An injectable nano-hydroxyapatite (n-HA)/glycol chitosan (G-CS)/hyaluronic acid (HyA) composite hydrogel for bone tissue engineering

Abstract

The aim of the present study was to fabricate an injectable nano-hydroxyapatite (n-HA)/glycol chitosan (G-CS)/hyaluronic acid (HyA) composite hydrogel and investigate its potent application in bone tissue engineering. The resultant composite hydrogel was thoroughly characterized by Fourier transform infrared spectroscopy (FTIR) and X-ray diffractometry (XRD). The developed composite hydrogels exhibited a porous structure (pore size: 100–350 μm) associated with many n-HA particles and their aggregation inside the hydrogel, as indicated by scanning electron microscopy (SEM). With an increase in HyA concentration, the porosity and swelling ratio of the hydrogels decreased accordingly. The developed composite hydrogels were susceptive to the enzymatic hydrolysis, which exhibited a faster degradation rate in PBS solution containing 2.5 mg ml⁻¹ lysozyme than that of in PBS (pH = 7.4). In vitro cytocompatibility was evaluated by using MC-3T3-E1 cells to confirm that the developed composite hydrogels were cytocompatible, non-toxic and cells were found to be attached and well spread out on the hydrogels after 7 days co-incubation. All these results suggest that the developed composite hydrogels has great potential for bone tissue engineering application.

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1. Introduction

The treatment of bone defects caused by many reasons such as trauma, disease and aging is the most important subject in the orthopedics. Until to now, autografting and allografting are still the first choice in the clinical treatment of bone defects, especially in large bone defects. However, autografting and allografting mainly suffer from shortcomings such as the shortages of donors, the risk of an immune response and transmitted diseases and so on.^1–3 Therefore, more and more attention have been oriented on the development of bone graft substitute (scaffold, hydrogel, membrane and so on) driven from the organic polymeric matrix and inorganic reinforcement.

Since the natural bone is made of nano-sized hydroxyapatite crystals with needle-like morphology, the nano-hydroxyapatite (n-HA) as inorganic reinforcement has been widely used for bone tissue engineering, due to its superior bioactivity and osteoconductivity.^4–7 However, n-HA by itself is unsuitable for bone tissue engineering, because it is very brittle and shows poor mechanical property.^8,9 Thus, more and more researchers are focusing on the development of n-HA/organic polymer hybrids composites.^7,10 In past several decades, a number of biodegradable polymeric matrixes such as poly(lactic acid) (PLA), poly(ε-caprolactone) (PCL), poly(D,L-lactic-glycolic acid) (PLGA), chitosan (CS), hyaluronic acid (HyA), alginate (Alg) and so on, have been widely employed to fabricate various types of composite scaffold/hydrogel via different techniques (freeze-drying, electrospinning, gas foaming) for bone tissue engineering.^11–15 Among these polymeric matrixes, natural-based materials have gained much interest than synthetic polymers due to their great biocompatibility, biodegradability and abundance resource. Chitosan (CS) as a natural biodegradable and biocompatible cationic polysaccharide, has been extensively used in the biomedical field and it has been reported to promote cell adhesion and functional expression of osteoblasts.^16,17 The simple n-HA/CS composite scaffold developed by freeze drying of mixture of n-HA and chitosan acidic aqueous solution resulted in the poor mechanical properties and unsatisfied cytocompatibility due to the positive nature of CS in physiological condition.^4,18 To overcome this inherent shortcoming, many kinds of water-soluble chitosan derivatives including carboxylated chitosan, PEGylated chitosan and so on, were successfully synthesized and physically composited with n-HA to form membrane/scaffold for various bone tissue application.^19–21 However, the poor mechanical property as well as rapid degradation behavior of physical composite scaffold/hydrogel had greatly limited its further in vivo application. In order to improve the mechanical properties and biological performance of composite scaffold/hydrogel, several strategies including photochemical cross-linking, “click” chemistry approaches and so on have been developed to generate the covalent composite scaffold/hydrogel.^22,23 More recently, Dahlmann et al. reported an in situ cross-linkable alginate and hyaluronic acid hydrogel based on the hydrazone reaction for myocardial tissue engineering.²⁴ Such hydrogel exhibited favorable properties such as non-toxic, excellent biocompatibility and biodegradability.

