Novel bioceramic-reinforced hydrogel for alveolar bone regeneration

Accepted Manuscript Novel Bioceramic-Reinforced Hydrogel for Alveolar Bone Regeneration Giorgio Iviglia, Clara Cassinelli, Elisa Torre, Francesco Baino, Marco Morra, Chiara Vitale-Brovarone PII: DOI: Reference:

S1742-7061(16)30399-3 http://dx.doi.org/10.1016/j.actbio.2016.08.012 ACTBIO 4375

To appear in:

Acta Biomaterialia

Received Date: Revised Date: Accepted Date:

22 February 2016 25 July 2016 10 August 2016

Please cite this article as: Iviglia, G., Cassinelli, C., Torre, E., Baino, F., Morra, M., Vitale-Brovarone, C., Novel Bioceramic-Reinforced Hydrogel for Alveolar Bone Regeneration, Acta Biomaterialia (2016), doi: http:// dx.doi.org/10.1016/j.actbio.2016.08.012

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Novel Bioceramic-Reinforced Hydrogel for Alveolar Bone Regeneration Giorgio Iviglia 1,2,*, Clara Cassinelli1, Elisa Torre1, Francesco Baino2, Marco Morra1, Chiara Vitale-Brovarone2 1

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Nobil Bio Ricerche Srl, 14037 Portacomaro (Italy)

Department of Applied Science and Technology, Institute of Materials Physics and Engineering, Politecnico di Torino, 10121 Torino (Italy) *

E-mail: [email protected] (Dr. G. Iviglia) Nobil Bio Ricerche srl

Via Valcastellana, 26 14037 Portacomaro (AT) Italy Tel. +39 0141 202547 – Fax +39 0141 278832

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ABSTRACT The osseointegration of dental implants and their consequent long-term success is guaranteed by the presence, in the extraction site, of healthy and sufficient alveolar bone. Bone deficiencies may be the result of extraction traumas, periodontal disease and infection. In these cases, placement of titanium implants is contraindicated until a vertical bone augmentation is obtained. This goal is achieved using bone graft materials, which should simulate extracellular matrix (ECM), in order to promote osteoblast proliferation and fill the void, maintaining the space without collapsing until the new bone is formed. In this work, we design, develop and characterize a novel, moldable chitosanpectin hydrogel reinforced by biphasic calcium phosphate particles with size in the range of 100300 µm. The polysaccharide nature of the hydrogel mimics the ECM of natural bone, and the ceramic particles promote high osteoblast proliferation, assessed by Scanning Electron Microscopy analysis. Swelling properties allow significant adsorption of water solution (up to 200 % of solution content) so that the bone defect space can be filled by the material in an in vivo scenario. The incorporation of ceramic particles makes the material stable at different pH and increases the compressive elastic modulus, toughness and ultimate tensile strength. Furthermore, cell studies with SAOS-2 human osteoblastic cell line show high cell proliferation and adhesion already after 72 h, and the presence of ceramic particles increases the expression of Alkaline Phosphatase activity after 1 week. These results suggest a great potential of the developed moldable biomaterials for the regeneration of the alveolar bone. Keywords: Composite, calcium phosphate, hydrogel, scaffold, alveolar bone regeneration

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1. INTRODUCTION The restoration of teeth by using titanium dental implants is nowadays a quite common procedure in oral surgery [1]. The American Association of Oral and Maxillofacial Surgeons recently reported that 69% of adults aged 35 to 44 have lost at least one of their permanent teeth and, by age 74, almost 26% of adults have lost all of their permanent teeth. More than 300,000 dental implants are placed per year worldwide and, until 2020, this number is expected to increase [1,2]. Tooth loss is a possible consequence of traumas or periodontal diseases, such as gingivitis, periodontitis or tissue decay. The rate of success of dental implants is around 98%, but it should be taken into account that the positive fate of a surgical procedure involving the insertion of a titanium screw still depends on the quality and quantity of alveolar bone which is present in the extraction site. In order to get a successful implant insertion, the alveolar ridge should have a minimum dimension, which is of 5 mm of width in the maxilla and of 10 mm of bone height [3]. If these dimensions are not available, an augmentation strategy will be necessary using grafting procedures. A huge alveolar bone loss of around 2 mm in vertical and almost 50% in horizontal occurs in the first 6 months after surgery; a continuous bone resorption occurs if no treatment is provided during the next year after the extraction procedure, and an average of 60% of bone loss could be reached in 3 years [4–9]. These data show the importance of implementing a strategy to augment the alveolar bone volume in order to provide a stable support for the implant and the future dental crown. The common use of autografts or allografts is affected by the morbidity and the risk of disease transmission associated with the donor site, thus the use of man-made bone graft biomaterials is more and more attractive [10–12]. A bone graft material should be biocompatible, degradable, osteoconductive, osteoinductive and should mimic the ECM of bone, in order to allow cell infiltration, proliferation and new bone formation. Hydroxyapatite (HA) and β-tricalcium phosphate (βTCP) are well known ceramic materials, widely used in bone tissue engineering as bone grafts in the form of particles or three-dimensional scaffolds [13–15]. HA (Ca10(PO4)6(OH)2) is the most abundant component in 3

native bone (around 65% of the inorganic phase) but, despite many studies showing osteoconduction stimulation and good mechanical properties, the slow degradation rate of HA limits its use as a bone filler only [16–18]. Combination of HA with βTCP allows the degradation properties to be managed, since βTCP materials have a degradation rate 3–12 times faster than HA, and the mechanical properties to be maintained good [19]. The ratio between HA and βTCP is a key point to obtain satisfactory mechanical strength and resorption kinetics and, ultimately, to stimulate osteointegration. Animal studies demonstrated that HA/βTCP biphasic materials in the rate of 25/75 wt.% allow new bone formation [15]. In the post-extractive site, it is important to use materials which can not only fill the void, but also be shaped in an easy manner and be not prone to migration, hence particles and threedimensional rigid scaffolds are not the best choice. To overcome this issue, we developed a pectin/chitosan-based polyelectrolyte hydrogel reinforced with biphasic calcium phosphate particles. Combining ceramics with polymeric materials provides – at least potentially – many advantages, in particular for dental practice, such as shape control, optimization of adhesion between implant and surrounding bone tissue, easy adaptation of the biomaterial to the extraction site, promotion of clot formation and prevention of ceramic particles migration [16,20–23]. Few composite products for bone regeneration are currently available on the market; perhaps the most popular is Bi-Oss Collagen® that exhibits an excellent biocompatibility but tends to resorb very quickly, with consequent loss of the mechanical properties and migration of the inorganic particles into the surrounding environment [24]. In the attempt to obtain composite biomaterials with better performances, in this work we developed a new hydrogel formulation based on two natural polymers, chitosan and pectin, to be combined with HA/βTCP particles. Chitosan is a natural polycationic material derived from chitin extracted from crustaceous exoskeleton, composed of β–(1,4)–glucosamine and N-acetyl-Dglucosamine; it has recently aroused great interest in bone tissue engineering applications due to its biocompatibility, intrinsic antibacterial nature, and ability not to stimulate a foreign body reaction 4

and, on the contrary, to promote cell adhesion, proliferation and differentiation [25,26]. Furthermore, chitosan has a backbone similar to glycosaminoglycan, the major component of bone ECM [21,27]. In order to stabilize the structure, we coupled chitosan with a natural polyanionic polysaccharide, pectin, which is a major component of cell walls of citrus or apple peel by-products. Pectin consists in a poly(D-galacturonic acid) chain, with a partially methoxylated carboxyl group [28]. We exploited the ionic interaction that occurs between chitosan and pectin [29–31] to develop a novel hydrogel mimicking an ECM-like environment for osteoblast cells, while biphasic calcium phosphate particles were added to mimic the inorganic phase of bone. Here, the physico-chemical, mechanical and in vitro biological properties of this new HA/βTCP-reinforced hydrogel were systematically investigated and a comparison with the base hydrogel was reported.

