction of negatively charged DXP with the –NH2 group of chitosan. Also, the intensity of the peak at 1752 cm–1 was reduced, suggesting the interaction of free protonated chitosan chains with DXP. The broad peak of –OH stretching vibrations in J Tissue Eng Regen Med (2015) DOI: 10.1002/term

Gellan–xanthan–baghdadite scaffold augments bone regeneration

chitosan (3100–3700 cm–1) became wide and more intense in DXP–CN nanoparticles, showing an increase in hydrogen bonding. These results suggest that both TPP and DXP interact with the acidic solution of chitosan through charge interactions and hydrogen bonding, finally giving solid nanoparticles.

3.2. Gelation time, swelling, rheological analysis and optimization of hydrogel-coating parameters Table 1 demonstrates that the gelation time of the GX and DXP–CN–GX hydrogels decreased with increasing crosslinker and GX polymer concentrations. The gelation times of DXP–CN–GX hydrogels were lower than those of the corresponding GX hydrogels with the same crosslinker and polymer concentration. These results suggest that both CaCl2 and DXP–CN nanoparticles are playing a role in the gelation of GX polymer. Of the different hydrogels tested, 0.5% gel with 2 mM CaCl2 and 0.7% gel with 1 and 2 mM CaCl2 (Table 1d, e, h) infiltrated homogeneously within the porous BD scaffold. Hydrogels with higher GX concentrations and high crosslinker concentrations (Table 1a, b, c, f, g) could not infiltrate within the BD scaffold and formed a hydrogel layer on the scaffold surface; this could be because of their instant gelation on coming into contact with the BD scaffold. Those hydrogels (DXP–CN–GX) that formed a homogeneous coating were further investigated for their viscoelastic nature. 0.5% GX hydrogel without DXP–CN nanoparticles crosslinked with 2 mM CaCl2 was also investigated by rheology analysis, to assess the effect of nanoparticles on viscoelastic properties. Typically, viscoelastic hydrogels have a storage modulus, G′, that is always higher than the loss modulus, G′′ (G′ > G′′), which represent the elastic and viscous behaviours of hydrogels, respectively. All CaCl2crosslinked DXP–CN–GX hydrogels had a storage modulus (G′) higher than the loss modulus (G′′) over the entire frequency range tested, confirming the viscoelastic nature of the hydrogels (Figure 2A). The storage modulus of

DXP–CN–GX hydrogel without CaCl2 was almost equal to its loss modulus (G′′ ≈ G′) and its value was almost 103 times less than the storage modulus (G′) of DXP– CN–GX hydrogel with crosslinker (Figure 2A), suggesting the role of the crosslinker in forming a viscoelastic hydrogel. DXP–CN–GX hydrogels with 2 mM crosslinker (0.5% and 0.7%) showed higher elasticity compared to the 1 mM (0.7%) CaCl2 DXP–CN–GX hydrogel. On comparison, the storage modulus of the 0.5% DXP–CN–GX hydrogel was higher than that of the GX hydrogel with same crosslinker concentration (see supporting information, Figure S3), further demonstrating the role of DXP–CN nanoparticles in improving the viscoelasticity of DXP–CN–GX hydrogels. Because of a suitable gelation time (7 ± 1 min) and a high viscoelasticity, 0.5% DXP–CN–GX hydrogel crosslinked with 2 mM CaCl2, which formed a homogeneous coating within the BD scaffold, was selected for further studies. Figure 2B demonstrates the percentage swelling rate of different DXP–CN–GX hydrogels after 24 h of incubation with PBS at the physiological condition (37°C, 100 rpm). The percentage swelling of hydrogels decreased with increasing GX polymer and the crosslinker concentrations, with the highest swelling rate observed with 0.5% DXP– CN–GX hydrogel with 1 mM CaCl2 (6471 ± 300%). We have calculated the percentage swelling rate of DXP– CN–GX–BD scaffold (1000 ± 243%; BD scaffold coated with 0.5% DXP–CN–GX hydrogel with 2 mM CaCl2), which was five times less than that of the hydrogel with the same concentrations (0.5% DXP–CN–GX, 2 mM CaCl2; 5100 ± 117% swelling), suggesting enhancement in the hydrogel crosslinking with the BD scaffold.

