1 Introduction

Metals and their alloys have wide applications in dentistry and medicine as a component of artificial implants [1, 2]. Titanium (Ti) and its alloys are widely used in orthopaedic and oral implants [3, 4]. However, uncoated Ti implants are not as effective at promoting the growth of bone-forming cells as are bioactive bioceramic-coated Ti implants [5, 6]. Furthermore, coatings with bioactive, osteoinductive, and antibacterial properties have been made to improve the surface characteristics of such implants [7–11].

There are numerous reports demonstrating the surface roughness of implants affects the rate of osseointegration [12]. Various methods have been developed in order to create a rough surface and improve the osseointegration of Ti dental implants [13]. Metal implants have been coated with layers of calcium phosphates mainly composed of hydroxyapatite. Following implantation, the release of calcium phosphate into the peri-implant region increases the saturation of body fluids and precipitates a biological apatite onto the surface of the implant. This layer of biological apatite might contain endogenous proteins and serve as a matrix for osteogenic cell attachment and growth [14, 15].

Bioactive glasses have been widely studied for their specific clinical applications due to their ability to promote bond-tissue formation at their surface and to bond with the surrounding living tissues [16–18]. Bonding to living bone seems to develop by a carbonate hydroxyapatite (CHA) layer formed on the surface of the bioactive glass [19, 20]. A very important characteristic of BGs is that they show a genetic control of osteoblasts. Ionic products of BG dissolution also increase proliferation of human osteoblasts and induce insulin-like growth factor II mRNA expression and protein synthesis [21]. Certain BGs have been shown to exhibit antibacterial activity. This may result from one of the surface reactions undergone by these glasses [22]. Particulate Bioglass® is one example to exert a considerable antibacterial effect on certain oral bacteria, including those associated with caries and periodontal diseases [23]. Recently, BG with the commercial name S53P4 and some other BGs such as sol–gel bioactive glasses have shown similar properties on a variety of clinically important aerobic pathogens [24]. The antibacterial action of a BG is affected by its chemical composition and the dissolution conditions in its surroundings. It has been shown that the antibacterial effect of BG 45S5 could be greatly enhanced by lowering its particle size and, as a consequence, by immediate release of alkaline species [25]. Our previous study showed that 58S (57.72 SiO2, 35.09 CaO, and 7.1 P2O5 by weight) and 63S (62.17 SiO2, 28.47 CaO, and 9.25 P2O5 by weight) BG nanoparticles have antibacterial activities. Especially, compared to the 63S BG nanoparticles, the 58S BG nanoparticles showed a stronger bactericidal effect on the studied pathogens with a lower minimum bactericidal concentration (MBC). In fact, the antibacterial activity of BGs has been suggested to be based on several factors including high pH levels and osmotic effects caused by the nonphysiological concentration of ions dissolved from glass. pH measurements revealed that the broth containing the 58S BG nanoparticles had higher pH levels as high as nine as compared with the 63S and 72S BG nanoparticles, which is the threshold concentration inducing antibacterial activity. The synergic effects of high calcium concentration and alkaline pH level may make the broth containing 58S a good antibacterial agent [26].

Synthetic biodegradable polyesters like polyglycolide, polylactide, and their copolyesters Poly (lactide-co-glycolide) (PLGA) have been widely utilized as scaffolds in tissue engineering [27]. These materials have many advantages such as being biodegradable and biocompatible while they are also easily processed into expected configurations [28]. Hollinger showed the osteogenic potential of PLGA [29]. Furthermore, it has been proposed that the use of PLGA in the composite coating provides a locking mechanism between the coating and the juxtaposed bone with time [30]. This study aimed at preparation and in vitro and in vivo evaluation of novel bioactive, biodegradable and antibacterial nanocomposite coating for the improvement of stem cells attachment and antibacterial activity as a candidate for dental implant applications.