2. MATERIALS AND METHODS 2.1 Materials β-tricalcium phosphate (βTCP), chitosan with medium molecular weight from crab shell (Mw = 400 kDa), pectin from citrus peel and all other chemicals were purchased from Sigma-Aldrich. HA particles (average size 5 µm) were purchased from Fluidinova. 2.2 Preparation of calcium phosphate particles HA and β-TCP were mixed in a percentage of 25 wt.% and 75 wt.% respectively, on the basis of a previous study by Morra et al. [15]. Briefly, the preparation of the ceramic slurry involved the mixing of the HA and β-TCP powders (31 wt.%) in ultrapure water (69%). Guar gum was dissolved previously in ultrapure water (3 wt.% of the water content), and used a pore forming agent. The slurry was then desiccated in oven at 80 °C for 8 h, and then shaped in circular disks using a hydraulic press (Mignon EA/SSN) with a pressure of 17.5 MPa for 10 s. The ceramic disks were then sintered in a furnace at 1100 °C for 1 h in air (heating rate 1 °C/min) [15]. Sintered disks were 5

ground in particles using a grinder machine (GM200, Retsch) and the ceramic powder was sieved in order to obtain a range of particles between 100 and 300 µm. Eventually, the biphasic calcium phosphate particles obtained were washed in ultrapure water and desiccated overnight at room temperature under laminar flow. 2.3 Preparation of composite hydrogel We dissolved 7 wt.% of pectin powder and 4 wt.% of chitosan powder in acetate buffer solution to obtain a pectin/chitosan ratio of 20/80. The solution was stirred for 6 h, then it was centrifuged and the polyelectrolyte (PEI) precipitated was collected. PEI complex (10 wt.%) was mixed using a Thinky Mixer (Retsch) with HA/βTCP particles (90 wt.%). The resulting slurry was poured in a circular shape polyester mold and freeze-dried with Lyo5P for 12 h, at 0.06 mbar and -56 °C in order to obtain pectin/chitosan_ceramic particles (PCC) composite hydrogels. For the pectin_chitosan based hydrogel (PC), we poured the PEI solution into a circular-shape polyester mold and freeze-dried the sample by Lyo5P for 12 h, at 0.06 mbar and -56 °C. 2.4 Chemical characterization Attenuated Total Reflectance Infrared spectroscopy (ATR-IR) was recorded using a Nicolet iS 10 produced by Thermo Scientific and equipped with a diamond crystal. Raw and processed materials (PC and PCC) were placed on the crystal and kept in place by a specific crimping tool. Experimental setup was conducted by acquiring 32 scans in the range of 500–4000 cm-1, with a resolution of 4 cm-1. 2.5 Morphological analysis The microstructure of the PCC and PC materials was studied in a nondestructive manner by microcomputed tomography (µ-CT), with a desktop µ-CT scanner (SkyScan 1174, Aartselaar, Belgium). The scanner was set at a voltage of 50 kV and a current of 800 µA; the sample was rotationally scanned (0-180°) at 7.55 µm pixel resolution. The exposure time per projection was 2300 ms and no 6

filter was used. No contrasting agents were used and the sample had a cylindrical size of 10 mm of diameter and 10 mm of height. Sample volume reconstruction in 3D was performed by making use of N-Recon software. We quantified the microstructural parameters of the samples in 3D by making use of CT-Analyzer software, after having binarized the 2D tomographic images (“slices”). The threshold levels of the grayscale images were equally adjusted for all the samples to allow a reliable measurement and comparison of the porosity. Scanning Electron Microscopy (SEM) analysis was performed to analyze the morphology of PCC and PC hydrogels. The samples were mounted on aluminum stubs and sputtered with gold at 15 mA for 2 min by using Agar Sputter Coater. The morphology of pores and the distribution of particles in the PEI structure were captured using a scanning electron microscope (EVO MA10 system, Zeiss) equipped with a micro-analysis system AZTec (Oxford University UK). 2.6 Mechanical characterization The mechanical properties of PCC and PC materials were evaluated through unconfined compression and cyclic compression testing by using Bose Electroforce equipped with a 100-N load cell. For unconfined compressive tests, PCC and PC samples were cut in a circular shape with 10 mm of diameter, and 10 mm of gauge length (n=3). The mechanical properties were assessed in both dry and hydrated conditions (soaked in PBS at 37 °C for 24 h). In preliminary tests no sample could reach a failure, so we decided to perform the compressive test until 40 % of the height of the sample, which is a sufficient value to see elastic and plastic deformation of the samples. A constant crosshead speed of 0.2 mm/min was applied. Stress (σ (MPa)) was obtained by dividing the applied force (N) by the cross-sectional area (mm2). Compressive elastic modulus was calculated from the linear region of the stress-strain curve between 0.5 and 5 % of strain, and the stress at 40% of strain was cautiously assumed as collapse stress (this is an approach commonly followed in the literature to assess the compressive strength of polymer/bioceramic composite scaffold [32,33]). The area under the stress-strain curve was calculated and related to the toughness of the sample. For cycling 7

testing, 100 loading and unloading cycles were performed on PC and PCC samples in hydrated condition (n=3). A first load until 16% of initial height was applied, and then a sinus cyclic loading curve with a frequency of 0.5 Hz and a constant crosshead speed of 1.2 mm/min was applied between 16 and 41% of initial height. Stress-strain curves for PCC and PC samples were reported; the amount of energy adsorbed and the percentage of recovery during the deformation cycle were calculated every 10 cycles. 2.7 Degradation properties The stability of PCC and PC hydrogels was analyzed at three different pH (2.5, 5.5 and 7.4) at 37 °C (n=3) following an experimental protocol proposed elsewhere [34]. Briefly, for the solution at pH 2.5, 21 g of citric acid monohydrate were dissolved in 500 ml of ultrapure water in a 1000 ml volumetric flask, then 200 ml of sodium hydroxide 1 M were added and the solution diluted to the mark with water; 40.4 ml of this solution were mixed with 59.6 ml of 0.1 M hydrochloric acid. For the solution at pH 5.5, 0.49 g of sodium phosphate dihydrate, 13.8 g of potassium phosphate and 0.49 g of sodium chloride were dissolved in 1000 ml of ultrapure water. The solution at pH 7.4 was prepared by dissolving 0.78 g of sodium phosphate dihydrate, 0.097 g of potassium phosphate and 4 g of sodium chloride in 500 ml of ultrapure water. The samples, three for each type, were cut in a cylindrical shape (10 mm of diameter and 2 mm of height) and immersed in the three different solutions, after recording their initial weight (W0). The degradation study was carried out for 1 week, during which samples were taken out at specific time intervals, freeze-dried and weighted (Wf). The mass loss percentage was calculated by using the following equation: Mass Loss (%) = [(W0 – Wf)/W0]x 100 (1) 2.8 Hydration properties For hydration kinetics, PC and PCC samples (10 mm diameter, 2 mm thickness, n=3) were allowed swelling in three different solutions (the same used for the degradation test pH 2.5, 5.5 and 7.4) 8