3.3. Mechanism of DXP–CN–GX hydrogel formation To check how different components of the hydrogel take part in formation of the hydrogel and affect its gelation, FTIR spectroscopy was performed with the final optimized formulation (as discussed in section 3.2); the results are given in Figures S4–S6 (see supporting information). The

Figure 2. (A) Storage modulus (G′) vs loss modulus (G′′) of different concentrations of DXP–CN–GX hydrogels. (a) G′ of 0.5% gel with 2 mM CaCl2; (b) G′ of 0.7% gel with 1 mM CaCl2; (c) G′ of 0.7% gel with 2 mM CaCl2; (d) G′ of 0.5% gel without CaCl2; (e) G′′ of 0.5% gel with 2 mM CaCl2; (f) G′′ of 0.7% gel with 1 mM CaCl2; (g) G′′ of 0.7% gel with 2 mM CaCl2; (h) G′′ of 0.5% gel without CaCl2. (B) Effect of varying the crosslinker and GX polymer concentration on the percentage swelling rate of DXP–CN–GX hydrogels Copyright © 2015 John Wiley & Sons, Ltd.

J Tissue Eng Regen Med (2015) DOI: 10.1002/term

R. R. Sehgal et al.

FTIR spectrum of the blank gellan powder showed a characteristic peak of carboxylate group stretching vibrations at 1619 cm–1 and stretching vibrations of the hydrogen bonded –OH group at 3422 cm–1. Similarly, xanthan powder exhibited characteristic peaks at 3434 cm–1 (–OH stretching), 1730 cm–1 (carbonyl group of ester) and 1620 cm–1 (carboxylate anion). In the FTIR spectrum of the GX hydrogel crosslinked with CaCl2, the carboxylate peak of gellan and xanthan appeared at 1630 cm–1 and the carbonyl group of esters in xanthan showed a peak at 1730 cm–1. Other than these two peaks, one extra peak at 1798 cm–1 has been observed. This may be due to stretching vibrations of a carboxylate group crosslinked with calcium ions. Other than that, increase in the intensity of the broad peak at 3430 cm–1 compared to pure polymers suggests possible enhancement in hydrogen bonding within the hydrogel. When DXP–CN nanoparticles were added into the GX hydrogel, the intensity of the extra peak at 1798 cm–1 was reduced (see supporting information, Figure S6) in DXP–CN–GX hydrogel compared to GX hydrogel, which indicates less ionic interaction of the GX carboxylate group with the calcium ions. Other than that, stretching vibrations of the carboxylate group at 1630 cm–1 in the GX hydrogel, bending vibrations of the –NH2 group at 1538 and 1558 cm–1 in DXP–CN nanoparticles and stretching vibrations of amide in DXP–CN nanoparticles merged into a broad peak, in the range 1612–1650 cm–1, in the DXP–CN–GX hydrogel. All these results suggest that crosslinking sites of the GX polymer are now distributed between calcium ions and positively charged DXP–CN nanoparticles. These results were also confirmed by rheology analysis and gelation time measurements, which showed a higher storage modulus of DXP–CN–GX hydrogel than GX hydrogel (see supporting information, Figure S3) and reduction in the gelation time (Table 1) in the presence of nanoparticles, suggesting the involvement of DXP–CN nanoparticles in formation of the hydrogel. A possible mechanism of hydrogel formation in the presence of calcium chloride crosslinker and DXP–CN nanoparticles is given as a schematic diagram (Scheme 2). It is well known that the gellan polymer forms a hydrogel in the presence of the crosslinker at low temperature (double helical formation through hydrogen bonding and further ionic crosslinking of carboxylate groups present in the double helix with positively charged ions) (Chandrasekaran and Thailambal, 1990). Xanthan is found in solution in a rigid chain conformation that also has negatively charged carboxylate groups (Rodriguez-Hernandez and Tecante, 1999). In the presence of calcium ions, one gellan double helix can ionically crosslink with the other gellan helix, or it can crosslink with the polymeric chain of xanthan and finally forms GX hydrogel. In the presence of DXP– CN nanoparticles, the GX polymer can also interact with the surface of the positively charged nanoparticles, as shown in Scheme 2. When this DXP–CN–GX hydrogel is coated within the baghdadite scaffold, it can further interact with the free calcium ions present on the scaffold surface. Copyright © 2015 John Wiley & Sons, Ltd.