2 Experimental procedure

2.1 58S BG nanoparticles synthesis

Starting materials used in this preparation were of analytical grade tetraethyl orthosilicate (TEOS), triethyl phosphate, Ca (NO3)2·4H2O (Aldrich), ethanol, and hydrochloric acid (Merck). All materials were used intact without further purification. The composition of studied bioactive glass belongs to the SiO2–CaO–P2O5 system with 58S (57.72 SiO2, 35.09 CaO, and 7.1 P2O5 by weight) composition. Bioactive glass nanopowders were prepared using the sol–gel technique. Ethanol was used as a dispersant to obtain the nanopowders. The chosen volume ratio of ethanol to TEOS was two. Proper amounts of deionized water (15 ml), 2 N hydrochloric acid (2.5 ml), and TEOS (20.49 ml) were dissolved in ethanol and stirred at room temperature for 30 min. Triethyl phosphate (2.08 ml) was then dissolved into the acid silica sol. After 20 min of stirring, Ca (NO3)2·4H2O (13 g) was added. The solution was stirred for an hour longer. The reaction mixture was transferred to a large Teflon container, which was then placed in an oven for aging at 60°C for 54 h. The aged gel was transferred into another Teflon vessel which was placed inside an especially designed drying chamber (cylindrical steel chamber capped with an aluminium foil with holes in it) with a proper amount of water. In the next step, the whole chamber was placed in an oven at 130°C for 52 h. Finally, the dried gel nanopowders were calcined at 600°C for 2 h.

2.2 Preparation of the nanocomposite coating

The solvent casting process was applied to coat the substrates. Commercially pure Ti (Grade 2), was cut into pieces 20 × 10 mm in size and used as the substrate. These substrates were ultrasonically cleaned first in acetone for 20 min, then in a 70% ethanol solution for 20 min, and finally in distilled water for 15 min. HA nanoparticles (Aldrich,USA) and the 58S BG nanoparticles were dispersed ultrasonically in 20 ml of chloroform (Merck, Germany) and added to PLGA (0.1 g/ml in CHCl3) (RESOMER® RG 502H, PLGA; 50/50wt% poly (lactic acid)/poly(glycolic acid); inherent viscosity 0.20 dl/g [25°C; 0.1 in CHCl3]) with stirring. Solutions with three ratios of the components (PLGA: nanoparticles 90:10, 85:15, and 80:20 by weight percent) were prepared to coat the specimens by dip-coating procedure. The dip-coating process was repeated three times for each sample. The substrates were soaked in the solutions at a speed of 2 cm/min. After each dip-coating step, samples were dried for 1 min and the procedure was repeated.

In the nanocomposite coatings equal amounts of HA and BG nanoparticles were used. The solvent was then allowed to evaporate at the room temperature (21°C) for 24 h. Scanning electron microscopy (SEM, Phillips XL 30) was used to study the microstructure and morphology of the coatings and X-ray diffraction (XRD, Philips X’ Pert-MPD System with Cu kα wavelength of 1.5418 Ǻ) technique was utilized to determine the composition of the coatings.

2.3 In vitro bioactivity evaluation of the nanocomposite coating

The assessment of in vitro bioactivity was carried out by soaking the samples in simulated body fluid (SBF) in sterilized polyethylene containers maintained at 37°C. The SBF experiment protocol was presented elsewhere [31, 32]. The samples were soaked in SBF for two different periods (7 and 14 days). In order to provide more favourable conditions for apatite deposition, the solution was renewed every 2–3 days.

Hence, the exchange of SBF leads to better simulation of the in vivo conditions, making the assay more precise and reliable. Next, the samples were removed from the solution, rinsed gently, first with pure ethanol, then with deionized water. Finally, they were dried at the room temperature for 3 h. The formation and growth of apatite layer on the samples were verified by SEM and XRD.

2.4 The coatings’ degradation studies

In order to study the effect of BG and HA nanoparticles on the degradation of the nanocomposite coating, in vitro degradation tests were carried out in phosphate buffer saline (PBS; pH 7.4) at 37°C. Each sample was placed in a vessel containing 50 ml of PBS and incubated for periods up to 60 days. The ratio of the sample mass to the PBS solution volume was chosen 6 mg/ml. The pH of the PBS solution was monitored every 2 days. The medium was changed every week. At each time interval (14, 30, and 60 days), samples were removed from the solution, washed with distilled water and air-dried overnight. Changes in the surface morphologies of the coatings during in vitro degradation were evaluated by SEM. In order to evaluate the effect of nanoparticles on the degradation of nanocomposite coating, the degradation of PLGA coating was investigated, simultaneously.