after recording the initial weight (W0) at 37 °C. The swollen disks were taken out of the solution at regular time intervals, blotted with filter paper to remove excess surface water, and their swollen weights (Ws) were noted. The water uptake by the PCC and PC networks was determined by the following equation: Hydration Degree (%) = [(Ws – W0)/Ws]x 100 (2) 2.9 In vitro studies Osteoblast-like SAOS-2 cells were used in cell growth experiments. Experimental cell culture medium (BIOCHROM KG, Berlin) consisted of Minimum Eagle’s Medium without l-glutamine, 10% fetal bovine serum, streptomycin (100 µg/l), penicillin (100 U/ml), and 2 mM l-glutamine in 250-ml plastic culture flasks (Corning™). Cells were cultured at 37°C in a humidified incubator equilibrated with 5% CO2. Cell suspension was obtained by adding 2 mL of a sterile 0.5% Trypsin-EDTA solution (GIBCO by LIFE TECHNOLOGY, ref 15400-054) to a 250 mL cell culture flask, resuspended in the experimental cell culture medium, and diluted to 1×105 cells/ml. For these experiments, 5 ml of the cell suspension were seeded into 6-well tissue culture polystyrene plates (9.6 cm2 of growth area; Falcon™), containing the samples. Experiments were performed in triplicate. After 72 h and 2 weeks, samples were carefully rinsed with PBS and fixed in 5% glutaraldehyde-PBS. Samples were dehydrated using increasing concentration of ethanol in water-ethanol solutions up to 100% ethanol. Dehydrated samples were gold sputter-coated (Agar Sputter Coater) and SEM analysis was performed using a EVO MA10 (Zeiss), equipped with a micro-analysis system AZTec (Oxford Instruments, UK). The expression of ALP gene as a cell differentiation marker was assessed by using the real time reverse transcription polymerase chain reaction (qRT-PCR). Osteoblast–like SAOS-2 cells were cultured as previously described and the total RNA was extracted after 1 week by using MagMax Total RNA Isolation Kit (Applied Biosystems, Milan, Italy) following the manufacturer’s instruction. RNA quality was assessed by checking the A260/A280 ratio (1.6–2.0). Then total RNA was used as a template for cDNA 9

synthesis using random hexamers as primer and multiscript reverse transcriptase (high capacity cDNA RT Kit from Applied Biosystems). cDNA amplification and relative gene quantification were performed by using Taq Man probe and primers from Applied Biosystems (Hs 01029144_m1, ALP). Real time PCR was performed in duplicate for all sample and targets on a Step-One instrument (Applied Biosystems) by using software Step-One version 2.1. PCRs were carried out in a total volume of 20 µl and the amplification was performed as follows: after an initial denaturation at 95 °C for 10 min, the PCR was run for 40 cycles at 95 °C for 15 s and at 60 °C for 1 min. We normalized the content of cDNA samples by making use of the comparative threshold (Ct) cycle method, consisting in the normalization of the number of target gene copies versus the endogenous reference gene GAPDH. The Ct is defined as the fractional cycle number at which the fluorescence generated by cleavage of the probe passes a fixed threshold baseline when amplification of the PCR product is first detected. For comparative analysis of gene expression, data were obtained using ∆CT method.

2.10 Inflammatory response The murine macrophage cell line J774.1A (European Collection of Cell Cultures) was maintained in Dulbecco’s modified Eagle’s medium (Gibco Invitrogen, Cergy-Pontoise, France) supplemented with 10% fetal bovine serum, penicillin (100 U ml−1), streptomycin (100 µg ml−1) and 4 mM lglutamine. Cells were grown in a 100% humidified incubator at 37 °C with 10% CO2 and passaged 2–3 days before use. The J774.1A cells (2 × 104 cells ml−1) were seeded into 6-well tissue culture polystyrene plates (9.6 cm2 of growth area; Falcon™), containing the samples. RNA extraction was performed on the cells grown on the polystyrene in presence or absence of PC or PCC scaffolds. The extraction protocol was already reported in the previous paragraph. Briefly, after 4 h, RNA was isolated from J774.1A cell line by using the MagMax-96 Total RNA Isolation Kit (Life Technologies), and then reverse transcripted by using the High-Capacity cDNA Reverse Transcription Kit (Applied Biosystems), according to the manufacturer’s instructions. Total RNA 10

concentration was measured by using a spectrophotometer (UV-1700, Shimadzu) and RNA quality was assessed by evaluating A260/A280 ratio ranging from 1.8 to 2.1. Real-time PCR was performed with the TaqMan Gene Expression Master Mix (Applied Biosystems) with the Applied Biosystems StepOne Plus instrument (Applied Biosystems). The primer sets for the Real-time PCR of mouse interleukin-1β (IL-1β), interleukin-6 (IL-6), interleukin-10 (IL-10), and glyceraldehyde-3-phosphate dehydrogenase (GAPDH) (Applied Biosystems Assay’s ID: Mm01336189_m1, Mm99999064_m1, Mm99999056_m1, Mm03302249_g1 respectively) were chosen from the collection of the TaqMan Gene Expression Assays (Applied Biosystems). The analysis was conducted using the method of comparative CT (∆CT), which was designed following the manufacturer’s instructions.

2.11 Statistical analysis Experimental data were presented as mean ± standard deviation. Statistical differences between groups were analyzed

using two-way ANOVA using Tukey’s multiple comparison test and

Student’s t-test. Statistical significance was represented as *p<0.05, **p<0.01, ***p<0.001, ****p<0.0001.

3. RESULTS AND DISCUSSION 3.1 Chemical characterization Figure 1 a shows the ATR-IR spectra of pectin powder, chitosan powder, pectin/chitosan polyelectrolyte complex (PC), HA/βTCP

particles and pectin/chitosan_HA/βTCP composite

hydrogel (PCC). The analysis confirms that all peaks belong to inorganic material; in particular, triply degenerated asymmetric stretching mode (ν3) of the P-O bond of the phosphate group is associated to the peaks at 1125 cm-1 for tricalcium phosphate and at 1025 cm-1-1010 cm-1 for hydroxyapatite [15,35]. Typical spectra of polysaccharide were shown in the case of pectin and chitosan powder. The region between 3700 cm-1 and 3000 cm-1 for pectin and chitosan is assigned 11