3.4. Physicochemical characterization of hydrogel-coated (DXP–CN–GX–BD) and uncoated BD scaffolds: phase, pore morphology, porosity and bioactivity analysis XRD spectra of BD scaffolds revealed the formation of pure Ca3ZrSi2O9 phase (Ramaswamy et al., 2008), while coated DXP–CN–GX–BD scaffolds did not show any characteristic peaks of BD on the surface (Figure 4B), demonstrating the formation of a homogeneous coating of hydrogel over the BD scaffold. SEM images (Figure 3) demonstrate that the DXP–CN–GX hydrogel, the BD scaffold and the BD scaffold coated with DXP–CN–GX hydrogel formed a highly porous, interconnected network with pore sizes in the range 200–800 μm (Figure 3A, C, D). Also, CN nanoparticles of size 100–200 nm were homogeneously dispersed inside the hydrogel (Figure 3B). The calculated percentage porosity of the BD scaffold was 79.30 ± 0.30%, which reduced to 76.27 ± 1.20% in the case of the coated DXP– CN–GX–BD scaffold (Table 2), suggesting that the coating phenomenon did not significantly impact the porosity of the scaffold. Cells and nutrients can easily migrate inside the macroporous structure (200–400 μm), a crucial step for the survival and growth of the encapsulated cells. Compressive tests were conducted for pure hydrogels (GX and DXP–CN–GX), uncoated BD scaffolds and the coated DXP–CN–GX–BD scaffolds and the results are shown in Table 2. The compressive strength of the coated DXP– CN–GX–BD scaffold was 2.5-fold higher than that for the uncoated BD scaffold and five-fold that of the the DXP– CN–GX hydrogel. A similar trend was obtained for compressive modulus, which was 23.10 ± 3.45 MPa for the DXP– CN–GX–BD scaffold (2.4-fold higher than the uncoated BD scaffold and 40-fold higher than the hydrogel). Bioactivity is an important property that enhances osteointegration within the scaffold through the formation of layers of apatite at the interface between the scaffold and the bone. Figure 4A shows SEM images of DXP–CN–GX–BD scaffolds at 0, 8 and 16 days of incubation in SBF solution. It can be clearly seen that the apatite started depositing on the surface of the scaffold after 8 days of incubation, and the area of deposition progressed as the incubation time increased, with the maximum deposition observed after 16 days of incubation. XRD analysis showed peaks at 2θ value of 29, 31 and 46 after 16 days of SBF incubation that were absent in day 0 control scaffolds, which further proves the deposition of biomimetic apatite on the DXP–CN–GX–BD scaffold (Figure 4B). The hydrogel coating and apatite deposition on the DXP–CN–GX–BD scaffold surface may potentially mask peaks of the BD scaffold after 16 days of incubation in SBF solution. Apatite deposition was further confirmed by FTIR and ICP–AES analysis. Figure 4C shows FTIR spectra of DXP–CN–GX–BD scaffolds at 0, 8 and 16 days of SBF incubation. All samples showed a characteristic peak of phosphate in the regions 580 cm–1, 620 cm–1 and 960–1050 cm–1, the intensity of which increased with increased incubation time, confirming the deposition of apatite over the scaffold surface. Also, the peak at 870cm–1 suggests the formation of carbonated J Tissue Eng Regen Med (2015) DOI: 10.1002/term

Gellan–xanthan–baghdadite scaffold augments bone regeneration

Scheme 2. Possible mechanism of DXP–CN nanoparticles and DXP–CN–GX hydrogel formation. (A) Mechanism of DXP–CN nanoparticles formation. (B) Molecular structure of low-acyl gellan gum. (C) Molecular structure of xanthan gum. (D) Proposed mechanism of DXP–CN–GX hydrogel formation. (E) FTIR spectrum of DXP–CN–GX hydrogel

hydroxyapatite (Hoppe et al., 2013). ICP–AES analysis showed a decrease in phosphorus concentration in SBF solution, which was directly proportional to the increase in duration of incubation. Similarly, there was an initial increase in calcium concentration, which was then decreased slightly with increased incubation time. These results demonstrate that the calcium released from the DXP–CN–GX–BD scaffold formed a SBF solution which was supersaturated with respect to apatite, leading to the precipitation of calcium and phosphate in the form of apatite. Further, the negatively charged surface of the hydrogel provides sites for the initial nucleation of apatite. Once initiated, these apatite molecules can grow further until a condition of equilibrium is reached (Figure 4D). These results collectively signify the bioactivity of the developed DXP–CN–GX–BD scaffolds. Copyright © 2015 John Wiley & Sons, Ltd.