2.5 Attachment and viability of hASCs on the prepared nanocomposite coating

Human adipose-derived stem cells (hASCs) were extracted from subcutaneous adipose tissue of a 35 years old healthy female patient with purity greater than 95%. Protocols were approved by the institutional ethical review board of Isfahan University of Medical Sciences, Iran. Tissues were harvested, as resected fat from lipectomy biopsies. The harvested tissues were processed according to standard stromal-cell extraction protocols [33, 34]. Adipose tissues were finely minced and then digested with 0.075% collagenase type I (Sigma Chemicals, St-Louis, MO) in DMEM F12-HAM medium supplemented with 10% FBS, 1% penicillin/streptomycin for 30 min at 37°C. The collagenase I was neutralized by adding an equal volume of control medium containing Dulbecco’s modified Eagle’s medium (DMEM, GibcoBRL, Carlsbad, CA), 10% fetal bovine serum (FBS, Sijiqing, Hangzhou, China), 1% penicillin/streptomycin. Floating adipocytes were discarded and the cells from the stromal–vascular fraction were pelleted, rinsed with media and centrifuged. The cell suspension was seeded in culture flasks and cultured at 37°C, 5% CO2. Culture medium was changed every 2 days. Having reached confluence, the cells were passaged. Cells at the third passage were used for the experiments. Samples were sterilized by exposure to UV light for 2 h and it was followed by their overnight soaking in 70% ethanol according to standard techniques [35]. The hASCs were enzymatically harvested from the flasks, counted, and seeded onto the sterile samples. In order to determine the cells’ attachment and viability, samples were located in a six-multiwell plate and then seeded at a cell density of 2.0 × 104 cells/cm2, in a complete culture medium. The uncoated Ti (20 × 10 mm) and the polystyrene standard culture plate were used as controls. On every sample (20 × 10 mm), 100 μl of a cell suspension was applied using great care and the cells were allowed to attach for 2 h to the underlying substrate, then 3 ml of culture medium was added carefully as well. At the end of the incubation periods (24, 48, and 72 h), samples were rinsed three times with PBS in order to remove non-adherent cells. The viability of the hASCs still adhering to the substrates were determined by MTT (3-(4,5-dimethylthiazol-2-yl)-2, 5-diphenyl tetrazolium bromide) assay. The MTT was prepared as a 2 mg/ml stock solution in PBS, sterilized by Millipore filtration, and kept in the dark. The MTT solution was added to each well. After 3 h incubation at 37°C, the supernatant was removed, and dimethyl sulfoxide (DMSO) was added to each well to dissolve the formazan crystals under shaking. The solutions were collected and put in 96 well plates. The optical density of the formazan solution was read on an ELISA plate reader (Hyperion, Florida, USA) at 540 nm. The absorbance was proportional to the number of viable cells attached to the substrates. The results from four individual experiments were averaged and reported as optical density values. Morphology of the cells incubated for 3 days was observed using SEM. Cells adhering to the substrates were washed with PBS. Subsequently, the cells were fixed with 2.5% glutaraldehyde in PBS for 1 h at 4°C. The specimens, after being thoroughly washed with PBS, were dehydrated using graded ethanol changes, gold splattered in vacuum, and examined using SEM. The results were statistically analyzed using ANOVA. The post hoc multiple comparisons were performed using Tukey’s test. A “P”-value < 0.05 was considered statistically significant.

2.6 PBGHA nanocomposite coating adhesion to K-wires

Ten commercially available K-wires (Synthes, Switzerland) made of stainless steel with a diameter of 0.8 mm served as substrate specimens and were coated with PBGHA nanocomposite coating with an optimum amount of nanoparticles. Total coating mass was determined with an electronic micro-balance (Sartorius AG, Göttingen, Germany, readability 0.01 mg). Then, without previous drilling, K-wires were incorporated proximally into rabbit tibiae (white mature male New Zealand rabbits.) as intramedullary rods and they were immediately explanted. After explantation and removal of adherent bone and bone marrow, the loss of coating mass (denoted LCM) was determined gravimetrically.

3 Results

3.1 Characterization of the nanocomposite coating

Figure 1 shows the SEM micrographs of the prepared nanocomposite coating. The PLGA was not able to surround the nanoparticles at concentrations above 10 wt% (15 and 20 wt%) in the dry film. At 10 wt%, a uniform nanoparticle distribution was observed along with some nanoparticle aggregates. The Cross-section image of the prepared coating which contains 10 wt% of the nanoparticles showed an interface with no defects (Fig. 1d). Figure 2 shows SEM micrographs of PBGHA coating after sterilization. Nanoparticles were exposed on the surface after sterilization because the PLGA was decomposed during sterilization in ethanol. Furthermore, this coating exhibited macropores after sterilization. XRD was used to provide clues to phases that existed in the nanocomposite coating (Fig. 3). XRD pattern of the prepared coating showed the peaks belonging to Ti and HA and an amorphous pattern is related to the PLGA and BG.

