Development of novel 3D scaffolds using BioExtruder by the incorporation of silica into polycaprolactone matrix for bone tissue engineering

Materials Today Communications 21 (2019) 100651

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Development of novel 3D scaffolds using BioExtruder by the incorporation of silica into polycaprolactone matrix for bone tissue engineering

T

Nandini A. Pattanashettia, Sara Biscaiab, Carla Mourab, Geoffrey R. Mitchellb, ⁎ Mahadevappa Y. Kariduraganavara, a b

Department of Chemistry, Karnatak University, Dharwad, 580003, India Centre for Rapid & Sustainable Product Development, Polytechnic Institute of Leiria, Marinha Grande, 2430-028, Portugal

A R T I C LE I N FO

A B S T R A C T

Keywords: Three-dimensional scaffolds BioExtruder Morphology Wettability Cytotoxicity Bone tissue engineering

Development of three-dimensional (3D) scaffolds has acquired a great importance for bone repair and tissue reconstruction. Thus, this paper addresses the development of 3D scaffolds by varying the content of silicon dioxide (SiO2) in polycaprolactone (PCL) matrix. The scaffolds were fabricated by employing a novel fused deposition modelling technique (BioExtruder). The physicochemical properties of the developed 3D scaffolds were systematically studied using various techniques. The thermal properties and stability of the PCL and its composites were assessed using differential scanning calorimetry and thermogravimetric analysis. The morphology of the developed scaffolds was evaluated using scanning electron microscopy and optical microscopy, and found that the pore size was increased from 270 to 320 μm with increasing the SiO2 content in the PCL matrix. The wettability of the developed scaffolds was assessed using contact angle meter. The scaffold incorporating 15 wt% of SiO2 exhibited the highest hydrophilic property as well as thermal stability. The Young’s modulus value determined using universal testing machine indicated that the scaffold developed with 15 wt% of SiO2 exhibits 101.59 MPa. To assess the performance of the scaffolds for tissue engineering applications, the invitro cytotoxicity and cell proliferation were systematically carried out using L929 Mouse Fibroblasts and MG63 Osteoblasts, respectively. It was found that the scaffolds did not show any toxic effects towards the cell growth, and the cell proliferation was greatly increased > 90% during 7 days of cell culture. Based on the results, it is concluded that the scaffold containing 15 wt% of SiO2 is of potential candidate for bone tissue engineering application.

1. Introduction Bone repair and tissue reconstruction are the challenging topics in the fields of tissue engineering and regenerative medicine. Though bone is known for its self-healing property, the large scale bone defects, such as bone infections, trauma and bone tumours, hinder the healing process [1–3]. Thus, external intervention or bone substitute is often required to repair or replace the defect tissues. As an alternative option for conventional autografts and allografts, bone tissue engineering has been proved as more advantageous and effective in terms of bone repair and reconstruction [4,5]. To address this, a biodegradable and biocompatible scaffold serves as a temporary skeletal frame which can mimic the properties of extracellular matrix (ECM), such as mechanical support, cell adhesion, proliferation and differentiation [5–9]. These biodegradable and biocompatible scaffolds induce bone tissue regeneration and undergo gradual degradation, and thereby replacing a

⁎

new bone tissue [5]. In particular, scaffolds for bone regeneration should provide a highly interconnected porous structure. These pores allow the transportation of nutrients, light and oxygen molecules to the inner parts of a scaffold and facilitate the cell growth, vascularisation and removal of waste material, and thereby maintaining a sufficient mechanical strength to provide the structural requirements of the substituted tissue [5,9,10]. Such porous bone scaffolds can be developed by various techniques, like solvent casting/particulate leaching [11,12], gas foaming [13], freeze drying [14], thermally induced phase separation [7], electrospinning [9,15] etc. However, the scaffolds with required interconnected porous structure for specific defects are difficult to manufacture with these techniques [5,7,12–14]. In recent years, the manufacturing of three dimensional scaffolds increased the hope for the production of scaffolds with closer similarities to bone matrix [16]. Among the manufacturing processes, 3D printing is a unique and rapid prototyping (RP) process which was developed in early 1990′s at

Corresponding author. E-mail address: [email protected] (M.Y. Kariduraganavar).

https://doi.org/10.1016/j.mtcomm.2019.100651 Received 27 May 2019; Received in revised form 1 September 2019; Accepted 17 September 2019 Available online 18 September 2019 2352-4928/ © 2019 Elsevier Ltd. All rights reserved.

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with definite porosity. Before the fabrication of scaffolds, the thermal properties and strength of PCL and its composites were studied using differential scanning calorimetry (DSC) and thermogravimetric analysis (TGA), respectively. All the scaffolds possessed uniform fibres with interconnected porous structure as evidenced by the images of optical and scanning electron microscopes. The wettability of the composites and the scaffolds was studied by water contact angle (WCA) measurement. Fourier transform infrared spectroscopy (FTIR) was used to characterize the intermolecular interaction between the components of the scaffolds. The crystallinity of the scaffolds was analysed by X-ray diffraction. The enhanced mechanical strength of the scaffolds was realized by universal testing machine. The cell viability and MTT assay was systematically performed to assess the cell viability and cell proliferation using L929 Mouse Fibroblasts and MG63 Osteoblasts, respectively. The results were discussed based on the structure, interaction and nature of the materials.