to the O–H stretching vibration (νOH), while the region between 3000–2800 cm-1 belongs to C-H stretching vibration (νCH) (Figure 1 a). Deeper analysis on pectin spectrum shows two bands associated with the stretching vibration at 1740 cm-1 of carbonyl group, corresponding to the methyl ester group (COOCH3) and carboxyl acid (COOH), while the band at 1606 cm-1 belongs to the stretching vibration of the carbonyl group of the carboxylate ion (COO-) (Figure 1 b). Concerning the chitosan spectrum, the band at 1647 cm-1 is due to the C=O stretching vibration of amide I, whilst the band at 1580 cm-1 is due to the NH bending amide II, maybe overlapped to the N-H vibration of the amine groups (Figure 1 b). Band assignment is consistent with available literature [36,37]. PC spectra show the formation of PEI complex; a shift of amine band to 1557 cm-1 due to the interaction between the positive charge of chitosan, NH3+, and the negative charge of pectin, COO-, was detected (Figure 1 b). As expected, the spectrum of PCC sample show bands associated to both the PEI complex (the band at 1557 cm-1) and the inorganic phase (bands between 950 cm-1 and 1140 cm-1) (Figure 1 b) [38]. 3.2 Morphological analysis SEM analyses at different magnifications for PC and PCC samples are reported in Figure 2. The addition of ceramic particles reduces the pore size and the total porosity, but reinforces the trabeculae and the 3D porous structure of the materials. SEM pictures suggest that the dispersed particles are well distributed in the PEI matrix. In vitro and in vivo studies from the available literature show that a porosity up to 50 %, and pore size in the range between 100 to 400 µm are the optimum for bone healing and show higher alkaline phosphatase activity and new bone formation compared with materials with a porosity lower than 100 µm and higher than 400 µ m [39]. A µCT study (Figure 3 a, b and c) confirmed the results of SEM analysis. Figure 3 a and 3 c reveal that the bioceramic particles are uniformly and homogeneously distributed throughout the sample volume. Most pores of PCC sample are in the range of 50-200 µm (the largest pores do not exceed 350 µm), while the PC scaffold has a broader distribution of pore size, with most pores 12

located between 200 and 400 µ m and the maximum pore size around 650 µm (Figure 3 b). Direct comparison of mean pore sizes (St.Sp) further confirms the decrease of pore dimension due to the introduction of ceramic particles (250 vs. 120 µ m). The total porosity decreases by adding ceramic particles, too, from around 74% for PC sample to 55% in the case of PCC scaffold. Threedimensional porous scaffolds should allow cell infiltration and facilitate vascular invasion. Furthermore, it is necessary that the ceramic particles are homogenous dispersed in the polymeric matrix – which was observed in this study (Figure 3 a and 3 c) – in order to promote uniform osteogenesis. Both micro-porosity and macro-porosity are important morphological properties [40,41]. It has been shown that microporosity plays an important role in influencing cell adhesion to the biomaterial surface, while macro-porosity promotes blood vessel infiltration, nutrients transportation and clot formation, which is important to stimulate the overall healing process, besides allowing cell infiltration and new bone growth [42–44]. 3.3 Mechanical characterization The incorporation of ceramic particles into soft hydrogel scaffolds is a valuable strategy to improve the mechanical properties of the base material [45,46]. A compressive test was performed on PC and PCC scaffolds, and the results are shown in Figure 4 a and in Table 1. The addition of HA/βTCP particles increases the elastic modulus, the compressive strength and the toughness of the material, both in as prepared and hydrated conditions. Compressive elastic modulus for PC scaffolds was 1005 ± 250 kPa in as prepared condition, instead it was twofold higher for PCC scaffold, 2559 ± 596 kPa. Hydration kinetics data show that the water uptake at pH 7.4 was higher in both cases, and the compressive elastic modulus decreases to 33 ± 8 kPa for PC material and to 65 ± 9 kPa for PCC scaffolds. Alveolar bone regeneration is a low-bearing application, where providing stability and three-dimensional shape for functional and aesthetic reasons is more important than mechanical strength [16,47]. PC materials show a soft structure, and the incorporation of ceramic particles inside the PEI matrix makes the material tougher (Figure 4 a). 13

For example, the toughness was 34.3 ± 4.0 kJ/m3 and 222.8 ± 1.9 kJ/m3 for PC and PCC scaffolds, respectively, in as prepared condition. There is a relative paucity of literature dealing with the toughness of bone tissue engineering scaffolds and, thus, comparison with previous works is difficult; however, it is interesting to point out that the toughness of PCC scaffold is ten times higher than that assessed for single-phase glass-ceramic porous scaffolds [48]. After 24 h at 37 °C in a solution with pH 7.4, both hydrated samples show a decrease of toughness (2.4 ± 0.5 kJ/m3 for PC scaffold and 12.0 ± 2.1 kJ/m3 for PCC scaffold) compared to as prepared materials. A similar behavior was also found by Liu et al. on bioglass scaffolds [49]. The toughness of PCC material in hydrated condition is comparable, as order of magnitude, to the value found by other authors for hydroxyapatite scaffolds [50]. Ceramic particles, as expected, reduce the porosity but reinforce mechanically the scaffolds, maintaining a range and a degree of porosity that should still allow cell infiltration and proliferation [40]. Cyclic stress is a common situation to which the scaffolds used for alveolar bone regeneration are subjected. The stability and integrity of the scaffolds are the most important properties that they should have, since material fragmentation could provoke an inflammatory response and a huge deformation could cause the collapse of the defect. Hydrogel PC scaffolds show a stable energy adsorbed (and recovery network) from the 5th to the 100th cycle; the addition of 90% of ceramic particles does not affect too much the trend, although the energy adsorbed was slightly higher for the PCC composite due to the reduced mobility of the PEI network (Figure 4 b, c). The results are affected by the deformation of the swollen surface of the samples that do not return instantaneously to the original shape after cessation of compression stress. After 100 cycles no fragmentation was detected for both samples; furthermore, PC and PCC materials, after being left without load to recover the (original) shape for 5 min, showed a variation of the final diameter with respect to the initial one of 1.4 ± 0.59 % and 7.52 ± 1.32 %, respectively (more than 90 % of recovery for both samples). This means that in vivo the PCC scaffold is able to maintain the shape and to avoid the collapse of the site. Furthermore, tooth extraction site has a non-regular shape, hence ensuring the 14

possibility to customize the dimension of the implant material in order to fit the socket is an important characteristic for dental practice. 3.4 Hydration studies at different pH A high degree of swelling allows cell infiltration into the scaffolds and maximizes the probability of cell growth in a three-dimensional structure [22]. Furthermore, high swelling behavior improves the capability of the scaffold to adsorb nutrients from the culture media. Chitosan/pectin complexes are formed by a ionic interaction between positive charges of chitosan and negative charges of pectin; hence, the PEI network exhibits a pH-sensitive swelling [34,51]. We tested the swelling behavior in three different solutions with three different pH (2.5, 5.5 and 7.4), and the results are shown in Figure 5. PC hydrogels showed a significantly different behavior depending on the pH of the solution. If the pH of the solution changes, the degree of interaction between pectin and chitosan changes, too, and the swelling increases or decreases depending on the degree of dissociation of the complex (Figure 5 a). At low pH value (2.5), pectin is neutralized and free positive charges (NH3+) appear in the network; COOH from pectin chain allows the swelling of the material until the value of 1460.42 ± 81.33 % of the initial mass, after 6 h. After 24 h, PC hydrogels soaked in a strongly acidic solution (pH 2.5) showed a huge increase in the water uptake; this behavior could be associated to the total degradation of pectin and to the loss of PEI network that allows the highest water adsorption, 2495.34 ± 33.87%. At alkaline pH (7.4), free negative charges appear inside the network, since at higher pH chitosan is partially neutralized and the PEI network loosens, thus leading to a hydration degree of 1880.2 ± 218.8 %, after 6 h. At pH 7.4, the partial neutralization of NH3+ has not caused a total loss of the PEI network, since chitosan has still some ionized groups bonded with COO– of the pectin chain. However, after 6 h, the swelling ratio started to decrease due to an initial degradation of the network. Since the pKa of pectin is 4.0 and that of chitosan is 6.0, at the pH of 5.5 over 99% of pectin is still in its ionized form and chitosan exists in both ionized (NH3+) and unionized forms (NH2). The swelling percentage of PC network in the solution with a pH of 5.5 was 1443.8 ± 14.4 % after 6 h; then, the capability of the network to adsorb the solution is 15