Encouraging by the work of Dahlmann et al.²⁴ in the present study, we attempted to adopt the hyaluronic acid (HyA) aldehydes to covalently cross-link glycol chitosan (G-CS) aqueous solution via hydrazone reaction to develop an injectable nano-hydroxyapatite/glycol chitosan/hyaluronic acid composite hydrogel for bone tissue engineering application. Nano-hydroxyapatite (n-HA) was physically dispersed throughout hydrogel that could improve the connection and permanence of the inorganic phase in scaffold and enhance osteoblast proliferation. We hypothesized that the developed nano-hydroxyapatite/glycol chitosan/hyaluronic acid composite hydrogel with favorable properties such as non-toxic, tunable mechanical properties, excellent biocompatibility and biodegradability, could serve as a bone substitute for bone reconstruction in vivo.

2. Materials and methods

2.1 Materials

Glycol chitosan (G-CS, G7553, degree of polymerization ≥ 400) and sodium periodate were provided by Sigma-Aldrich (USA). Hyaluronic acid (HyA; M_w ∼ 1 000 000) was obtained from Shangdong Freda Co, Ltd (China). Nano-hydroxyapatite (n-HA; ∼20 nm) was purchased from Beijing DK nano technology Co, Ltd (China). Other regents used in this manuscript were all of analytical grade.

2.2 Oxidation of hyaluronic acid to aldehydes and its characterization

Inspiring by the result of previous study,²⁴ HyA aldehydes were synthesized by the oxidation of vicinal diols within the monomeric units of HyA using periodic acid oxidation assay. Briefly, 5 g HyA was dissolved into 200 ml distilled water solution followed by the addition of 1.6 g sodium periodate in 10 ml aqueous solution for reaction 2 h under dark condition. Thereafter, the reaction was quenched by the addition of equivalent amount of glycol. Finally, the resultant mixture was dialyzed against with distilled water for 3 days and lyophilized at −60 °C for 72 h (FD-1D-50; Beijing Boyikang Laboratory Instruments Co., Ltd) to yield HyA aldehydes powder for further application. The obtained HyA aldehydes was characterized by ¹H-NMR and the degree of oxidation was quantified by ¹H-NMR.²⁴

2.3 Preparation of nano-hydroxyapatite/glycol chitosan/hyaluronic acid composite hydrogel

The n-HA/G-CS/HyA composite hydrogel was developed by the following procedure and the original weight ratios of three components were presented in Table 1. Briefly, the calculated weight of glycol chitosan was dissolved into distilled water solution followed by the addition of n-HA powder to form homogeneous slurry with the aid of continuous ultrasound. Thereafter, a certain amount of HyA aldehydes aqueous solution was added into the above slurry to obtain n-HA/G-CS/HyA hydrogels. Finally, the obtained n-HA/G-CS/HyA hydrogels were lyophilized at −60 °C for 72 h to obtain the n-HA/G-CS/HyA scaffold for further characterization.

Table 1 The components of n-HA/G-CS/HyA composite hydrogel and corresponding gelation time

Code	Final n-HA (mg ml^−1)	Final G-CS (mg ml^−1)	Final HyA aldehydes (mg ml^−1)	Gelation time (s)
S1	0.5	5	0.5	105 ± 5
S2	0.5	5	1	60 ± 8
S3	0.5	5	1.5	25 ± 5

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2.4 Characterization

2.4.1 FTIR analysis. The FTIR spectra of various composite scaffolds were recorded by a FTIR instrument (Perkin Elmer, USA) within scanning range of 400–4500 cm⁻¹ by using KBr pellets.