Massachusetts Institute of Technology (MIT), Cambridge by Sachs and coworkers [17]. From this technique, the scaffolds could be directly manufactured from a set of 3D data without the requirement of an additional mould. The resulting scaffolds exhibit interconnected porosity, designed shape and controlled chemistry [8,18]. In addition, 3D printing is an additive manufacturing (AM) technology, wherein successive layers of the material are laid down in different shapes [5,8]. A large number of additive processes include fused deposition modelling, stereolithography, selective laser sintering and some of the extrusion based technologies [8]. In the manufacturing of 3D scaffolds, selection of biomaterials plays a prime role for the repair and replacement of human tissues and organs. This is done based on several factors, like biocompatibility, biodegradability, surface properties and mechanical strength [8,9]. A large number of polymers are being experimented to manufacture tissue engineering scaffolds, which includes PCL, poly(ethylene glycol), ploy (lactic acid), poly(vinyl alcohol) and polyurethane [9]. Among these, PCL is widely used synthetic degradable polymer for bone tissue engineering. PCL is a non-toxic semicrystalline aliphatic polyester. It is a tissue compatible, biodegradable and bioresorbable polymer with slow hydrolytic degradation via bulk and surface erosion. It exhibits low melting point at around 60 °C and high thermal stability with decomposition temperature of 350 °C. Thus, it provides an unusual thermal and structural stability along with a good processability [19,20]. With such characteristics, PCL enables its melt processing to obtain 3D porous scaffolds with highly interconnected porous network [19]. Though PCL has the desired properties for 3D scaffolds, its high hydrophobicity and low degradability makes it less suitable for longterm applications [8,19–21]. Thus, the development of composite scaffolds is gaining considerable interest by combining the PCL with other materials like hydroxyapatite, β-tricalcium phosphate (β-TCP), etc. Calcium phosphate was considered as one of the promising material as a bone substitute in view of its high biocompatibility, bioactivity, low setting temperature and adequate stiffness, but it fails due to lack of macropores and low porosity which result in lower resorbtion rate and mechanically brittle [22]. As an alternative to this, an idea was employed to make use of a biologically relevant metal oxide as dopant. Literature reveals that the use of dopants like silicon, zinc, magnesium and strontium have the ability to enhance the biological responses [23–29]. Among the dopants, silicon has been noted as an important trace element which is predominantly absorbed by humans in the form of orthosilicic acid and is found in numerous tissues including bone, tendons, aorta, liver and kidney [30,31]. The deficiency of silica can lead to skull and peripheral bone deformities, poorly formed joints, defects in cartilage and collagen, and disruption of mineral balance in the femur and vertebrae [32–37]. It is also proved that it is an important element in osteogenesis which can stimulate the cellular activities, such as proliferation, differentiation and mineralization of cells [23–26]. Silica is widely recognised as an essential component of bioactive ceramics in skeletal therapies and also facilitate bone repair at the wounded site [27]. Silicon (up to 0.5%) is often localized in active growth areas and played an important role in bone formation and calcification. For instance, Karlsson et al. [28] proposed that silica chelates form an essential step in the formation and mineralization of hard tissues. Similarly, β-TCP scaffolds doped with SiO2/ZnO; increased the mechanical strength and facilitated the cellular proliferation [29]. Thus, silicon plays an important physiological role in bone formation and particularly, the nano ranged bioactive silica particles have large surface area which can form a tighter interface with the polymer matrix [19,32–35]. Understanding the pros and cons of aforesaid PCL material and the importance of silicon, we have incorporated SiO2 as a dopant into the PCL matrix. The PCL and its composites were subjected to BioExtruder for the development of 3D scaffolds. The amount of SiO2 was varied and optimized in the PCL matrix with a sole desire to obtain better scaffolds

2. Materials and methods 2.1. Materials Polycaprolactone (Mw = 50,000 Da) was purchased from Capa™ 6500 Perstorp Caprolactone, Cheshire, UK. Silicon dioxide nanopowder (5–15 nm) was purchased from Sigma Aldrich Chemie, GmbH, Germany. Dichloromethane (DCM) and dimethyl sulfoxide (DMSO) procured from Sigma Aldrich, USA were employed as solvents. All the chemicals were of reagent grade and used as received. L929 Mouse Fibroblasts were procured from European Collection of Authenticated Cell Cultures (ECACC), Sigma Aldrich, Inc. Germany for cell viability assays. Cell proliferative study was performed using MG-63 Osteoblasts, procured from American Type Culture Collection (ATCC, Manassas, USA). Trypsin phosphate versene glucose solution (TPVG) and phosphate buffered saline was purchased from HIMEDIA, Mumbai, India. An MTT assay was carried out using yellow tetrazolium MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide) purchased from Sigma Aldrich, USA. Dulbecco’s Modified Eagle’s Medium and Fetal Bovine Serum were respectively obtained from Gibco® and HyClone®, USA for the assays. 2.2. Preparation of PCL/SiO2 composite material The composites were prepared with different weight percent of PCL/SiO2. A 20 g of PCL was dissolved in 20 ml of dichloromethane by adding PCL in small portions in the solvent which was stirred continuisuly using a magnetic stirrer for 1 h at room temperature. After the complete dissolution of PCL, the silica nanoparticles were added slowly to the solution and stirred further for 1 h. The resulting mixture was a highly viscous fluid which was casted on a glass plate in the form of pellets and allowed to air dry overnight. The amount of SiO2 with respect to PCL was varied from 0, 5, 10 and 15 wt%. After the complete evaporation of the solvent, the dried pellets were cut and kept in a vacuum desiccator until further use. These dried pellets were used for the fabrication of 3D scaffolds using an extruder based fused deposition modelling technique developed at Centre for Rapid and Sustainable Product Development (CDRSP), Polytechnic Institute of Leiria, Portugal as shown in Fig. 1. This extrudes the molten thermoplastic material through a nozzle [19,34,35]. It is basically a solvent free technique in which polymer is allowed to melt directly above its melting point and extruded in the form of filaments with a width of approximately 300 μm [19,31,34,35]. The preparation scheme of scaffold using PCL and SiO2 is shown in Fig. 2. Based on the weight% of SiO2, the resulting scaffolds were designated as PCL, PS1, PS2, and PS3 as shown in Table 1. An attempt was also made to incorporate the higher weight% of SiO2, but it was not possible to successfully extrude the composite 2