3.5. In vitro release experiments DXP in vitro release was determined using the dialysis bag method and the results are shown in Figure 5. DXP release from the control DXP–GX hydrogels and DXP–BD scaffolds demonstrated burst release, with 88.65 ± 6.5% of the DXP released from DXP–GX hydrogels within 4 h and 92 ± 8% from DXP–BD scaffolds within 1 h. A more sustained DXP release was observed from chitosan nanoparticle-based formulations, i.e. DXP–CN nanoparticles, DXP–CN–GX hydrogels and hydrogel-coated DXP–CN–GX– BD scaffolds, which showed 63.7 ± 3%, 39.4 ± 10% and 17.8 ± 3.34%, release, respectively, within 4 h (Figure 5C). After the initial release, DXP–CN nanoparticles showed 90 ± 6% release over a period of 30 h. This release pattern of the DXP from DXP–CN nanoparticles was more sustained J Tissue Eng Regen Med (2015) DOI: 10.1002/term

R. R. Sehgal et al.

Figure 3. Cross-sectional SEM analysis of hydrogels and scaffolds. (A, B) DXP–CN–GX hydrogel at different magnifications. (C) BD scaffold. (D) DXP–CN–GX–BD scaffold

Table 2. Assessment of porosity, compressive strength and compressive modulus of BD scaffold before and after hydrogel coating Samples 0.5% GX, 2 mM CaCl2 0.5% DXP–CN–GX, 2 mM CaCl2 BD scaffold DXP–CN–GX–BD scaffold

Porosity (%)

Compressive strength (MPa)

Compressive modulus (MPa)

93.20 ± 0.34 91.05 ± 1.10 79.30 ± 0.30 76.27 ± 1.20

0.031 ± 0.01* 0.034 ± 0.01* 0.06 ± 0.02* 0.15 ± 0.02

0.55 ± 0.03* 0.50 ± 0.17* 9.57 ± 0.62* 23.10 ± 3.45

*Significantly different compared to DXP–CN–GX–BD scaffold (p < 0.05) (n = 4).

Figure 4. Physicochemical characterization of DXP–CN–GX–BD scaffolds in SBF. (A) SEM images after (a) 0, (b) 8 and (c) 16 days in SBF; scale bar = 20 μm. (B) XRD spectra of (a) BD scaffold and DXP–CN–GX–BD scaffolds after (b) 0, (c) 8 and (d) 16 days in SBF. (C) FTIR spectra of DXP–CN–GX–BD scaffolds in SBF. (D) ICP–AES analysis of 0, 8 and 16 days DXP–CN–GX–BD scaffolds in SBF Copyright © 2015 John Wiley & Sons, Ltd.

J Tissue Eng Regen Med (2015) DOI: 10.1002/term

Gellan–xanthan–baghdadite scaffold augments bone regeneration

Figure 5. In vitro release kinetics of DXP from different types of hydrogels and scaffolds. (A) Cumulative increase in concentration of the DXP from samples in PBS solution at 37°C. (B) Cumulative release (%) of DXP from samples in PBS at 37°C. (C) Comparison of # cumulative release (%) as observed after 4 h at 37°C for different samples in PBS; different to DXP–GX controls (p < 0.05); *different to DXP-BD scaffolds (p < 0.05, n = 3)

when the nanoparticles were encapsulated within hydrogels (DXP–CN–GX) and hydrogel-coated BD scaffolds (DXP– CN–GX–BD) compared with direct release from DXP–CN nanoparticles. DXP–CN–GX–BD scaffolds and DXP–CN–GX hydrogels revealed DXP sustained release of 78 ± 6% and 83.5 ± 7.7% in 126 and 112 h, respectively.

all Z scan images, demonstrating that the hydrogel and hydrogel-coated BD scaffold provide a conducive environment that supports cell viability inside its hydrophilic environment (Figure 6). The viability of cells was further confirmed by capturing epifluorescent microscopic images of viable cells at low resolution (×2.5) after 21 days of cell incubation (Figure 6E–H).

3.6. Cell viability

3.7. Cell proliferation study

The percentage viability of MG 63 cells with hydrogelcoated DXP–CN–GX–BD scaffolds and uncoated BD scaffolds extract was found to be > 90% after 24 h of incubation, suggesting the biocompatibility of the scaffolds for further cell encapsulation (see supporting information, Figures S1, S2). To check the viability of cells within scaffolds after 3D encapsulation, a live/dead assay was performed after 21 days of cell incubation. Confocal images in Figure 6A–D demonstrate that the majority of the cells were viable (green fluorescence) within the different constructs tested, with very few dead cells (red fluorescence), suggesting that the scaffolds provide a healthy and biocompatible environment for cell growth. The cells maintained a round morphology within hydrogel-based scaffolds (GX and DXP–CN–GX) and converted to the normal spindle-shaped morphology in the presence of BD and DXP–CN–GX–BD scaffolds, suggesting the role of BD scaffold stiffness in maintaining cell morphology. Viable cells were evident in