2.4.2 XRD detection. XRD analysis of various composite scaffolds was performed with a X-ray diffractometry (XRD; D8 Advance Davinci, Bruker) by using Cu Kα radiation in the range of 2θ = 5–60°.

2.4.3 SEM observation. The micro-morphology of various composite scaffolds were observed with a scanning electron microscopy (SEM; S-4800, Hitachi). Prior to observation, the composite scaffolds were sputtered with gold for better conductivity.

2.4.4 Porosity test. The porosity of various composite scaffolds were determined by the liquid displacement method as described previously.²⁵ Briefly, the cylindrical scaffolds were generated in a cylindrical mould and the original weight (W₀) as well as volume (V₀) were measured. Thereafter, the scaffolds were immersed into dehydrated alcohol for 24 h until saturation and the samples were weighted again (W₂₄). Finally, the porosity was calculated by the following equation (eqn (1)):ρ represented the density of dehydrated alcohol (ρ = 0.79 g ml⁻¹).

2.4.5 Swelling test. To measure the swelling ratio of hydrogels, lyophilized hydrogel samples were weighted (W₀) and immersed into 20 ml phosphate buffer solution (PBS, pH = 7.4) in a 50 ml BD tube for a period of swelling study. After 24 h of incubation, the hydrated hydrogel samples were carefully removed from the medium and the residual PBS on the surface of hydrogels were cleaned by a filter paper, and weighted (W₂₄). Finally, the swelling ratio of hydrogels was calculated from the following equation (eqn (2)):

Swelling ratio (%) = (W₂₄ − W₀)/W₀ × 100% (2)

2.5 In vitro degradation test

In vitro degradation behavior of various composite hydrogels over time were performed in PBS solution (pH = 7.4) and PBS solution (pH = 7.4) containing 2.5 mg ml⁻¹ lysozyme, respectively. Briefly, the scaffold samples were weighted (W₀) and then immersed into 15 ml PBS solution (pH = 7.4) or PBS solution (pH = 7.4) containing 2.5 mg ml⁻¹ lysozyme for in vitro degradation study at 37 °C. After 1, 2, 3 and 4 weeks of incubation, the samples were withdrawn and washed with distilled water followed by the drying at a vacuum drier overnight and weighting (W_t). The degree of degradation was expressed as the weight loss (%) and calculated by the following equation (eqn (3)):where W₀ was the original weight of samples and W_t represented the weight of sample after incubation for specific time.

2.6 In vitro cytocompatibility test

The attachment and spreading nature of MC-3T3-E1 cells on the n-HA/G-CS/HyA hydrogel were evaluated by using Live/Dead staining assay. Briefly, the HyA aldehydes and G-CS were dissolved separately into sterile phosphate buffer saline (PBS; pH = 7.4) to form 3 mg ml⁻¹ HyA aldehydes aqueous solution and 10 mg ml⁻¹ G-CS aqueous solution, respectively. Thereafter, a certain amount of n-HA powder was added into 10 mg ml⁻¹ G-CS aqueous solution to obtain n-HA/G-CS slurry giving final n-HA concentration at 1 mg ml⁻¹. In situ cross-linking hydrogels were prior formed in the 6-well plates by mixing the equal volume of 3 mg ml⁻¹ HyA aldehydes aqueous solution and n-HA/G-CS slurry to partially cover the bottom of well, followed by the incubation with MC-3T3-E1 cells at density of 1 × 10⁵ cells per well in DMEM medium. After 1, 3, 5 and 7 days of co-incubation, the viability of cells contacted with hydrogels was visualized with a Live/Dead assay and measured by MTT assay. The cells without any treatment was used as reference. For the cells encapsulated into hydrogels, the cells were firstly suspended into n-HA/G-CS slurry, followed by the addition of 3 mg ml⁻¹ HyA aldehydes aqueous solution in equal volume to obtain the cell seeded composite hydrogel in 6-well plates (1 × 10⁵ cells per well). After 1, 3, 5 and 7 days of co-incubation, the viability of cells in the hydrogels was visualized with a Live/Dead assay. The Live/Dead assay was performed as following procedure: the cells or cells seeded composite hydrogel were incubated with Live/Dead working solutions containing 2 μM calcein acetoxymethyl (calcein AM) and 4 μM ethidium homodimer-1 (EthD-1) for 30 minutes in a dark place at room temperature, and then viewed under the fluorescence microscope.