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Fig. 1. The BioExtruder system employed for the fabrication of scaffolds.

material even after adjusting the process parameters.

Table 1 Fabrication of scaffolds with different weight percent of PCL and SiO2.

2.3. Fabrication of scaffolds A BioExtruder technique was used to fabricate the three-dimensional block models of the scaffolds using computer aided design (CAD) software (Solid Works, Dassault Systemes, S. A.). Layer by layer assembly with 0/90° lay down pattern was employed for the development of scaffolds to obtain highly interconnected porous network of square channels as shown in Fig. 3. A regular filament distance (FD) of 650 μm was maintained through a software. The nozzle tip with a diameter of 0.3 mm was used for extrusion. This maintains the quadrangular internal pore geometries with porosity of 350 μm. The pore size ranging between 300 and 350 μm is known to be optimum for bone tissue engineering and the same was maintained in the present study during the fabrication of scaffolds [8,29,31,38]. Through a series of trials, the scaffolds of desired porosity were fabricated by varying the temperature (above their melting temperature) at the thermocouples. The deposition velocity of 15 mm/s and screw rotation velocity of 50 rpm were maintained during the fabrication of scaffolds. The scaffolds of two different dimensions, measuring (length x width x height) 20 mm x 20 mm x 7.84 mm and 20 mm x 20 mm x 3.36 mm were employed for all the compositions.

Scaffold

PCL (wt%)

SiO2 (wt%)

Melt temperature in the extruder (°C)

PCL PS1 PS2 PS3

100 95 90 85

0 5 10 15

90 90 100 100

3.2. Thermal analysis of the composite material The composite material is to be subjected to various temperatures during the melt extrusion process. Generally, this temperature has to be slightly higher than the melting temperature of the composite material. It therefore becomes an important criterion to find out the exact melting temperature and the degradation temperature of the composite material prior to the fabrication of scaffolds. Thus, the following thermal analyses have been carried out. 3.2.1. Differential scanning calorimetry Thermal properties of PCL and its composites were investigated using Perkin-Elmer differential scanning calorimeter (SDT Q600, TA Instruments-Waters LLC, USA). The samples of around 8–10 mg were heated from ambient to 120 °C at a heating rate of 10 °C/min. 3.2.2. Thermogravimetric analysis Thermal stability of PCL and its composites were investigated using Perkin-Elmer TGA/DTA thermogravimetric analyzer (SDT Q600, TA Instruments-Waters LLC, USA). Around 8–10 mg of the samples was heated from ambient to 500 °C at a heating rate of 10 °C/min.

3. Characterization 3.1. Fourier transform infrared spectroscopy (FT-IR) To study the interaction between SiO2 and PCL, the Fourier transform infrared spectroscope (Nicolet, Impact-410, USA) was employed. The composite material and potassium bromide were mixed and ground well for the preparation of pellets and recorded the spectra in the range of 400-4000 cm−1.

3.3. Morphology and porosity of scaffolds The porosity and fibre diameter of the scaffolds were observed and measured using optical microscope (MICROS, CROCUS, Austria). The surface morphology of the scaffolds was assessed by scanning electron Fig. 2. Preparation scheme of scaffolds using PCL and SiO2.

3

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Fig. 3. An illustration of scaffold design: (a) the base layer and (b) succeeding layers with 0/90° lay down pattern; (c) 3D view of the scaffold and (d) frontal view of the developed porous scaffold.

was calculated using Eq. (1), by dividing the load F by the initial crosssectional area of the scaffold (A = lw) [37]:

microscope (SEM, FEI Quanta FEG 600, UK). Except PCL, other samples were examined without coating under low vacuum conditions. However, the PCL sample was vacuum dried and then sputtered with gold.

σ=

F A

(1)

The strain ԑ was evaluated using Eq. (2) by the ratio between the change in height (Δh) and initial height (ho) of the scaffold:

3.4. Wettability of the scaffolds To study the wettability of the scaffolds, the water contact angle was measured by a sessile drop method at room temperature using static contact angle goniometer (Optic Tensiometer, Theta Software: One attention 1.0). It was realized that the interconnected porous structure of the three-dimensional scaffolds would play a major role during the wettability measurement of the scaffolds. To overcome this difficulty, thin membranes of all the compositions were prepared by a solution casting method. The membranes thus obtained were subjected to goniometer for the measurement of contact angle.