Fig. 1
figure 1

Scanning electron microscopy micrographs of the PBGHA nanocomposite coating: a 10 wt%, b 15 wt%, c 20 wt% nanoparticles, and d cross-section of the nanocomposite coating (10 wt% nanoparticles)

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Fig. 2
figure 2

Scanning electron microscopy micrographs of the PBGHA coating after sterilization. Nanoparticles were exposed on the surface and the coating exhibited macropores

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Fig. 3
figure 3

X-ray diffraction pattern of the prepared nanocomposite coating (10 wt% nanoparticles). The peaks belong to both Ti plus HA. An amorphous pattern belongs to the PLGA and BG

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3.2 In vitro bioactivity evaluation

Scanning electron microscopy analysis shows the effect of PBGHA nanocomposite coating on HA deposition during increasing immersion times in SBF. It was observed that PBGHA nanocomposite coating was able to nucleate more HA formation on its surface compared to Ti substrate. After 7 days, apatite deposits were observed and detected (by SEM and XRD) on the nanocomposite coating, but was not detected on the Ti substrates. After soaked in SBF for 14 days, HA deposits covered all the surface of PBGHA nanocomposite coating (Fig. 4a). However, at this time, Ti substrate showed less bioactivity and some HA deposits were observed on its surface (Fig. 4b). XRD patterns of the nanocomposite coating surface and Ti substrates before and after 14 days immersion in SBF are shown in Fig. 5a, b. As it was observed, after 14 days immersion in SBF, the diffraction peaks of apatite deposits formed on the surface of the prepared nanocomposite coating were sharp (Fig. 5a). But, the diffraction peaks of apatite deposits on Ti substrates were very weak (Fig. 5b).

Fig. 4
figure 4

Scanning electron microscopy micrographs of the PBGHA nanocomposite coating (10 wt% nanoparticles) (a), and the titanium substrate (b) after 14 days immersion in SBF

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Fig. 5
figure 5

X-ray diffraction of the PBGHA nanocomposite coating (10 wt% nanoparticles) (a), and titanium substrate (b) before and after immersion in SBF for 14 days

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3.3 Degradation studies

Figure 6 shows morphological changes of the PBGHA and PLGA coatings before and after degradation in PBS. Results showed that the morphology of the sterilised coatings before degradation was porous (Fig. 6a, d). After degrading for 14 days, more pores could be seen on the surface of the coatings (Fig. 6b, e). Pores were formed by PLGA degradation. During degradation, the dissolution of nanoparticles and their aggregates could create pores of similar dimension to them and their aggregates and some of the nanoparticles were exposed on the surface (Fig. 6c). Nanoparticles aggregates dissolution made pores with greater dimensions than pores formed by pure PLGA degradation. After 30 days, the PLGA coating showed more pore formation (Fig. 6f), and it was observed that it swelled in some parts and lost its adhesion to the substrate. PBGHA nanocomposite coating considerably degraded after about 60 days immersion in PBS (Fig. 7).

Fig. 6
figure 6

Scanning electron microscopy micrographs of the PBGHA nanocomposite coating (10 wt% nanoparticles) (a, b, c), and PLGA coating (d, e, f) before and after degradation for 14 days (b, e), and 30 days (c, f). Bar = 50 μm

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Fig. 7
figure 7

Scanning electron microscopy micrographs of the PBGHA nanocomposite coating (10 wt% nanoparticles) after degradation for 60 days in PBS. The coating considerably degraded after about 60 days immersion in PBS. Remains of the PBGHA coating was observed on the titanium substrate

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3.4 The effect of the nanocomposite coating on the attachment, viability, and proliferation of the hASCs

Figure 8a shows MTT assay results at 24, 48, and 72 h of incubation. At 24 h significant statistical difference was observed between the PBGHA nanocomposite coating and the control (P < 0.05). By 48 and 72 h, the results obtained for the PBGHA nanocomposite coating were significantly different from those obtained for the Ti substrates and the control (P < 0.05). The hASCs showed greater attachment proliferation and viability on the nanocomposite coating than that of Ti substrates and control. Figure 8b shows morphology of the cells adhered to the nanocomposte coating after 3 days incubation. The nanocomposite coating supported the attachment of hASCs.