Proliferation of MG 63 cells was performed by MTT assay. A significant (p < 0.05) increase in cell proliferation was evident when cells were cultured for 7 and 14 days on CN–GX scaffolds, compared to GX scaffolds (Figure 7A). These results suggest that the addition of CN nanoparticles inside a GX scaffold plays an active role in cell growth. Figure 7A demonstrates that hydrogel-coated scaffolds (DXP–CN–GX–BD) had statistically (p < 0.05) more cell proliferation compared to BD-only scaffolds after 7 days of culture. The results obtained indicate that the coating of DXP–CN–GX hydrogel inside a BD scaffold improves earlier cell proliferation of MG 63 cells compared to the BD-only scaffold.

Copyright © 2015 John Wiley & Sons, Ltd.

3.8. Osteoblastic differentiation study The ALP activity of cells was statistically higher (p < 0.05) with the DXP–CN–GX hydrogel after 7 days of incubation J Tissue Eng Regen Med (2015) DOI: 10.1002/term

R. R. Sehgal et al.

Figure 6. Live/dead staining of MG 63 cells with calcein AM (green, live cells) and propidium iodide (red, dead cells) grown within hydrogels and scaffolds after 14 days (confocal images and Z scan) and 21 days (epifluorescent images) of incubation: confocal images, scale bar = 200 μm, resolution = ×10; epifluorescent images, scale bar = 200 μm, resolution = ×2.5; Z scan step size, 60 μm up to depth of 300 μm and resolution = ×10

compared to other hydrogels (GX, DXP–GX, CN–GX) (Figure 7B), while DXP–GX hydrogel did not promote ALP activity, compared to GX hydrogel, at any stage of proliferation. These results suggest that the controlled delivery of DXP from the nanoparticulated DXP–CN–GX hydrogel promoted the earlier differentiation of MG 63 cells. Furthermore, coating of this DXP-delivering hydrogel (DXP–CN–GX hydrogel) inside the porous BD scaffold (DXP–CN–GX–BD scaffold) significantly enhanced the ALP activity of MG 63 cells after 7 and 14 days of culture, compared to the BD-alone scaffold (p < 0.05), verifying that the coating of hydrogel improved the differentiation capability of the BD scaffold. Calcium deposition is an important feature of bone cell differentiation. Figure 7C shows that cells grown on GX hydrogel deposited the least amount of calcium compared to all other scaffolds (p < 0.05) after 21 days of cell incubation. Moreover, calcium deposition by cells cultured within the DXP–CN–GX hydrogel was statistically Copyright © 2015 John Wiley & Sons, Ltd.

greater (p < 0.05) compared to DXP–GX and CN–GX hydrogels (Figure 7C). These results confirm the role of sustained delivery of DXP and also polymeric chitosan nanoparticles in mineral deposition by MG 63 cells (Figure 7C). Coating of these hydrogels (DXP–CN–GX) with the high calcium deposition capability inside porous BD (DXP–CN–GX–BD) scaffolds resulted in an approximately 50% increment in cell calcium deposition after 21 days of incubation, compared to blank BD scaffolds (p < 0.05) (Figure 7C). To work as a bone tissue-engineered graft, a scaffold should be able to maintain the cell phenotype and should be able to secrete bone-specific markers. In order to evaluate the gene expression of extracellular bone-specific markers from MG 63 cells grown on different scaffolds, fluorescent immunostaining techniques were performed. Osteocalcin and collagen type 1 are well known proteins secreted by mature bone cells. Osteocalcin (Figure 8Aa–f) and collagen type 1 (Figure 8Ag–l) immunolocalization J Tissue Eng Regen Med (2015) DOI: 10.1002/term