ԑ=

Δh ho

(2)

The Young’s modulus (ԑ) of each sample was calculated from the slope of the initial linear portion of the average stress-strain curve. 3.7. Cytotoxicity of the scaffolds The in-vitro cell viability of the scaffolds was carried out using L929 Mouse Fibroblasts according to ISO standard (10993-5:2009). The samples (5 mm х 5 mm х 2 mm) were stabilized overnight in 70% ethanol under UV light and then washed with phosphate buffer saline (PBS). L929 Mouse Fibroblasts were cultured in Dulbecco’s Modified Eagle’s Medium (DMEM), supplemented with 10% Fetal Bovine Serum (FBS). The cells were then seeded on 24-well plates in order to obtain approximately 80% of confluence. The approximate population/cell count was 1.5 х 105 cells/well. The plates were incubated overnight at 37 °C under 21% O2 and 5% CO2 atmosphere. Two assays namely, direct contact (qualitative) assay and extract (quantitative) assay were performed on all the scaffolds. For the extract assay, the scaffolds were submerged in the DMEM medium and left in the incubator to obtain the leachings. After 12 h of incubation, the medium of the cells was then discarded and the leachings were further used to submerge the cells. For the direct contact assay, the sterilised scaffolds were directly placed on the top of the cells in the well plate. Both the plates were left in the incubator for another 24 h. In case of direct contact assay, the viability of the cells towards scaffolds (Live or Dead) was observed by taking the photographs through optical microscope. However, an MTT protocol (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide) was performed for the leachings of the extract assay. This test can detect the conversion of 3-(4,5-dimethylthiazolyl-2-yl)-2,5-diphenyltetrazolium

3.5. Powder X-ray diffraction analysis (PXRD) The X-ray diffraction analysis was performed by powder X-ray diffractometer (Bruker D8 Advances, UK) at room temperature. The X-ray source was Ni-filtered Cu-Kα radiation (40 kV, 30 mA). The samples were mounted on a sample holder and diffraction was scanned in the reflection mode at an angle of 2θ over a range of 10°-60° at a constant speed of 5°/min. 3.6. Mechanical testing of the scaffolds The mechanical testing of the 3D BioExtruded scaffolds was carried out in order to obtain the compression moduli of the porous scaffolds at room temperature using the testing machine (Zwick Roell Z100, USA) with a load of 10 K N and 1 mm/min of compression velocity. The scaffolds were cut to obtain the specimens in a rectangular shape of dimensions 4 mm х 4 mm х 7.84 mm. The load was applied upon the specimens and the scaffold was compressed until the strain limit of the machine was reached. Ten specimens for each composition were tested and stress-strain curves were obtained. For each sample, the stress (σ) 4

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bromide to formazan. The fresh medium was used as a negative control while the latex was used as a positive control. In brief, 100 μl of MTT (5 mg/10 ml of MTT in PBS) was added to each well containing the leachings of the extract assay. The plates were gently shaken and incubated for 4 h at 37 °C in 5% CO2 atmosphere. The mitochondrial dehydrogenase enzyme of viable cells cleaves the tetrazolium ring to an insoluble purple formazan. The supernatant was removed and 100 μl of DMSO was added and the plates were gently shaken to solubilize the intracellularly formed formazan. Later, the scaffolds were removed and the absorbance was recorded using a microplate reader at a wavelength of 540 nm. The cell proliferation was assessed by the generation of purple colour, which is directly proportional to the number of metabolically active proliferated cells. The cells proliferative potential on each scaffold was calculated using the formula: Viable/proliferative (%) = (Absorbance of Control – Absorbance of Sample) /(Absorbance of Control) х 100 3.8. Cell proliferation in the scaffolds To study the cell proliferation, each scaffold (3 mm х 3 mm х 2 mm) was sterilised with 70% ethanol for 24 h and rinsed twice with a plain DMEM medium and pre-conditioned. All these steps were performed under sterile condition. The MG-63 osteoblasts cell line procured from ATCC was cultured in DMEM supplemented with 10% inactivated FBS, 1% penicillin (100 IU/ml) and streptomycin (100 μg/ml) under a humidified atmosphere of 5% CO2 at 37 °C until confluent. The cells were dissociated with TPVG solution containing 0.2% trypsin, 0.02% EDTA and 0.05% glucose in PBS. The viability of the cells was checked using trypan blue and further centrifuged. Further, the MG-63 cells were seeded in a 96-well plate at a rate of 50,000 cells/well and incubated for about 24 h at 37 °C under 5% CO2 incubation. The scaffolds (3 mm х 3 mm х 2 mm) were placed in each well of the 96 well plate. The MG-63 monolayer cells were trypsinized and the cell count was adjusted to 5.0 х 105 cells/ml using DMEM containing 10% FBS. To each well of the 96 well microtiter plate, 100 μl of the diluted cell suspension (50,000 cells/well) was seeded on each scaffold. The cells seeded on cell culture plate without scaffold served as the control group. The plates were then incubated at 37 °C for 3 days in 5% CO2 atmosphere. Further microscopic examination was carried out and recorded the observations for every 24 h and noticed the cell attachment and proliferation. After 72 h, the test solutions in the wells were discarded and an MTT assay was performed as similar to the procedure adopted in extract assay of the cytotoxicity study. The same procedure was repeated for the 5th and 7th day, and calculated the percent of cell proliferation based on the absorbances.