2.7 Statistical analysis

Experimental results were expressed as mean ± SD. Data was analyzed by one-way ANOVA for comparison of any two groups. p < 0.05 was considered as statistical significance.

3. Results and discussion

3.1 Synthesis and characterization of HyA aldehydes

It is well known to us that the sodium periodates could selectively oxidize the vicinal diol group leading to the formation of two aldehydes. As shown in Fig. 1, it is clearly observed that the presence of aldehyde groups was confirmed by the chemical shift at 4.89 ppm, 5.00 ppm and 5.10 ppm, which was in accordance with the previous studies.^21,24 The degree of oxidation was 10.2%, as quantified by comparing the integrals of aldehyde hydrate with the signal of the HyA backbone (2.01 ppm of acetamide for HyA).

Fig. 1 ¹H-NMR spectrum of HyA aldehydes in D₂O.

3.2 Preparation of nano-hydroxyapatite/glycol chitosan/hyaluronic acid (n-HA/G-CS/HyA) hydrogel

The formation of nano-hydroxyapatite/glycol chitosan/hyaluronic acid (n-HA/G-CS/HyA) composite hydrogel was developed by a simple mixing method (Scheme 1). By the addition of HyA aldehydes aqueous solution to glycol chitosan/n-HA slurry resulted into the sol–gel transition of system due to the covalent cross-linking between –NH₂ of glycol chitosan and –CHO of HyA aldehydes. The gelation behavior of system greatly depended on the content of glycol chitosan and HyA aldehydes. By increasing the content of HyA aldehydes, the gelation time decreased accordingly (Table 1). It should be noted that the gelation behavior of system was not thermosensitive, but was time-dependent.

Scheme 1 The formation of nano-hydroxyapatite (n-HA)/glycol-chitosan (G-CS)/hyaluronic acid (HyA) aldehydes hydrogel.

3.3 Characterization

3.3.1 FTIR analysis. In order to evaluate the possible chemical interaction between the three components, the FTIR spectra measurements were performed. As shown in Fig. 2A, the distinctive absorption band at 3571 cm⁻¹ was attributed to the stretching vibration of hydroxyl groups of n-HA, while the absorption bands at 604 cm⁻¹ and 1040 cm⁻¹ were attributed to the phosphate group (–PO₄³⁻).^26,27 Additionally, the bonds of CO₃²⁻ (1422 cm⁻¹ and 1470 cm⁻¹) also appeared in n-HA.²⁸ For blank hydrogel, the typical bands at 1642 cm⁻¹ and 1456 cm⁻¹ that corresponded to amide group and amino group were observed.¹³ Intensity at 1030 cm⁻¹ was attributed to C–O bond stretching. As comparison with spectra of n-HA and blank hydrogel, the nano-hydroxyapatite/glycol chitosan/hyaluronic acid hydrogel exhibited both characteristic bands of n-HA and blank hydrogel, but the absorption at 1642 cm⁻¹ in blank hydrogel (amide I) shifted a little to low wavelengths (1640 cm⁻¹), which might be attributed to the results of possible intra-hydrogen bonds interaction among the three components.²⁸ These results confirmed the effective existence of n-HA in hydrogel samples.

Fig. 2 (A) FTIR spectra of blank hydrogel, n-HA powder, S1, S2 and S3; (B) XRD spectra of blank hydrogel, n-HA powder, S1, S2 and S3.