reduced compared to the other cases. Specifically, due to the presence in the network at pH 5.5 of intramolecular H-bonding between COOH3 and OH, which makes the PEI matrix more stable, the PC hydrogel after 24 h at pH 5.5 show a significantly lower swelling degree (687.69 %) compared to the same material at the pH of 7.4 (around 1450 %) and 2.5 (around 2500 %). Addition of HA/βTCP particles stabilizes, in general, the swelling properties of the PC-based hydrogel (maximum swelling degree of the composite around 300 %) (Figure 5 a). The swelling behavior between PC and PCC materials is significantly different, while no statistically significant differences were assessed among PCC samples during the hydration kinetics in the three different solutions (pH of 2.5, 5.5, 7.4) (Figure 5 b). As expected, the maximum water uptake for PCC scaffold was lower compared to the PC hydrogel in all the conditions considered: for example the swelling degree for PCC scaffold at pH 7.4 was about 275.05 ± 35.45 % after 6 h, almost 10 fold lower than the value reported for PC hydrogel at the same pH. The maximum percentage of solution uptake for PCC material was reached after 2 h and was stable until 24 h. As already observed for PC hydrogels, also in the case of PCC samples the lowest value of water uptake was shown for the sample soaked in the solution at pH 5.5. Providing a stable material could allow the dental technician to use the scaffold independently of the environment (especially the pH), since it will work always in the same manner avoiding unexpected behaviors (Figure 5 c). Furthermore, the swelling of the PCC scaffold is high enough to ensure cell infiltration and nutrient transportation, and a sufficient mechanical strength is still maintained. The incorporation of ceramic particles into the PC matrix reduces the swelling capability of the materials, which is maybe due to the interaction between the particles and the network. In particular, pectin carboxylic groups could be ionically crosslinked by calcium ions (Ca2+), thereby forming the so-called “egg box” structure, where a divalent cation is bonded with different carboxylic anions [28]. Furthermore, the high percentage of rigid, non-deformable ceramic particles (which cannot swell) used in this work (90 wt.%) reduces the influence of the PC matrix in the swelling properties. In alveolar bone regeneration, the use of sponges is desirable, since they 16

could easily fit into the alveolar bone defect, be simply cut by scissors or a lancet, and be easily molded in the periodontal cavity (Figure 5 d) [16,43,52]. The addition of calcium phosphate particles does not influence the moldability of the PC hydrogel base material (Figure 5 d). 3.5 Degradation properties at different pH After implantation, the composite hydrogel should degrade, and new bone should form and replace it. Furthermore, during degradation the osteconductive HA/βTCP ceramic particles come into direct contact with the newly formed bone and should further promote osseointegration. The degradation behaviors of PC and PCC samples have been tested in three different conditions of pH (2.5, 5.5, 7.4) for 1 week and the results are shown in Figure 6. During the degradation process, the mechanisms are similar to those obtained during the hydration kinetics, since the solutions used were the same [53]. PC hydrogel shows the highest degradation rate at pH 2.5 with almost 100% mass loss after 1 week; this mass loss was twice compared to that of PC treated at pH 7.4 and almost three times with respect to the hydrogel soaked at pH 5.5, which lost around 40% of the initial mass (Figure 6 a, b). At the pH of 2.5, the PEI network loosens due to the neutralization of the pectin. The same behavior is reported for the PCC sample, that lost around 40 % of the initial mass after 1 week in the solution at the pH of 2.5. These results are due in part to the neutralization of COO– ions, and in part to the dissolution of ceramic particles in a highly acidic environment. Furthermore, the ionic interaction between the two polymer chains is reversible and non-permanent, so in aqueous solution this association is gradually reduced and then lost with time. PCC scaffolds tested in the solution at pH of 5.5 and 7.4 showed a similar percentage of initial mass loss, around 10 %. The addition of ceramic particles makes the materials more resistant and less susceptible to pH variation (Figure 6 c).

17

3.6 In vitro studies Adhesion, spreading and morphology of osteoblast-like cells cultured for 14 days on PC and PCC are shown in Figure 7. SEM images show that, after 72 h, SAOS-2 cells seeded on the PC scaffold surface exhibit a round shape and low density, followed by high proliferation and formation of a network of cells. In particular, SEM images after 2 weeks show the formation of large cords of cells, which prefer to stay together rather than adhere and spread on the surface of the materials (Figure 7 a). The addition of ceramic particles increases the adhesion and proliferation of osteoblast cells at 72 h, and the SEM images reveal that, after 2 weeks, SAOS-2 cells are completely adhered on the surface and infiltrated inside the porosity, forming bridges and creating a strong and consistent layer of cells (Figure 7 b). The higher attachment and proliferation of cells on the PCC materials are due to the presence of the calcium phosphate phase that conducts osteoblast proliferation [54] Furthermore, the addition of micro-particles increases the roughness and the area of the surface, which promote the osteoblast proliferation and adhesion [23]. Since PC and PCC materials show different cell morphology at two weeks but both demonstrated cytocompatibility and high proliferation, we thus performed ALP gene-expression after 7 days of SAOS-2 cell culture. The results are reported in the bar graph in Figure 8 and demonstrated that the presence of HA/βTCP particles in the PCC composite scaffolds promotes ALP activity, which is 3-fold higher than that of PC-based hydrogels. In vitro studies show a good biocompatibility of PCC scaffold, furthermore the addition of ceramic particles could promote alveolar bone formation and infiltration inside the PCC composite hydrogel. 3.7 Inflammatory response Gene expression analysis on IL1β, IL6 and IL10 is reported in Figure 9. Inflammatory response was evaluated on the cells grown on the polystyrene in presence of PC, PCC and without the scaffold. The test was performed in order to evaluate the possible inflammation which could be provoked on the tissue surrounding the implanted material. Degradation products, chemical 18