Fig. 4. The FTIR spectra of PCL and its SiO2 composites.

C–O bands of PCL. 4.2. Thermal analysis Among the different parameters varied during the fabrication of scaffolds, the temperature fixed at three different zones also played an important role for uniform and easy extrusion of the composite material. To select and set the appropriate temperature at three zones, the thermal properties of PCL and its composite materials were investigated by DSC and resulting thermograms are presented in Fig. 5. The melting temperature of pure PCL was found to be 60.91 °C and this was gradually shifted to higher temperature with the incorporation of SiO2. It was particularly shifted to 61.18, 62.1 and 62.2 °C for 5, 10 and 15 wt% SiO2, respectively. This was mainly attributed to the incorporation of silica and its weight percent in the PCL matrix. The incorporated nanosized silica established strong interaction and thereby increased the thermal stability of the resulting material by way of acting as a barrier for mass transport to the volatile products generated during melting process. Similarly, the thermal stability of PCL and its composites was

3.9. Statistical analysis Results were averaged and expressed as mean ± standard deviation. The differences between the scaffolds and the treatment days in cell proliferation were statistically evaluated using ANOVA. A value of P < 0.05 was considered statistically significant. 4. Results and discussion 4.1. FT-IR analysis The FTIR spectra of pure PCL and silica doped samples were recorded and the resulting spectra are presented in Fig. 4. In PCL spectrum, the bands appeared at 2857 and 2925 cm−1 were attributed to symmetric and asymmetric stretching vibrations of CH2 groups, respectively. The band appeared at 1732 cm-1 was assigned to the carbonyl stretching (C=O) of the ester linkages. Upon comparing the spectrum of PCL with the spectra of PS1, PS2 and PS3 samples, it can be seen that there was no significant difference in the bands. This is expected, since Si-O bands of SiO2 also appear at the same frequency of

Fig. 5. The DSC thermograms of PCL and its SiO2 incorporated composites. 5

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4.3. Fabrication of the scaffolds In order to mimic the bone structures, three-dimensional scaffolds were fabricated using BioExtruder. The scaffolds were developed by single lay-down pattern of 0/90° with a constant filament distance (FD) of 650 μm to obtain square interconnected pores of regular dimensions. To achieve smooth and uniform filaments of the scaffolds with required porosity, two major parameters, such as temperature and voltage, were varied during the fabrication process. Based on the results obtained by DSC and TGA thermograms, the temperatures were set between 70 and 100 °C during the fabrication of scaffolds. By varying these two parameters, the scaffolds with required porosities, i.e., in the range of 270–320 μm were obtained. An utmost care was taken to achieve the required range of pore sizes. 4.4. Morphology of the scaffolds Fig. 6. The TGA patterns of PCL and its SiO2 incorporated composites.

The variables of the scaffolds, such as pore size and porosity, are the key parameters to determine the tissue engineered constructs. These parameters particularly to promote the bone formation play a vital role in cell attachment, cell migration and cell infiltration, and thus have a direct impact on vascularisation and osteoconduction. Interconnected porous structure allows nutrients molecules to transport inside the scaffold which in turn facilitates the cell ingrowth. The rate of degradation and mechanical strength of the scaffolds are also directly dependent on the porosity of the scaffolds, because higher porosity always increases the surface area per unit volume of the scaffolds. According to the literature, the pore size between 100 and 150 μm is minimum requirement for bone formation. However, the pore size larger than 350 μm is reported to be a good range for the enhanced bone formation and vascularisation [5,9]. The morphological evaluation of the developed three-dimensional scaffolds was carried out using optical microscope and scanning electron microscope. The optical microscopic images of all the developed scaffolds are presented in Fig. 7. The scaffolds developed from BioExtruder presented a well defined

evaluated by TGA and the thermograms are presented in Fig. 6. It is observed that all the samples underwent degradation in a single step and no loss occurred even up to 300 °C. The PCL underwent degradation at around 325 °C. However, it was shifted to 370 °C after the addition of 5 wt% of SiO2 and this was further increased to 425 and 450 °C for 10 and 15 wt% SiO2, respectively. This suggests that the addition of silica played an important role in enhancing the degradation behaviour as well as thermal stability of the composites. This apparently due to a better adhesion between silica nanopowder and the polymer matrix, and this hinders the diffusion of volatile molecules across the matrix during the melting process. Based on these results, suitable temperature was chosen and maintained at three different zones of the BioExtruder. However, the zones were set slightly above the melting temperatures of the samples during the fabrication of the scaffolds. Based on the thermal stability data, we conclude that the developed composites could be of suitable candidates for regeneration of bones.

Fig. 7. The optical microscopic images of the developed scaffolds. 6

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could able to achieve which plays an important role during the cell growth and proliferation for bone tissue engineering.