Gellan–xanthan–baghdadite scaffold augments bone regeneration

Figure 7. (A) Effect of different scaffold components on MG 63 cell proliferation and (B) alkaline phosphatase activity of MG 63 cells seeded within them after 7 and 14 days of culture. (C) Semiquantification of calcium deposition of MG63 cells seeded within different hydrogels and scaffolds for 7 and 21 days. (D) Optical images of different hydrogels and scaffolds stained with alizarin red after 21 days of culture; (a, c, e, g) MG 63 cultured GX hydrogel, DXP–CN–GX hydrogel, BD scaffold and DXP–CN–GX–BD scaffold, respectively; # (b, d, f, h) their respective controls without cell seeding. Different to GX control (p < 0.05); *different to BD scaffold (p < 0.05); ×× different to DXP–GX and CN–GX hydrogel (p < 0.05); n = 4

is depicted using confocal images of different types of cellcultured scaffolds. Minimal fluorescence intensities for osteocalcin and collagen type 1 were obtained with the GX hydrogel, which increased significantly in other hydrogels (DXP–GX, CN–GX and DXP–CN–GX hydrogel) compared to the GX hydrogel (Figure 8B, C). The coating of DXP–CN–GX hydrogel over the BD scaffold had a significant and more pronounced effect on osteocalcin and collagen type 1 secretion.

4. Discussion Bone loss is related to high economic burden because of the scarcity of autografts. As an alternative to that, the use of a tissue-engineering scaffold encapsulating osteoblasts or stem cells and bioactive signalling molecules, which can guide the growing cells towards osteodifferentiation, is an area of intense research (Bessa et al., 2008). Critical parameters for the delivery of these molecules within scaffolds include their short half-lives and the maintenance of biological activities (Bessa et al., 2008). Clinically available tissue-engineered scaffolds, such as INFUSE® and OP-1®, have shown promising results in a few cases but, due to poor control over biological activity, high Copyright © 2015 John Wiley & Sons, Ltd.

initial doses of growth factors are required to trigger stem cells towards osteodifferentiation, which in turn increases the cost of the scaffolds (Vo et al., 2012). Therefore, the present study explored the use of a chemically stable, osteogenic drug, DXP, to promote the differentiation of osteoblasts to overcome the above-mentioned issues. Although DEXA is effective at very low concentrations (10–100 nM; Guzmán-Morales et al., 2009), its long-term use at high concentrations (>400 ng/ml) has been shown to cause severe toxicity (Kim et al., 2003), emphasizing the importance of limiting cell exposure to high levels of DEXA. Major drawbacks in the existing literature are the use of organic solvents (Son et al., 2011; Kim et al., 2003), high doses of DEXA (Son et al., 2011; Kim et al., 2003), the exposure of cells to UV light (Nuttelman et al., 2006) and the use of nanoparticles for growth of MSCs in culture medium without encapsulation within the scaffold (Oliveira et al., 2009), which are limiting their clinical translation. The feasibility of incorporating biocompatible hydrophilic DXP–CN nanoparticles and osteoblast cells within the nanostructured 3D matrix of gellan–xanthan blend hydrogel at 37°C, and its coating within the BD scaffold, has been investigated in this study towards the development of an ideal bone tissue-engineered scaffold. In our study, we used a thermoresponsive biocompatible hydrogel J Tissue Eng Regen Med (2015) DOI: 10.1002/term

R. R. Sehgal et al.

Figure 8. Immunohistochemistry of scaffolds cultured with MG 63 cells for 21 days. (A) Confocal images showing production of osteocalcin from (a) GX, (b) DXP–GX, (c) CN–GX, (d) DXP–CN–GX, (e) BD and (f) DXPCN–GX-BD scaffolds; and collagen type 1 from (g) GX, (h) DXP–GX, (i) CN–GX, (j), DXP–CN–GX, (k) BD and (l) DXP–CN–GX–BD scaffolds; scale bar = 50 μm; resolution = ×40. Com# parative analysis of the percentage area covered with (B) osteocalcin and (C) collagen type 1; different to BD control (p < 0.05); *different to GX control (p < 0.05); n = 4

mixture of gellan and xanthan gum, whose gelation time depends upon the temperature of the solution and the concentration of the polymer and crosslinker. Earlier work by different authors claimed that the thermoresponsive nature of low-acyl gellan gum, whose setting temperature varies in the range 30–40°C, depended upon the ions and polymer concentration (Dyondi et al 2013; Oliveira et al 2010). At high temperature, gellan remains in the form of a randomly distributed chain, while at low temperature it forms a double helical structure through hydrogen bonding, leading to sol–gel transition (Chandrasekaran and Thailambal, 1990). Divalent cations (crosslinkers) further strengthen the hydrogel’s elasticity and stiffness through ionic crosslinking between the adjacent chains, whose strength varies according to the crosslinker concentration. In our system, a combination of the gellan and xanthan polymer along with the crosslinker CaCl2 and DXP–CN nanoparticles has been optimized for their effects on the gelation time and infiltration within the BD scaffold; 0.5% GX polymer with 2 mM CaCl2 has been used for the coating of the BD scaffold, because of its feasible gelation time (7–8 min) and high viscoelasticity at physiological temperatures (Figure 2A), which allows homogeneous filling and coating of the hydrogel within the scaffold. Copyright © 2015 John Wiley & Sons, Ltd.