composition and swelling behavior could influence the response of the surrounding tissue; if the material does not stimulate any cytokines expression, a foreign body reaction is avoided. As shown in Figure 9, no material elicits a pro-inflammatory response, and the expressions of IL1β and IL6 are comparable between the cells grown in presence of PC/PCC and on the control (polystyrene alone). A slight increase in the expression of IL10 was reported for PCC material compared with the polystyrene control, conversely no significance was assessed to the increase of IL10 for PC material. Bone resorption occurs when inflammatory mediators reach a critical concentration that depends on the expression of pro-inflammatory cytokines, such as the interleukin (IL) family, of which IL1β is the most studied member due to its role in acute and chronic inflammatory and autoimmune disorders. On the opposite site, the inflammatory level is controlled by the expression of anti-inflammatory cytokines, such as IL10. In normal physiological conditions, there is a balance between bone formation and bone resorption, but this balance can be altered in certain inflammatory conditions. This equilibrium is altered by the action of pro-inflammatory cytokines, such as IL1β that induces osteoclastogenesis, while counteraction by anti-inflammatory mediators, such as IL10, inhibits osteoclastogenesis. An anti-inflammatory response seems to occur in both PC and PCC scaffolds, although a slight increase of IL10 gene expression was detected in the latter compared to the former (Figure 9). The hypothesis is that this anti-inflammatory behavior could be due to the properties of pectin, which has been shown to exhibit an anti-inflammatory action in many studies [55–57]. Further and in-depth studies have to be done in order to fully elucidate this behavior, but this preliminary work confirms that PCC could be used as a safe bone graft for alveolar bone regeneration, without the risk of inducing inflammation in the surrounding tissue. 4. CONCLUSIONS Alveolar bone regeneration is a procedure that requires a biomaterial with some peculiar characteristics such as mouldability in order to fill the irregular void, biodegradability, swelling capacity to promote infiltration of nutrients, absence of inflammatory response on the surrounding 19

tissue, osteoconduction and stimulation of new bone formation. Composite scaffolds comprising a pectin/chitosan base hydrogel filled with HA/βTCP particles (PCC) was successfully prepared and characterized. The results show that PCC material can be easily molded to fit the defect size/shape and has a good stability in different pH condition, with a high swelling degree (up to 300 % of the initial weight). Mechanical characterization demonstrated that the addition of ceramic particles increases the mechanical strength compared to the base hydrogel (toughness of PCC scaffold was around 220 kJ/m3, and compressive elastic modulus was 2.5 MPa in dry condition). The scaffold morphology and porosity as well as the presence of osteoconductive HA/βTCP micro-particles promote highly osteoblast adhesion and proliferation with 2-fold higher ALP gene expression at 1 week compared to PC scaffold. Gene expression results demonstrated that PCC scaffold elicits antiinflammatory and pro-osteogenic responses; this results confirm that PCC biomaterial could be a promising tool for application in alveolar bone regeneration. Acknowledgments GI would like to acknowledge Nobil Bio Ricerche srl and Politecnico di Torino, Department of Applied Science and Technology, for the PhD fellowship. This research was funded by Nobil Bio Ricerche srl.

20

REFERENCES [1]

L. Gaviria, J.P. Salcido, T. Guda, J.L. Ong, Current trends in dental implants, J. Korean Assoc. Oral Maxillofac. Surg. 40 (2014) 50.

[2]

A. Gupta, M. Dhanraj, G. Sivagami, Status of surface treatment in endosseous implant: a literary overview, Indian J. Dent. Res. 21 (2010) 433–438.

[3]

G.M. Raghoebar, R.H. Batenburg, A. Vissink, H. Reintsema, Augmentation of localized defects of the anterior maxillary ridge with autogenous bone before insertion of implants, J. Oral Maxillofac. Surg. 54 (1996) 1180–5; discussion 1185–6.

[4]

J. Liu, D.G. Kerns, Mechanisms of guided bone regeneration: a review, Open Dent. J. 8 (2014) 56–65.

[5]

F. Van Der Weijden, F. Dell’Acqua, D.E. Slot, Alveolar bone dimensional changes of postextraction sockets in humans: A systematic review, J. Clin. Periodontol. 36 (2009) 1048– 1058.

[6]

L. Schropp, A. Wenzel, L. Kostopoulos, T. Karring, Bone healing and soft tissue contour changes following single-tooth extraction: a clinical and radiographic 12-month prospective study, Int. J. Periodontics Restorative Dent. 23 (2003) 313–323.

[7]

A. Tallgren, The continuing reduction of the residual alveolar ridges in complete denture wearers: A mixed-longitudinal study covering 25 years, J. Prosthet. Dent. 89 (2003) 427– 435.

[8]

S. Bernstein, J. Cooke, P. Fotek, H.-L. Wang, Vertical bone augmentation: where are we now?, Implant Dent. 15 (2006) 219–228.

[9]

F.G. Draenert, D. Huetzen,

A. Neff, W.E.G. Mueller, Vertical bone augmentation 21

procedures: Basics and techniques in dental implantology, J. Biomed. Mater. Res. - Part A. 102 (2014) 1605–1613. [10] F.-M. Chen, Y. Jin, Periodontal tissue engineering and regeneration: current approaches and expanding opportunities, Tissue Eng. Part B. Rev. 16 (2010) 219–255. [11] J.D. Bashutski, H.L. Wang, Periodontal and Endodontic Regeneration, J. Endod. 35 (2009) 321–328. [12] S. Elangovan, S. Srinivasan, S. Ayilavarapu, Novel regenerative strategies to enhance periodontal therapy outcome, Expert Opin. Biol. Ther. 9 (2009) 399–410. [13] D. Tadic, M. Epple, A thorough physicochemical characterisation of 14 calcium phosphatebased bone substitution materials in comparison to natural bone, Biomaterials 25 (2004) 987–994. [14] H. Yuan, M. Van Den Doel, S. Li, C. A. Van Blitterswijk, K. De Groot, J.D. De Bruijn, A comparison of the osteoinductive potential of two calcium phosphate ceramics implanted intramuscularly in goats, J. Mater. Sci. Mater. Med. 13 (2002) 1271–1275. [15] M. Morra, G. Giavaresi, M. Sartori, A. Ferrari, A. Parrilli, D. Bollati, R. R. Y. Baena, C. Cassinelli, M. Fini, Surface chemistry and effects on bone regeneration of a novel biomimetic synthetic bone filler, J. Mater. Sci. Mater. Med. 26 (2015). [16] T. Matsuno, K. Omata, Y. Hashimoto, Y. Tabata, T. Satoh, Alveolar bone tissue engineering using composite scaffolds for drug delivery, Jpn. Dent. Sci. Rev. 46 (2010) 188–192. [17] L. Zhang, N. Hanagata, M. Maeda, T. Minowa, T. Ikoma, H. Fan, X. Zhang, Porous hydroxyapatite and biphasic calcium phosphate ceramics promote ectopic osteoblast differentiation from mesenchymal stem cells, Sci. Technol. Adv. Mater. 10 (2009) 025003.