3.3.2 XRD analysis. Fig. 2B shows the XRD spectra of n-HA, bank hydrogel and composite hydrogels. The main diffraction peaks of n-HA at 2θ = 25.8, 28, 29, 31.8, 32.9, 34.1 and 40° were observed, while the main peak of bank hydrogel was at 2θ = 23.5°. In the case of composite hydrogels, the new peak at 31.8° appeared, corresponding to the specific peak of n-HA.^11,19 Although the intensity of n-HA crystal decreased in the composite hydrogels, the crystalline structure of n-HA was well retained in the composite hydrogels, suggesting that the n-HA was successfully incorporated into hydrogels.

3.3.3 Morphology of hydrogel. The interior morphology of the lyophilized hydrogel samples were characterized by SEM. As shown in Fig. 3, all composite hydrogels samples exhibited interconnected porous structure with pore size in the range of 100–350 μm. With an increase in HyA concentration, the pore size of hydrogel decreased accordingly, which might be induced by the higher cross-linking density of hydrogel.²⁹ More careful observation in high magnification (×1500) showed that there were many n-HA particles and its aggregation on the wall of pore.

Fig. 3 SEM images of the freeze-dried composite hydrogels samples. (A) S1 hydrogel (×300); (A-1) S1 hydrogel (×1500); (B) S2 hydrogel (×300); (B-1) S2 hydrogel (×1500); (C) S3 hydrogel (×300); (C-1) S2 hydrogel (×1500). Red arrow indicated the n-HA.

3.3.4 Porosity test. As an ideal scaffold for tissue engineering application, the developed scaffold should be a porous structure, allowed the nutrients and oxygen to penetrate freely to the inner regions of the scaffold. Meanwhile, the porous structure of scaffold also should provide the suitable positions for cell migration and further proliferation. Fig. 4A showed the porosity of the various hydrogels. The hydrogels from S1, S2 and S3 exhibited the porosity of 71.87 ± 3.74%, 61.52 ± 0.76% and 60.44 ± 1.49%, respectively. The porosity was significantly higher in S1 hydrogel compared with those of S2 and S3 hydrogel (P < 0.05), indicating that the component concentration had potent influence on the porous structure of hydrogel. When the HyA concentration increased from 0.5 mg ml⁻¹ to 1.5 mg ml⁻¹, the cross-linking degree of hydrogel increased accordingly corresponding to the result of SEM observation, yet resulting in the low level of porosity of hydrogels. This result was consistent with the previous studies, which claimed that the porosity of scaffold decreased significantly with an increase of the content of chemical cross-linker.^29,30

Fig. 4 (A) Porosity of various composite hydrogels samples; (B) swelling ratio of various composite hydrogels samples. *P < 0.05 indicated statistically significant difference between S1 and S2, S3 (n = 3). (C) The rate of weight loss of various composite hydrogels samples as a function with time in PBS (pH = 7.4); (D) the rate of weight loss of various composite hydrogels samples as a function with time in 2.5 mg ml⁻¹ lysozyme aqueous solution.

3.3.5 Swelling test. The swelling behavior is also one of important factors to be investigated for composite hydrogel. The swelling ratios of all hydrogels were determined by weighing over time (Fig. 4B). In a period of 72 h swelling study, the swelling equilibrium was achieved rapidly at 24 h. As depicted in Fig. 4B, it was clearly observed that the swelling ratio of S1, S2 and S3 after 24 h of incubation were 5993 ± 100%, 4231 ± 319% and 3745 ± 384%, respectively. As is well known to us, the swelling ratio of hydrogel is mainly modulated by its microstructure and the hydrophilic–hydrophobic property of its component. And it is also influenced by the by the environmental conditions, such as temperature, pH condition and so on.^31,32 The swelling ratios of hydrogels were found to be dependent on the component concentration. More specifically, the swelling ratio of hydrogel decreased with an increase in HyA concentration, which might be attributed to the higher HyA concentration resulted in the more compact microstructure of hydrogel, yet decreasing the available space for water retention during the swollen state.³³ Similar result was also reported by Zhao et al., whom demonstrated that the decreased of pore size of hydrogel would lead to the decrease of swelling ratio.³⁴