4.5. Wettability of the scaffolds The wettability of the samples generally depends on the interconnected porous structure of the scaffolds. But, the difference in pore sizes of the 3D structured samples could interfere in the results of wettability measurements. In order to know the pore size and accurate hydrophobic/hydrophilic nature of the developed PCL and its composite materials, the water contact angle (WCA) was determined on the porous three-dimensional scaffold structures (Fig. 10a) and also upon the composite films (Fig. 10b). It can be seen from the Fig. 10a that the hydrophobic nature of pure PCL was reduced by the incorporation of SiO2 and its wt% in the PCL matrix. The pure PCL scaffold exhibited WCA of 83.49° with hydrophobic nature, whereas the inclusion of silica resulted in a decreased WCA of 82.38, 82.22 and 79.13° for PS1, PS2 and PS3, respectively. Apart from the characteristic properties of materials, the interconnected porous channels also play a major role in determining the wettability of 3D scaffolds. This is clearly evidenced from the images of both optical microscopic and SEM. In order to understand the role of only the characteristic properties of material on the contact angle of water, we developed membranes of pure PCL and its composites. The WCA for pure PCL was found to be 82.15°. For PS1, PS2 and PS3 composite membranes, it was found to be 70.12, 64.17 and 35.47°, respectively. The addition of silica introduces the functional groups in the PCL matrix, which could form the hydrogen bonding with the water molecules and thus show a hydrophilic nature. The formation of hydrogen bonding becomes prominent with the increase of silica content in the PCL matrix and thereby demonstrated an enhanced hydrophilic nature. On the whole, it was concluded that the incorporation of SiO2 into the PCL matrix has greatly reduced the

Fig. 8. Variation of pore size and fibre width with respect to SiO2 in PCL scaffolds.

geometry with uniform interconnected porous structure as evidenced from the microscopic images. It is observed that the extruded filaments possessed a regular geometry with decreasing fibre width from 310 to 270 μm. Also, the filaments appeared to have a good adhesion with adjacent filaments. It was further observed that with increase in silica content, the extruded filaments became thinner and smooth, and thus reducing the filament width. Thus, the porosity of the scaffolds was increased as clearly noticed from Fig. 8. The similar observations were made from the SEM images of the scaffolds presented in Fig. 9. The filaments were observed to be thinner and smoother with the increase of SiO2 content and thereby increased the pore size of the scaffolds. Thus, the overall porosity of the scaffolds was found to be in the range of 270–320 μm, wherein the porosity increased from PCL to PS3. On the whole, an optimum range of pore size

Fig. 9. The SEM images of PCL, PS1, PS2 and PS3 scaffolds. 7

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Fig. 11. The PXRD patterns of PCL, PS1, PS2 and PS3 scaffolds. Fig. 10. (a) Contact angle of PCL, PS1, PS2 and PS3 scaffolds. (b) Contact angle of PCL, PS1, PS2 and PS3 membranes.

normal biomechanical function. Thus, the mechanical properties of the scaffolds must match the physical demand of the surrounding bone. Further, the mechanical properties of human bone vary depending on the location and function. Therefore, the mechanical properties of the scaffolds should be tailored to match the requirement of the defect site. To address this, the polymer-metal oxide composite scaffolds were developed to fulfil the requirement of mechanical properties for bone tissue engineering. The strength/stiffness of the scaffolds was determined by mechanical compression testing. The mean stress-strain curves of scaffolds were plotted as shown in Fig. 12. A typical stress-strain curve comprises of 3 distinct regions, such as

hydrophobicity of the pure PCL scaffolds and thus helps to improve the cell adhesive property of the scaffolds. In a similar study by Ganesh et al. [39], PCL/SiO2 nanofibrous composite scaffolds fabricated by electrospinning technique demonstrated a higher contact angle of 118° and 114° respectively for 0.5% and 1% SiO2 incorporation, and thereby exhibiting a hydrophobic nature of the scaffolds. Thus, due to the three dimensional structure of the developed scaffolds, the scaffolds in the present study are hydrophilic in nature which plays an important role for adhesion of cells and their proliferation. 4.6. PXRD analysis Fig. 11 shows the XRD patterns of pure PCL and its SiO2 incorporated scaffolds. Two strong and sharp diffraction peaks appeared at 2θ = 21.4 and 23.8° in all the samples, which were respectively attributed to the diffraction of (110) and (200) crystallographic planes of semi-crystalline PCL. However, the intensity of these two peaks was significantly lowered with increasing the content of SiO2 in PCL matrix. This indicates that the crystallinity of the samples was decreased with increasing the silica content in the scaffolds. This apparently favours the cell proliferation owing to the predominance of the amorphous domains in the scaffolds. 4.7. Mechanical strength of the scaffolds For bone tissue engineering, the characteristic feature of the scaffolds plays an important role to provide the mechanical integrity at the defect site till the bone tissue is completely repaired and restored the

Fig. 12. The stress-strain curves of the scaffolds. 8

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a linear elastic region, a plastic region of roughly consant stress and a region of steeply rising stress. From the curves, it was observed that pure PCL scaffolds exhibited higher strain% with lower stress values, whereas the incoroporation of silica had a decreasing effect on the starin at break values. On the other hand, the silica doped scaffolds exhibited higher stress values as compared to pure PCL. This clearly reveals that though pure PCL is a tough material, but the strength and toughness of the composite scaffolds increased with increasing the silica content in the PCL matrix. The Young’s modulus values were also calculated from the slopes of the linear portions of the stress-strain curves and are presented graphically as shown in Fig. 13. It is again observed that the incorporation of silica has greatly increased the Young’s modulus values of the scaffolds. The pure PCL exhibited a comparatively lower Young’s modulus value of 54.13 MPa, while PS1, PS2 and PS3 scaffolds showed the improved values of 78.55, 86.72 and 101.59 MPa, respectively. These results revealed that the 3D printed

Fig. 13. The Young’s Modulus values of the composite scaffolds.