Fig. 8
figure 8

a MTT assay results for the PBGHA nanocomposite coating (10 wt% nanoparticles). The absorbance values are proportional to the number of viable cells. Difference from titanium + P < 0.05, and difference from control ‡ P < 0.05, b scanning electron micrograph of the confluent layer of hASCs growing on the surface of the PBGHA nanocomposite coating (10 wt% nanoparticles) at 3 days of culturing (scale bar represents 50 μm)

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3.5 PBGHA nanocomposite coating adhesion to K-wires

After implantation, LCM averaged 3.88 ± 0.21% for PBGHA nanocomposite-coated K-wires. Thus, about 96% of the nanocomposite coating mass remained attached to the K-wires during intramedullary implantation. No adhesive failure between the coating and the substrate was observed, and LCM was mainly due to abrasion with cohesive failure within the nanocomposite coating itself.

4 Discussion

Events leading to the integration of an implant into bone, and hence determining the performance of the device, take place largely at the tissue-implant interface. Development of this interface is complex and involves numerous factors. These include not only implant-related factors, such as the material, shape, topography, and surface chemistry, but also the mechanical loading, surgical technique, and patient variables, such as bone quantity and quality. A goal of current implantology research is the design of devices that induce controlled, guided, and rapid healing [36]. One strategy to improve implant integration is the modification of well-established biomaterials by means of surface coating [37]. For example, Chang et al. [38] showed direct bonding between a HA coating on a Ti substrate and the newly formed bone, whereas non-coated Ti did not. Sato et al. [30] showed enhanced osteoblast adhesion on hydrothermally treated HA/titania/PLGA sol–gel Ti coatings.

In the present study, the PBGHA nanocomposite coating was prepared on the Ti substrates. Our hypothesis was suggested that the presence of the 58S BG and hydroxyapatite nanoparticles could be effective in enhancing the biological behaviour of the surface. The results from our previous study revealed that the 58S BG nanoparticles (35.09 wt% CaO) exhibits antibacterial activity even at concentrations below those commonly used in clinical applications [26]. pH measurements showed that the broth containing 58S reached a pH level as high as nine, which is the threshold concentration inducing antibacterial activity. The cytotoxicity of the 58S BG nanoparticles was evaluated. Over the past few years, substantial efforts have begun in the study of the toxic effects of nanomaterials on the environment and living systems. The easiest simple solution in assessing nanomaterial toxicology is utilizing various mammalian cells to test for viability or increase/decrease in a designated inherent biological pathway against chosen engineered nanomaterials. For instance, tetrazolium salt based assays [e.g., MTT (Mosmann 1983)] are readily available commercially and straightforward for use [39]. Therefore, MTT assay was used to evaluate the cytotoxity of the 58S bioactive glass nanoparticles. By 72 h, the cells were exposed to the 58S BG nanoparticles extracts showed significantly greater viability and proliferation than the control did. It was concluded that the extract of the 58S BG nanoparticles could induce cell proliferation. This was in agreement with previous studies that have reported the ionic products of BG dissolution increase proliferation of human osteoblasts [21]. Furthermore, the releasing of silicilic acid by the 58S BG nanoparticles in culture medium causes its alkalinization, due to ion exchange [26]. This increased silicon concentration could increase collagen production as an important parameter in bone repair. It has been shown that fibroblasts, a cell type that has the same embryologic origin of osteoblasts, internalize silicon without detectable cell damage. It was also demonstrated that increased silicon concentration increased collagen production [40, 41].

In the nanocomposite coating preparation, both the 58S BG and HA nanoparticles were used to achieve their advantages, simultaneously. The presence of the BG and HA nanoparticles could have a synergistic effect for early cell capture. The results showed that nanoparticles content in the nanocomposite coating could not exceed 10 wt% due to the formation of non-uniform coating on the substrates. For nanoparticle contents of 15 wt%, and especially 20 wt%, many particle aggregates were observed throughout the specimen. The phase separation phenomenon resulting from this aggregation would then induce early failure at the interface and thus cause the mechanical properties of the composites to deteriorate [42, 43]. Nanoparticles were exposed on the surface after the coating sterilization process because the PLGA decomposed during sterilization in ethanol. Furthermore, these coatings exhibited macropores after sterilization. The coating showed a porous structure that may provide a suitable surface topography for cell adhesion. In fact, open porous structure of void spaces is capable of accommodating tissue ingrowth and it has proposed that the micrometer level pore size is essential for cellular adaptation, sufficient nutrient permeation, and new bone formation [44].