3.4 In vitro degradation test

Controllable degradation behavior of scaffold was very important issue for the bone tissue engineering application. The degradation of hydrogels were evaluated by examining the weight loss over time in either PBS (pH = 7.4) or PBS solution containing 2.5 mg ml⁻¹ lysozyme (Fig. 4). As presented in Fig. 4C, all the hydrogels samples exhibited various extents of weight loss in PBS solution (pH = 7.4) during 4 weeks study. It was found that the degradation rate decreased with the order of S1 > S2 > S3. It seems that hydrogels with the lower crosslinking densities resulted in the faster degradation properties.³⁴ At initial 2 weeks, the rapid degradation of hydrogels might be attributed to the leaching of un-crosslinking components (glycol chitosan and HyA aldehydes) in hydrogel.^19,30,35 Thereafter, the degradation rate decreased accordingly, indicating that the developed hydrogels were resistant to the hydrolytic degradation in PBS (pH = 7.4) solution. On the contrary, the degradation rate of all hydrogels accelerated significantly in PBS solution containing 2.5 mg ml⁻¹ lysozyme, being completely dissolved after 4 week of degradation (Fig. 4D). This result indicated that the developed hydrogels were susceptive to the enzymatic hydrolysis and might be suitable for bone tissue engineering application as it was implanted in vivo on the condition of some enzymolysis.³⁵

3.5 In vitro cytocompatibility test

For in vitro cytocompatibility test, the MC-3T3-E1 cells were contacted directly to the composite hydrogel. As shown in Fig. 5, with the time proceeding, the cells contacted with the hydrogels adhered to, spread and grew on the cell culture dish with normal cell morphology. Live/Dead staining showed that the cells viability (green: live cells; red: dead cells) in the presence of hydrogel leachables and the presence of n-HA in composite hydrogel did not influence its cytocompatibility with almost cells were green. The cell viability of hydrogel after 1, 3, 5 and 7 days incubation remained over 90%, as indicated MTT assay (Fig. S1†). Regarding of the encapsulation of cells into hydrogels (Fig. 6), it clearly observed that the majority of cells were green after 1 day incubation, suggesting the covalent crosslinking environment were none toxic to the osteoblasts cells. Recent studies also claimed that a in situ hydrogelation system based on alginate aldehydes and hyaluronic acid had excellent cytocompatibility in presence of viable myocytes.²⁴ After 7 days co-incubation, more green cells in the composite hydrogels as well as an increase in cell aggregation were observed, suggesting that the encapsulated cells could spread and grow well on the hydrogels. All these results indicated that the developed composite hydrogel exhibited satisfied cytocompatibility suitable for bone tissue engineering application.

Fig. 5 Light micrographs of cells contacted with the composite hydrogels after 1, 3, 5 and 7 days incubation and Live/Dead staining of cells contacted with the composite hydrogels after 1, 3, 5 and 7 days incubation. (Green: live cells; red: dead cells).

Fig. 6 Cell viability in composite hydrogel after 1, 3, 5 and 7 day incubation (green: live cell; red: dead cell).

4. Conclusion

In this paper, n-HA/G-CS/HyA composite hydrogels were successfully fabricated via hydrazone crosslinking strategy and then thoroughly characterized by FTIR and XRD. With an increase in HyA concentration, the pore size, porosity and swelling ratio decreased accordingly. The developed composite hydrogels were susceptive to the enzymatic hydrolysis with complete degradation after 4 weeks study, as indicated by in vitro degradation test. In vitro cytocompatibility test showed that the developed composite hydrogels were non-cytotoxic and cells were found to be attached and spread well on the hydrogels after 7 days co-incubation. All these results indicated that the developed n-HA/G-CS/HyA composite hydrogels might be a promising candidate for bone tissue engineering applications.

Acknowledgements

The research was supported by the National Natural Science Youth Foundation of China (Grant No. 81101395 & No. 81401871).

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Footnote

† Electronic supplementary information (ESI) available. See DOI: 10.1039/c5ra26160k

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