Carbohydrate Polymers 126 (2015) 130–140

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Biomineralized biomimetic organic/inorganic hybrid hydrogels based on hyaluronic acid and poloxamer Hyun Wook Huh a , Linlin Zhao b , So Yeon Kim b,c,∗ a b c

College of Pharmacy, Chungnam National University, Daejeon 305-764, South Korea Graduate School of Energy Science and Technology, Chungnam National University, Daejeon 305-764, South Korea Department of Chemical Engineering Education, College of Education, Chungnam National University, Daejeon 305-764, South Korea

a r t i c l e

i n f o

Article history: Received 19 December 2014 Received in revised form 9 March 2015 Accepted 16 March 2015 Available online 23 March 2015 Keywords: Hydrogel Biomineralization Hyaluronic acid Poloxamer

a b s t r a c t A biomineralized hydrogel system containing hyaluronic acid (HA) and poloxamer composed of a poly(ethylene oxide)/poly(propylene oxide)/poly(ethylene oxide) (PEO–PPO–PEO) block copolymer was developed as a biomimetic thermo-responsive injectable hydrogel system for bone regeneration. Using HA and poloxamer macromers with polymerizable residues, organic/inorganic HA/poloxamer hydrogels with various compositions were prepared and subjected to a biomineralization process to mimic the bone extracellular matrix. An increase in HA content within the hydrogels enhanced intermolecular chelation with calcium ions, leading to an increase in nucleation and growth of calcium phosphate in the hydrogels. After the biomineralization procedure, a crystalline formation was observed within and on the surface of the hydrogel. All of the HA/poloxamer hydrogel samples exhibited relatively high water content of greater than 90% at 25 ◦ C, and the water content was influenced by the HA/poloxamer composition, biomineralization, and temperature. In particular, the HA/poloxamer hydrogel was injectable through a syringe without demonstrating appreciable macroscopic fracture at room temperature, whereas it was more opaque and adopted a more rigid structure as the temperature increased because of the increasing hydrophobicity of poloxamer. The enzymatic degradation behavior of the hydrogels depended on the concentration of hyaluronidase, HA/poloxamer composition, and biomineralization. The release kinetics of model drugs from HA/poloxamer hydrogels was primarily dependent on the drug loading content, water content, biomineralization of the hydrogels, and ionic properties of the drug. These results indicate that biomineralized HA/poloxamer hydrogel is a promising candidate material for a biomimetic hydrogel system that promotes bone tissue repair and regeneration via local delivery of drugs. © 2015 Elsevier Ltd. All rights reserved.

1. Introduction Natural bone is a composite biomaterial that functions as a connective tissue and is responsible for the mechanical support of muscular activity. Its extracellular matrix (ECM) has a unique composition of calcium phosphate minerals, primarily in the form of hydroxyapatite, that provide strength and support to the bone tissue (Bose, Roy, & Bandyopadhyay, 2012; Holzwarth & Ma, 2011; Hutmacher, 2000). In recent years, significant progress has been made in the development of synthetic bone graft materials, including bioceramics, biopolymers, and composite materials, that share the benefits of being amenable to different designs and

∗ Corresponding author at: Department of Chemical Engineering Education, College of Education, Chungnam National University, Daejeon 305-764, South Korea. Tel.: +82 42 821 5892; fax: +82 42 821 8864. E-mail address: [email protected] (S.Y. Kim). http://dx.doi.org/10.1016/j.carbpol.2015.03.033 0144-8617/© 2015 Elsevier Ltd. All rights reserved.

modifications to meet specific needs (Bose et al., 2012; Holzwarth & Ma, 2011; Hutmacher, 2000; Liu, Lim, & Teoh, 2013; Wang et al., 2007). In particular, a new development in biomaterials is the synthesis of calcium phosphate in polymer matrices to produce biomimetic composites that can initiate osteogenesis when implanted in bony sites (Dogan & Oner, 2006; Nuttelman, Benoit, Tripodi, & Anseth, 2006; Song, Saiz, & Bertozzi, 2003). Biomineralized polymeric matrices prepared by the bioinspired precipitation of inorganic minerals from aqueous solution at near-ambient or physiological conditions have emerged as promising tools to uncover the complex roles of the cellular microenvironment in regulating the diverse phenotypic activities of stem and progenitor cells related to the development, repair, regeneration, and remodeling of bone tissue (Abdelkebir et al., 2012; Dogan & Oner, 2006; Furuzono, Taguchi, Kishida, Akashi, & Tamada, 2000; Ma & Feng, 2011; Nuttelman et al., 2006; Song et al., 2003; Taguchi, Muraoka, Matsuyama, Kishida, & Akashi, 2001; Yokoi, Kawashita, Kikuta, & Ohtsuki, 2010). Most previous studies have focused on apatite