Poly (lactide-co-glycolide)/bioactive glass/hydroxyapatite biodegradable nanocomposite coating could be effective especially at the early stages to induce bone apposition and problems such as non-desirable adhesion to the implant, associated with coatings such as plasma-sprayed HA in which coating could be solved [45]. In fact, a biodegradable coating will degrade as new bone grows and some time after surgery it will degrade entirely. In degradation studies it was supposed that nanoparticles dissolution could form voids in the PBGHA nanocomposite coating. Furthermore, degradation of PLGA could lead to expose more nanoparticles and their aggregates on the surface and consequently, increase the size of the pores in the nanocomposite coating. During degradation, the dissolution of nanoparticles could create pores of similar dimension to the original nanoparticles, which will facilitate oligomer diffusion in the composite. This enhancement in oligomer diffusion through the sample surface, together with the buffering effect of dissolution compounds, could reduce the autocatalysis degradation of PLGA. The swelling of the nanocomposite coating could become more homogeneous without the formation of a surface shell found in PLGA. PBGHA nanocomposite coating considerably degraded after about 60 days of immersion in PBS. Our previous study showed that the 58S BG nanoparticle was effective at buffering; producing a higher Ca2+ ion concentration and pH value in the dissolution medium [26].This buffering effect could dominate the degradation rate of the nanocomposite.

Bioactivity test showed that PBGHA nanocomposite coating provides a suitable surface for bone-like apatite sediments formation. Also, it was showed that after 7 days, samples with pure PLGA coating induced the deposition of bone-like apatite (not shown).

In addition to the differentiated bone cells (osteoblasts, osteoclasts, and osteocytes), bone and adjacent tissues contain a number of less differentiated cells. These undifferentiated cells are of great importance for proper bone healing or anchorage of an implant, since they can be recruited to form osteoprogenitor cells and, with time, develop into differentiated bone cells [46]. An adult stem cell type known as hASCs share many of characteristics of its marrow counterpart including extensive proliferative potential and the ability to undergo multilineage differentiation along classical mesenchymal lineages: adipogenesis, chondrogenesis, osteogenesis, and myogenesis [47]. In this experiment, we used hASCs to evaluate their attachment and viability on the prepared nanocomposite coating containing 10 wt% nanoparticles. The results showed excellent attachment and viability of hASCs on the PBGHA nanocomposite coating. These cells can differentiate into osteoprogenitor cells and, with time develop into differentiated bone cells [46]. Nano-HA on PLGA showed a positive modulation on early osteoblast capture compared to pure PLGA. In fact, nano-HA impeded proliferation, but the cell functionality such as alkaline phosphatase (ALP) and protein expressions were improved. Although nano-HA decreased the cell number, it may have a chemoattractant role for filopodia formation due to selective protein absorptions [48]. Figure 8b showed the positive effect of nanoparticles for filopodia formation and cell capture.

It was reported that the existence of BG particles could enhance cell attachment and spreading. The composites with BG particles were hydrophilic, which could induce a more wettable surface than pure PLLA [49]. The synergetic effects of the 58S BG and HA nanoparticles in the PLGA matrix could enhance early cell capture, proliferation, and differentiation.

Various test methods have been developed to study coating stability. However, these in vitro tests mainly focus on solitary stresses like bending forces and tensile, shear, or compression loads without simulating the in vivo situation, where forces act in a complex manner and combined stresses predominate [10]. ASTM C-633-01, for example, is a standard test to evaluate adhesion and cohesion strength of inorganic coatings [50]. Bonding agents with epoxy resins and organic solvents are applied to fix the substrate coating to the loading fixture. Penetrating bonding agents may interact with biodegradable polymer coatings like PBGHA nanocomposite coating, invalidating mechanical stability or even dissolving the polymer. Stability testing is therefore recommended for every specific application. PBGHA nanocomposite coating proved great stability on K-wires during intramedullary implantation. Abrasion occurred due to flaking of small polymer particles. About 96% of the nanocomposite coating mass remained attached to the K-wires which presented high stability and adhesion on K-wires. Further investigations, especially in vivo tests are being performed in order to guarantee the prepared coating in dental applications.

5 Conclusion

Results showed that using 10 wt% nanoparticles (5 wt% HA and 5 wt% BG) in the nanocomposite coating could provide the desired uniform coating. In addition, the study of in vitro bioactivity showed rapid formation of bone-like apatite on the prepared coating. PBGHA nanocomposite coating considerably degraded after about 60 days of immersion in PBS. The hASCs showed excellent attachment and viability on the PBGHA nanocomposite coating. Furthermore, no adhesive failure between the coating and the substrate was observed, and LCM was mainly due to the abrasion with cohesive failure within the nanocomposite coating itself. Therefore, bioactive PBGHA nanocomposite coating provides an ideal surface for the stem cells attachment, viability, and proliferation and it could induce antibacterial activity, simultaneously.