22

[18] M. Alcaide, M.C. Serrano, R. Pagani, S. Sánchez-Salcedo, M. Vallet-Regí, M.T. Portolés, Biocompatibility markers for the study of interactions between osteoblasts and composite biomaterials, Biomaterials 30 (2009) 45–51. [19] S. Schaefer, R. Detsch, F. Uhl, U. Deisinger, G. Ziegler, How Degradation of Calcium Phosphate Bone Substitute Materials is influenced by Phase Composition and Porosity, Adv. Eng. Mater. 13 (2011) 342–350. [20] J. A. Killion, S. Kehoe, L.M. Geever, D.M. Devine, E. Sheehan, D. Boyd, C. L. Higginbotham, Hydrogel/bioactive glass composites for bone regeneration applications: Synthesis and characterisation, Mater. Sci. Eng. C. 33 (2013) 4203–4212. [21] L. Peng, Preparation and Evaluation of Porous Chitosan/Collagen Scaffolds for Periodontal Tissue Engineering, J. Bioact. Compat. Polym. 21 (2006) 207–220. [22] M. Peter, N.S. Binulal, S. V. Nair, N. Selvamurugan, H. Tamura, R. Jayakumar, Novel biodegradable chitosan-gelatin/nano-bioactive glass ceramic composite scaffolds for alveolar bone tissue engineering, Chem. Eng. J. 158 (2010) 353–361. [23] S. Sowmya, P.T.S. Kumar, K.P. Chennazhi, S. V Nair, H. Tamura, R. Jayakumar, Biocompatible β -chitin Hydrogel / Nanobioactive Glass Ceramic Nanocomposite Scaffolds for Periodontal Bone Regeneration, Trends Biomater. Artif. Organs. 25 (2011) 1–11. [24] M.G. Araújo, B. Liljenberg, J. Lindhe, Dynamics of Bio-Oss Collagen incorporation in fresh extraction wounds: an experimental study in the dog, Clin. Oral Implants Res. 21 (2010) 55– 64. [25] P.K. Dutta, J. Duta, V.S. Tripathi, Chitin and Chitosan: Chemistry, properties and applications, J. Sci. Ind. Res. (India). 63 (2004) 20–31. [26] J. Finlay, L. Miller, J. Poupard, A review of the antimicrobial activity of chitosan, J. 23

Antimicrob. Chemother. 52 (2003) 18–23. [27] E. Khor, L.Y. Lim, Implantable applications of chitin and chitosan, Biomaterials 24 (2003) 2339–2349. [28] P. Sriamornsak, Chemistry of Pectin and Its Pharmaceutical Uses : A Review, Silpakorn Univ. J. Soc. Sci. Humanit. Arts. 3 (2003) 206–228. [29] R.K. Mishra, A. K. Banthia, A. B. Majeed, Pectin based formulations for biomedical applications: A review, Asian J. Pharm. Clin. Res. 5 (2012) 1–7. [30] F. Munarin, S.G. Guerreiro, M.A. Grellier, M.C. Tanzi, M.A. Barbosa, P. Petrini, P. L. Granja,

Pectin-based

injectable

biomaterials

for

bone

tissue

engineering.,

Biomacromolecules 12 (2011) 568–77. [31] L. Liu, M.L. Fishman, J. Kost, K.B. Hicks, Pectin-based systems for colon-specific drug delivery via oral route, Biomaterials 24 (2003) 3333–3343. [32] P. Gentile, M. Mattioli-Belmonte, V. Chiono, C. Ferretti, F. Baino, C. Tonda-Turo, C. Vitale-Brovarone, L. Pashkuleva, R.L. Reis, G. Ciardelli, Bioactive glass/polymer composite scaffolds mimicking bone tissue, J. Biomed. Mater. Res. - Part A. 100 A (2012) 2654–2667. [33] P. Gentile, V. Chiono, C. Tonda-Turo, C. Mattu, F. Baino, C. Vitale-Brovarone, G. Ciardelli, Bioresorbable glass effect on the physico-chemical properties of bilayered scaffolds for osteochondral regeneration, Mater. Lett. 89 (2012) 74–76. [34] F. Bigucci, B. Luppi, T. Cerchiara, M. Sorrenti, G. Bettinetti, L. Rodriguez, V. Zecchi, Chitosan/pectin polyelectrolyte complexes: Selection of suitable preparative conditions for colon-specific delivery of vancomycin, Eur. J. Pharm. Sci. 35 (2008) 435–441. [35] S. Koutsopoulos, Synthesis and characterization of hydroxyapatite crystals: a review study 24

on the analytical methods, J. Biomed. Mater. Res. 62 (2002) 600–12. [36] S.S. Rashidova, R.Y. Milusheva, L.N. Semenova, M.Y. Mukhamedjanova, N.L. Voropaeva, S. Vasilyeva, R. Faizieva, I.N. Ruban, Characteristics of Interactions in the Pectin-Chitosan System, Chromatographia 59 (2004) 779–782. [37] G. Morris, S. Kök, S. Harding, G. Adams, Polysaccharide drug delivery systems based on pectin and chitosan, Biotechnol. Genet. Eng. Rev. 27 (2010) 257–284. [38] P. Coimbra, P. Ferreira, H.C. de Sousa, P. Batista, M.A. Rodrigues, I.J. Correia, M.H. Gil, Preparation and chemical and biological characterization of a pectin/chitosan polyelectrolyte complex scaffold for possible bone tissue engineering applications, Int. J. Biol. Macromol. 48 (2011) 112–8. [39] E. Tsuruga, H. Takita, H. Itoh, Y. Wakisaka, Y. Kuboki, Pore size of porous hydroxyapatite as the cell-substratum controls BMP-induced osteogenesis, J. Biochem. 121 (1997) 317–324. [40] V. Karageorgiou, D. Kaplan, Porosity of 3D biomaterial scaffolds and osteogenesis, Biomaterials 26 (2005) 5474–5491. [41] J.R. Woodard, A.J. Hilldore, S.K. Lan, C.J. Park, A.W. Morgan, J.A.C. Eurell, S.G. Clark, M.B. Wheeler, R.D. Jamison, A.J.W. Johnson, The mechanical properties and osteoconductivity of hydroxyapatite bone scaffolds with multi-scale porosity, Biomaterials 28 (2007) 45–54. [42] A. Scabbia, L. Trombelli, A comparative study on the use of a HA/collagen/chondroitin sulphate biomaterial (Biostite) and a bovine-derived HA xenograft (Bio-Oss) in the treatment of deep intra-osseous defects, J. Clin. Periodontol. 31 (2004) 348–55. [43] T. Matsuno, T. Nakamura, K. Kuremoto, S. Notazawa, T. Nakahara, Y. Hashimoto, T. Satoh, Y. Shimizu, Development of beta-tricalcium phosphate/collagen sponge composite for bone 25

regeneration, Dent. Mater. J. 25 (2006) 138–44. [44] F. Baino, C. Vitale-Brovarone, Three-dimensional glass-derived scaffolds for bone tissue engineering: current trends and forecasts for the future, J. Biomed. Mater. Res. A. 97 (2011) 514–35. [45] H.H.K. Xu, C.G. Simon, Fast setting calcium phosphate-chitosan scaffold: Mechanical properties and biocompatibility, Biomaterials 26 (2005) 1337–1348. [46] A.K. Gaharwar, C. Rivera, C.J. Wu, B.K. Chan, G. Schmidt, Photocrosslinked nanocomposite hydrogels from PEG and silica nanospheres: Structural, mechanical and cell adhesion characteristics, Mater. Sci. Eng. C. 33 (2013) 1800–1807. [47] D. a. Wahl, J.T. Czernuszka, Collagen-hydroxyapatite composites for hard tissue repair, Eur. Cells Mater. 11 (2006) 43–56. [48] C. Vitale-Brovarone, G. Ciapetti, E. Leonardi, N. Baldini, O. Bretcanu, E. Verné, F. Baino, Resorbable glass-ceramic phosphate-based scaffolds for bone tissue engineering: synthesis, properties, and in vitro effects on human marrow stromal cells, J. Biomater. Appl. 26 (2011) 465–89. [49] X. Liu, M.N. Rahaman, G.E. Hilmas, B.S. Bal, Mechanical properties of bioactive glass (1393) scaffolds fabricated by robotic deposition for structural bone repair, Acta Biomater. 9 (2013) 7025–34. [50] H.-W. Kim, J.C. Knowles, H.-E. Kim, Hydroxyapatite porous scaffold engineered with biological polymer hybrid coating for antibiotic Vancomycin release, J. Mater. Sci. Mater. Med. 16 (2005) 189–95. [51] A. Ghaffari, K. Navaee, M. Oskoui, K. Bayati, M. Rafiee-Tehrani, Preparation and characterization of free mixed-film of pectin/chitosan/Eudragit RS intended for sigmoidal 26