Although hydrogels are excellent candidates in terms of biocompatibility and 3D cell encapsulation, a major problem remains, i.e. that the hydrogel-based scaffolds do not allow cells to adequately adhere and spread, due to the lack of stiffness of the soft polymeric hydrogel (Geckil et al., 2010). Osteoblasts attach, spread and are biologically more active on stiff surfaces compared to soft surfaces (Discher et al., 2005). We obtained a 2.4-fold increment in the compressive modulus of the BD scaffold after hydrogel coating compared to the uncoated one, and a 40-fold increment compared to hydrogels, suggesting that the hydrogel coating over the BD scaffold will provide a stiff surface for cell growth and will improve the performance of ceramic BD or pure hydrogel scaffolds. We previously demonstrated the bioactivity of the BD scaffold for bone tissue engineering (Ramaswamy et al., 2008). Deposition of apatite on the DXP–CN–GX–BD scaffold follows typical bio-glass scaffold behaviour, as discussed by Ohtsuki et al. (1992). In our study, the deposition of apatite can be attributed to initial calcium secretion from the DXP–CN–GX–BD scaffold, the formation of a supersaturated solution, initiation of apatite precipitation and, finally, nucleation over the negatively charged COO– hydrogel surface. Release of DXP from DXP–CN nanoparticles was more sustained compared to J Tissue Eng Regen Med (2015) DOI: 10.1002/term

Gellan–xanthan–baghdadite scaffold augments bone regeneration

DXP–GX hydrogels, due to charge interactions between the positively charged chitosan and the negatively charged DXP. Coating of the GX hydrogel over the DXP–CN nanoparticles further provided additional diffusion barriers for the controlled release of the drug. The cumulative concentration of DXP released from the DXP–CN–GX–BD scaffold was 1.22 ± 0.22 μg/ml after 1 day and 2.11 ± 0.28 μg/ml after 2 days in PBS medium (Figure 5A). Keeping in mind the DXP toxicity levels, low doses of DXP–CN nanoparticles (200× dilution, with DXP concentration 6.14 ng/ml after 1 day and 10.5 ng/ml after 2 days) was used for all cell culture experiments, which is within the non-toxic range for DXP. Enhanced cell proliferation was evident inside CN–GX hydrogels compared to GX hydrogels, while no significant difference in proliferation was observed inside DXP–CN– GX hydrogels compared to GX hydrogels at day 14 of culture. Enhanced proliferation in the CN–GX hydrogel confirmed the role of chitosan in enhancing cell proliferation, as previously suggested by others (Oliveira et al., 2006; Custodio et al., 2010). Moreover, because of the sustained release of DXP from DXP–CN–GX hydrogels, which can work as a switch for cell differentiation rather than proliferation, we obtained no difference in the proliferation rate of cells within DXP–CN–GX hydrogels compared to GX hydrogels. A similar kind of behaviour has been reported previously (Shi et al., 2010), where DEXAloaded scaffolds showed less cell growth compared to control unloaded scaffolds. The enhanced growth of cells in DXP–CN–GX–BD scaffolds compared to BD scaffolds may have been due to the biomimetic environment provided by the nanoparticulated hydrogel-coated scaffold (Rajam et al., 2011; Levengood and Zhang, 2014; Guzmán-Morales et al., 2009). ALP activity and calcium deposition are important features representing the earlier and later stages of cell differentiation. The incorporation of nanosized particles (DXP–CN) into hydrogels (DXP–CN–GX) resulted in enhanced ALP activity and higher calcium deposition of MG 63 cells compared to GX and DXP–GX hydrogels, suggesting that DXP released from the nanoparticulated hydrogel was biologically more active compared to direct loading within the hydrogel. Compared to previous studies that delivered high doses (20–200 μg/scaffold) of DEXA (Shi et al., 2010; Kim et al., 2003; Nuttelman et al., 2006), in the present study we demonstrated an improvement in cell differentiation at a low dose of the hydrophilic form of DXP (0.16 μg/scaffold). The bioactivity of BD scaffolds was further enhanced when the scaffolds were coated with the nanostructured hydrogel (DXP– CN–GX) inside the porous matrix, as reflected in enhanced ALP activity (two-fold higher) and higher calcium deposition (two-fold higher) at 7 and 21 days of culturing MG63 cells on DXP–CN–GX hydrogel-coated BD scaffolds, compared to BD scaffolds. In one earlier approach for the delivery of MSCs from the hydroxyapatite (HAP) scaffold, MSC differentiation was promoted ex vivo in culture medium supplemented with DEXA–dendrimer nanoparticles (rather than incorporation of nanoparticles within HAP Copyright © 2015 John Wiley & Sons, Ltd.