Fig. 14. The optical microscopic images of the cell viability on the scaffolds and the controls. 9

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scaffolds with the incorporation of SiO2 into PCL matrix has greatly enhannced the mechanical strength of the resulting scaffolds, and thus proved its potentiality for bone tissue regeneration. In a similar study by Castro et al. [40] the mechanical properties of the electrospun scaffolds were greatly improved by the incorporation of silica nanoparticles up to 33%, whereas with higher amounts of silica in the scaffolds the tensile modulus was decreased. The incorporation of silica does not necessariliy enhance the mechanical propeties of the scaffolds in every case. But, it all depends on the supporting polymer and the scaffold fabrication technique. For instance, Sowjanya et al. [41], observed a decrease in Young’s modulus value (8.16 MPa) of the scaffold fabricated by freeze drying method due to SiO2 incoproration into chitosan/alginate biocomposite matrix. The effect of some other bioactive inorganic materials was also studied by Gao et al. [42], wherein 0.5% SiO2 and 1% MgO was incorporated into β-TCP and the scaffolds were fabricated by selective laser sintering method. The compressive strength of the resulting scaffolds were 5.74 MPa and 902 MPa, respectively for SiO2 and MgO incorporated scaffolds. A maximum cmpressive strength of 10.43 MPa was obtained for the (0.5% SiO2 + 1% MgO + β-TCP) composite scaffold. But, in the present study, 15 wt% of SiO2 incorporation has resulted into a maximum Young’s modulus of 111.83 MPa. Thus, we can conclude that the BioExtruder is a supportive technique for the incorporation of higher content of inorganic materials into the polymer matrix, which further resuls into an improved mechanical properties.

Fig. 16. The cell proliferation of MG-63 osteoblasts on the scaffolds.

show any toxic effects on the growth of fibroblast cells. Thus, it can be further stated that neither of the composite material nor the use of solvent has any toxic effect on the cellular growth and cell viability. In order to study the cell proliferation of MG-63 osteoblasts on the developed scaffolds, an MTT assay was performed by culturing the cells and observations were recorded on 3rd, 5th and 7th day, the resulting data are presented in Fig. 16. From the data, it is clear that all the scaffolds exhibited cell proliferation in the range of 88–90% even after 3 days of culture. Further, the fluorescence values of all the scaffolds were increased with increasing the duration of cell culture from 5 to 7 days. At 5th day, the cell proliferation of PS3 scaffold was significantly higher than those of PS1 and PS2. However, the cell proliferation of pure PCL is almost on par with PS3. A similar trend was observed even at the 5th and 7th day of cell culture. Among the scaffolds studied, pure PCL showed better cell viability (99%) and cell proliferation (90–97%) as compared to other silica incorporated scaffolds. This reveals that though there was no toxic effect in any of the cases, but the incorporation of SiO2 has little inhibitory effect for cell viability and proliferation. When a comparative study was made among the SiO2 incorporated scaffolds, the PS3 demonstrated the highest cell viability and proliferation. This can be attributed to the fact that PS3 has extreme interconnected porous structure (∼316 μm) and lower water contact angle, and thereby supported the cells to penetrate and grow freely into the scaffold. Consequently, the cell proliferation was enhanced. This is expected, since the pore size and wettability of the scaffolds play an important role for cellular penetration and migration within the scaffolds [31,34]. The effect of wettability on the cell viability of the scaffolds was also observed in a study of nanosilica incorporated PCL composite nanofibrous scaffolds, wherein, the cell viability decreased to 84% due to higher contact angle of 114° of PCL/ SiO2(1%) composite [39]. Thus, in the present study, the 3D interconnected porous structure of PS3 scaffold demonstrated the highest cell proliferation and could be of potential candidate for bone regeneration. Literature reveals a number of studies wherein bioactive inorganic materials are embedded into the polymer matrix in order to enhance the properties of the scaffolds to be suitable for bone tissue engineering. But, achieving both the properties of mechanical stability and cell proliferation simultaneously is a matter of difficulty. For instance, Lee et al. [43], fabricated the multilayer PCL/silica composite scaffolds using three-axis robot system, and the developed scaffolds showed an increased cell proliferation with increase in SiO2 content while compromising the tensile strength value up to 3.2 MPa. In another study of chitosan/alginate/SiO2 biocomposite scaffolds, the SiO2 incorporation into chitosan/alginate has resulted to ≥100% cell proliferation with MG-63 rat osteoblasts, while the Young’s modulus decreased from 8.99 MPa to 8.16 MPa [41]. However, in the present study, both the properties which are necessary for the bone tissue engineering are achieved by judiciously incorporating the SiO2 into PCL matrix.