drug delivery, Eur. J. Pharm. Biopharm. 67 (2007) 175–86. [52] Z. Sheikh, C. Sima, M. Glogauer, Bone Replacement Materials and Techniques Used for Achieving Vertical Alveolar Bone Augmentation, Materials (Basel) 8 (2015) 2953–2993. [53] P.-H. Chen, T.-Y. Kuo, J.-Y. Kuo, Y.-P. Tseng, D.-M. Wang, J.-Y. Lai, H.-J. Hsieh, Novel chitosan–pectin composite membranes with enhanced strength, hydrophilicity and controllable disintegration, Carbohydr. Polym. 82 (2010) 1236–1242. [54] S. V. Dorozhkin, Biphasic, triphasic and multiphasic calcium orthophosphates, Acta Biomater. 8 (2012) 963–977. [55] N.L. Wang, H. Kiyohara, T. Matsumoto, H. Otsuka, M. Hirano, H. Yamada, Polyclonal antibody against a complement-activating pectin from the roots of Angelica acutiloba, Planta Med. 60 (1994) 425–429. [56] M.H. Sakurai, T. Matsumoto, H. Kiyohara, H. Yamada, B-cell proliferation activity of pectic polysaccharide from a medicinal herb, the roots of Bupleurum falcatum L. and its structural requirement, Immunology 97 (1999) 540–547. [57] H. Salman, M. Bergman, M. Djaldetti, J. Orlin, H. Bessler, Citrus pectin affects cytokine production by human peripheral blood mononuclear cells, Biomed. Pharmacother. = Biomédecine Pharmacothérapie 62 (2008) 579–82.

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FIGURE CAPTIONS: Figure 1. Chemical characterization: (a) ATR-IR spectra of pectin and chitosan powders, HA/βTCP ceramic particles, PC hydrogels, PCC composite scaffolds; (b) ATR-IR spectra focus between 2000 cm-1 and 500 cm-1, where the shift of the amine band to 1557 cm-1 and the bands belonging to the inorganic phase (950 cm-1 and 1140 cm-1) are clearly visible. Figure 2. SEM investigation of PC and PCC scaffolds. Highly interconnected pores are shown for PC scaffold; the incorporation of ceramic particles reduces the porosity and pore size of the material. Figure 3. Non-destructive tomographical analysis: (a) µCT images reveal uniform distribution of the ceramic particles in the pectin_chitosan PEI matrix, thus confirming the results from SEM investigation; (b) Pore size distribution of PC and PCC scaffold, the addition of ceramic scaffold reduces the pore size. (c) 2-D development of PC and PCC scaffolds, obtained by placing the analysis planes [xy] and [yz] in the mid-length, and moving the [xz] plane from the bottom to the middle and until the top of the length of the scaffold; 3-D reconstruction of the central (approximately mid-length) cross-section along the [xy], [xz] and [yz] orthogonal planes is also shown. Figure 4. Compressive test: (a) Stress–strain curves for PC and PCC scaffolds in as prepared and hydrated conditions. (b) loading and unloading curves for PC and PCC samples subjected to cyclic strain (100 cycles) in as prepared and fully hydrated conditions; (c) energy absorbed and recovery percentage (monitored every 10 cycles) for PC and PCC samples (the incorporation of calcium phosphate particles increases the strength of PCC scaffold, but does not reduce its recovery ability which is similar to that of PC material). The data are represented as mean ± standard deviation (n=3).

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Figure 5. Swelling properties: (a) swelling kinetics of PC and PCC scaffolds tested in different pH conditions (2.5, 5.5 and 7.4) over 24 h; (b) summary of the swelling degrees at 24 h; (c) optical images show the swollen sample at different pH, and (d) proof of the easy mouldability of composite scaffolds, along with the high hydration properties (water colored with few drops of red ink was used for this qualitative hydration test) and the capacity to recover the initial shape after a compressive stress. The results are reported as mean ± standard deviation (n=3), (*p<0.05, **p<0.01, ***p<0.001, ****p<0.0001, two-way ANOVA with Tukey’s multiple comparison test). Figure 6. Degradation properties at different pH: (a) degradation behavior of the PC and PCC materials over 1 week in different pH conditions; (b) summary of mass losses after soaking for 1 week; (c) optical images of PC and PCC scaffolds during the degradation test. The results are reported as mean ± standard deviation (n=3), (*p<0.05, **p<0.01, ***p<0.001, ****p<0.0001, twoway ANOVA with Tukey’s multiple comparison test). Figure 7. Biological tests with SAOS-2 cells: adhesion and proliferation of osteoblast-like cells were investigated by SEM analysis on (a) PC and (b) PCC scaffolds after 72 h and 2 weeks of culture. HA/βTCP particles promote adhesion and proliferation of SAOS 2 cells after 72 h and a confluent layer of cells was detected after 14 days. Figure 8. ALP gene expression after 1 week of osteoblast-like SAOS-2 cell culture. The results are reported as mean ± standard deviation (n=4), (*p<0.05, **p<0.01, Student’s t-test). Figure 9. Inflammatory response of murine macrophage cells surrounding the scaffolds (gene expression of IL1β, IL6 and IL10 of the cells grown in contact with PC and PCC materials). The results are reported as mean ± standard deviation (n=4) (*p<0.05, two-way ANOVA with Tukey’s multiple comparison test).

29

TABLES AS PREPARED

SAMPLES

Compressive modulus [kPa]

Stress40%

HYDRATED

Toughness

Compressive modulus

3

[kPa]

[kJ/m ]

[kPa]

Toughness Stress40% [kJ/m3] [kPa]

PC

1005 ± 250

127 ± 6

34.3 ± 4.0

33 ± 8

13 ± 3

2.4 ± 0.5

PCC

2559 ± 596

1525 ± 61

222.8 ± 1.9

65 ± 9

121 ± 27

12.0 ± 2.1

Table 1. Compressive modulus, collapse stress (at 40% strain) and toughness for PC and PCC materials in both dry and hydrated conditions.

30

Figure1

Figure 2

Figure 3

Figure 4

Figure 5

Figure 6

Figure 7

Figure 8

Figure 9

Statement of Significance The positive fate of a surgical procedure involving the insertion of a titanium screw still depends on the quality and quantity of alveolar bone which is present in the extraction site. Available materials are basically hard scaffold materials with un-predictable behavior in different condition and difficult shaping properties. In this work we developed a novel pectin-chitosan hydrogel reinforced with ceramic particles. Polysaccharides simulates the Extracellular Matrix of natural bone and the extensive in vitro cells culture study allow to assess that the incorporation of the ceramic particles promote a pro-osteogenic response. Shape control, easy adaption of the extraction site, predictable behavior in different environment condition, swelling properties and an anti-inflammatory response are the significant characteristics of the developed biomaterial.