scaffolds) before their translation in the animal model (Oliveira et al., 2009). In our approach, co-encapsulation of cells and DXP–CN nanoparticles within the biocompatible gellan and xanthan hydrogel, and its coating over the BD scaffold, provide the feasibility of in situ differentiation of cells rather than their ex vivo culture with DEXAcontaining nanoparticles. The effect of DEXA on collagen type 1 secretion reported in the literature has shown variable findings, with both increase and decrease in expression reported in different studies (Jabbarzadeh et al., 2007; Beresford et al., 1994). However, the increase in osteocalcin was strongly associated with late osteoblastic differentiation. In this study, the immunohistochemistry results demonstrated higher expression of both collagen type 1 and osteocalcin on DXP–CN–GX–BD scaffolds compared to all controls, suggesting that the controlled release of DXP from these scaffolds enhanced cell differentiation. This study has demonstrated the efficacy of the DXP– CN–GX nanoparticulated hydrogel and BD scaffold in promoting the proliferation and differentiation of MG 63 cells, suggesting that the intracellular controlled delivery of DXP through biocompatible chitosan nanoparticles within the nanostructured hydrogel may be a promising strategy for inducing osteo-differentiation.

4. Conclusions This study has demonstrated that the coating of baghdadite scaffolds with a nanostructured hydrogel of gellan and xanthan, containing DXP-encapsulating chitosan nanoparticles, provided sustained DXP release and improved the bioactivity of the scaffolds. Such scaffolds, which cause sustained dexamethasone delivery, appear promising for bone regeneration. In future, the ex vivo and in vivo differentiation potential of MSCs embedded inside 3D networks of a hydrogel-coated baghdadite scaffold should be evaluated.

Conflict of interest The authors declare no conflicts of interest.

Acknowledgements The authors would like to acknowledge the Centre for Research in Nanotechnology and Science (Indian Institute of Technology Bombay), the Industrial Research and Consultancy Centre (Indian Institute of Technology Bombay), the Indian Institute of Technology (Bombay), the National Health and Medical Research Council (Australia), the Australian Research Council (Australia) and the Rebecca Cooper Foundation (Australia) for providing instruments and cell culture facilities. R.R.S. acknowledges CSIR (India) for the award of a Junior Research Fellowship. The authors acknowledge Associate Professor Colin Dunstan for proofreading the manuscript. No kind of ethical approval was required for this study. J Tissue Eng Regen Med (2015) DOI: 10.1002/term

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Supporting information The following supporting information may be found in the online version of this article: Figure S1. Percentage viability of MG 63 cells treated with different concentrations of BD scaffold extract compared to cell growth on the control tissue culture plate

Copyright © 2015 John Wiley & Sons, Ltd.

J Tissue Eng Regen Med (2015) DOI: 10.1002/term

Gellan–xanthan–baghdadite scaffold augments bone regeneration

Figure S2. Percentage viability of MG 63 cells treated with different concentrations of DXP–CN–GX–BD scaffold extract compared to cell growth on the control tissue culture plate Figure S3. Comparison of storage and loss moduli of DXP–CN–GX and GX hydrogels with and without CaCl2 Figure S4. FTIR spectra of chitosan, TPP powder, blank chitosan nanoparticles, DXP and DXP-loaded chitosan nanoparticles (DXP–CN) Figure S5. FTIR spectra of gellan gum, xanthan gum and 0.5% gellan:xanthan (9:1) hydrogel (GX) crosslinked with 2 mM CaCl2 Figure S6. FTIR spectra of 0.5% GX (9:1) hydrogel crosslinked with 2 mM CaCl2, DXP–CN nanoparticles and DXP–CN nanoparticles loaded 0.5% GX (9:1) hydrogel crosslinked with 2 mM CaCl2

Copyright © 2015 John Wiley & Sons, Ltd.

J Tissue Eng Regen Med (2015) DOI: 10.1002/term