4.8. Cell viability and cell proliferation The interaction between cell and the scaffold acts as a minimum requirement for cell attachment and cell proliferation into the scaffolds. The scaffolds that are intended to be used for tissue engineering application should not release any toxic agents. In order to know this, L929 Mouse Fibroblasts were cultured onto the developed scaffolds to evaluate their potential cytotoxicity. Similarly, in-vitro cell culture study was conducted to investigate the biocompatibility of the scaffolds in terms of cell viability and cell proliferation. The obtained microscopic images are presented in Fig. 14. From the images, it is noticed that the seeded cells were uniformly distributed and adhered to the walls of the scaffolds. Further, it is observed that the PCL and its silica doped scaffolds (PS1, PS2 and PS3) did not exhibit any significant differences in the cell viability as compared to the cell growth on the negative control. Similarly, the percentage of cell viability was calculated for all the scaffolds as well as for the controls, and the resulting data are presented in Fig. 15. It is observed that the cell viability of PCL scaffold and negative control are almost same. However, a marginal decrease in the cell viability was noticed in the SiO2 incorporated scaffolds, but this is insignificant. This clearly suggests that the developed scaffolds did not

Fig. 15. The cell viability of the scaffolds and controls. 10

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5. Conclusions

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In the present study, 3D polymeric scaffolds were fabricated using a novel fused deposition modelling technique. The PCL was employed as a main source of scaffold material and the content of SiO2 was varied and optimized suitably in the PCL matrix. The solvent method was particularly employed to enhance the homogeneity between PCL and SiO2. The incorporation of SiO2 and its content in the PCL matrix have greatly enhanced the thermal stability of the scaffolds as evidenced from TGA and DSC data. All the developed scaffolds exhibited well adhered uniform fibres and highly interconnected porous structure as noticed from the images of optical and scanning electron microscopes. The silica incorporated scaffolds have produced smooth and thin fibres and thereby enhanced the pore size of the scaffolds up to ∼320 μm. The increased interconnected porous channels resulted to increased hydrophilic nature of the scaffolds as indicated by the water contact angles. Thus, the 3D structure of PS3 scaffold exhibited the lowest water contact angle of 79.13°. This was explained based on the creation of additional functional groups by the incorporation of silica in the PCL matrix and the subsequent formation of hydrogen bonding with the water molecules. The nano ranged silica particles have larger surface area and they formed a tighter interface with the PCL matrix. This resulted into a higher stress value of silica doped scaffolds as observed by the stress-strain curves of the scaffolds. Thus, the mechanical strength of the developed 3D scaffolds was increased with increasing the silica content in the scaffolds and obtained a maximum Young’s modulus value of 101.59 MPa for PS3 scaffold. Further, the developed scaffolds were proved to be viable and non-toxic to the cells as observed by the cell viability assay of L929 Mouse Fibroblasts. The incorporation of osteogenic silica content into PCL matrix and the resulting well interconnected porous structure of the scaffolds demonstrated a remarkable increase of cell proliferation as evidenced by culture study. On the whole, the PS3 scaffold made of 15 wt% silica and 75 wt% PCL demonstrated the highest cell proliferation as well as mechanical strength and thereby achieved the important requirements for bone tissue engineering. Thus, the developed PS3 scaffold proved to be a potential candidate for bone regeneration. Declaration of Competing Interest None. Acknowledgements Authors wish to thank the DST, New Delhi, for extending financial support under DST-PURSE-Phase-II Program [Grant No. SR/PURSE PHASE-2/13(G)]. Authors remain grateful to Dr. Saeed Mohan, University of Reading, UK, for helping in recording the microscopic images. References [1] V. Mourin˜o, A.R. Boccaccini, Bone tissue engineering therapeutics: controlled drug delivery in three-dimensional scaffolds, J. R. Soc. Interface 7 (2010) 209–227. [2] H. Seitz, W. Rieder, S. Irsen, B. Leukers, C. Tille, Three-dimensional printing of porous ceramic scaffolds for bone tissue engineering, J. Biomed. Mater. Res. Part B Appl. Biomater. 74B (2005) 782–788. [3] A.C. Jones, C.H. Arns, A.P. Sheppard, D.W. Hutmacher, B.K. Milthorpe, M.A. Knackstedt, Assessment of bone ingrowth into porous biomaterials using MICRO-CT, Biomaterials 28 (2007) 2491–2504. [4] A.J. Salgado, O.P. Coutinho, R.L. Reis, Bone tissue engineering: state of the art and future trends, Macromol. Biosci. 4 (2004) 743–765. [5] S. Bose, S. Vahabzadeh, A. Bandyopadhyay, Bone tissue engineering using 3D printing, Mater. Today 16 (2013) 496–504. [6] K. Rezwan, Q.Z. Chen, J.J. Blaker, A.R. Boccaccini, Biodegradable and bioactive porous polymer/inorganic composite scaffolds for bone tissue engineering, Biomaterials 27 (2006) 3413–3431. [7] D.W. Hutmacher, Scaffolds in tissue engineering bone and cartilage, Biomaterials 21 (2000) 2529–2543. [8] X. Li, R. Cui, L. Sun, K.E. Aifantis, Y. Fan, Q. Feng, F. Cui, F. Watari, 3D-Printed

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