H.W. Huh et al. / Carbohydrate Polymers 126 (2015) 130–140

formation on gels, scaffolds, fibers, or grafted films (Abdelkebir et al., 2012; Dogan & Oner, 2006; Furuzono et al., 2000; Nuttelman et al., 2006; Song et al., 2003; Taguchi et al., 2001). Several studies have shown that the role of chemical motifs can be specifically modulated by other matrix variables, such as pore size and the density of hydrophobic motifs, for mineralization of a three-dimensional matrix (Bleek & Taubert, 2013; Phadke, Zhang, Hwang, Vecchio, & Varghese, 2010; Sasaki, Yataki, & Maeda, 1998). However, a biomineralized hydrogel system for local delivery of drugs and bone regeneration has not yet been systematically examined. Recently, there has been growing interest in injectable in situ gel-forming polymeric hydrogel systems for tissue engineering applications (Balakrishnan, Joshi, Jayakrishnan, & Banerjee, 2014; Lee, Chung, & Kurisawa, 2009; Matanovic, Kristl, & Grabnar, 2014; Ni et al., 2014; Radhakrishnan, Krishnan, & Sethuraman, 2014). From a clinical perspective, injectable scaffolds are extremely desirable because they permit minimally invasive delivery of the construct and thereby reduce surgical risks (Matanovic et al., 2014; Radhakrishnan et al., 2014). The ideal injectable polymer scaffold for tissue engineering would satisfy the following criteria: it should be pliable enough at room temperature to allow the incorporation of cells and macromolecules, be amenable to minimally invasive implantation procedures, demonstrate in situ stabilization upon placement in the body, improve the mechanical integrity of the matrix to better support tissue or organ formation, be minimally toxic, and be capable of interacting with the host tissue and biological environment on a molecular level (Kim, Chung, Gilbert, & Healy, 2005; Kim & Healy, 2003; Matanovic et al., 2014; Radhakrishnan et al., 2014). The main objective of this study was to prepare a biomimetic thermo-responsive injectable hydrogel system to improve bone regeneration by local delivery of protein drugs including bone morphogenetic proteins. HA is a glycosaminoglycan component present in the ECM and synovial fluid of joints. It provides soft connective tissues with a counteracting force to resist compression by absorbing water and hence is the major component of the ECM in load-bearing joints (Collins & Birkinshaw, 2013; Leach & Schmidt, 2005; Price, Berry, & Navsaria, 2007). It has been reported that acidic ECM proteins that are attached to collagen scaffolds play important template and inhibitory roles during the mineralization process (Dogan & Oner, 2006; Nuttelman et al., 2006; Song et al., 2003). Thus, the acidic groups of HA serve as binding sites for calcium ions and align them in an orientation that matches the apatite crystal lattice (Amosi et al., 2012; Dogan & Oner, 2006; Song et al., 2003). Poloxamer, a non-toxic PEO–PPO–PEO triblock copolymer, plays an important role in in situ gelling systems (Elluru, Ma, Hadjiargyrou, Hsiao, & Chu, 2013; Niu et al., 2009). Through a sol–gel transition, an aqueous solution of poloxamer at a low temperature becomes a hydrogel at a higher temperature. This reversible transition has led to the application of poloxamer in some pharmaceutical fields, including fertility control, pain management, and controlled drug delivery through the ophthalmic or rectal route. PEO–PPO–PEO copolymers can exist in aqueous solutions in the form of single chains, micelles, or physical gels as a result of their amphiphilic properties. The block copolymer molecules can self-assemble into micelles in aqueous solutions at concentrations greater than the critical micelle concentration and at temperatures greater than the critical micelle temperature (Elluru et al., 2013; Niu et al., 2009). In addition, at higher polymer concentrations the close packing of micelles results in the formation of gel-like ordered structures (Elluru et al., 2013; Niu et al., 2009). In this study, we focus on a biomineralized organic/inorganic hybrid hydrogel system that assembles a matrix that is chemically and morphologically similar to natural bone matrix and has injectable and drug-release properties suitable for enhancing

131

bone tissue regeneration. We designed hydrogels composed of HA, a biocompatible natural polymer with acidic side groups, and poloxamer, a synthetic polymer with thermo-responsive sol–gel transition properties. HA and poloxamer macromers with polymerizable residues were synthesized. Crosslinked HA/poloxamer hydrogels were then prepared and mineralized using a ureamediated method. To evaluate the feasibility of biomineralized HA/poloxamer hydrogels as a thermo-responsive injectable system, we characterized their morphology, water content, and enzymatic degradation kinetics. In addition, in vitro release behaviors of model drugs from the HA/poloxamer hydrogel were investigated. 2. Materials and methods 2.1. Materials Hyaluronic acid (HA, Mw ∼ 1.6 × 106 Da), poloxamer (PEO–PPO–PEO; Mw ∼ 5800; 30% PEO), triethylamine (TEA), glycidyl methacrylate (GM), acryloyl chloride, ammonium peroxydisulfate (APS), hydroxyapatite, and N,N,N ,N tetramethylethylenediamine (TEMED) were purchased from Sigma (St. Louis, MO, USA). Urea, hyaluronidase, cefazolin, theophylline, and sodium hydroxide were obtained from Aldrich (Milwaukee, WI, USA). Hydrogen chloride was purchased from Junsei Chemical Co., Ltd. (Tokyo, Japan). Dulbecco’s phosphatebuffered saline (PBS) was purchased from GIBCO BRL (Grand Island, NY, USA). Distilled and deionized water was prepared using a Milli-Q Plus System (Millipore, Bedford, MA, USA). All other chemicals used were reagent grade and were used as purchased without further purification. 2.2. Methods 2.2.1. Synthesis of HA and poloxamer macromers HA and poloxamer macromers with polymerizable residues were synthesized (Niu et al., 2009; Oudshoorn, Rissmann, Bouwstra, & Hennink, 2007). Briefly, 1 g of HA was dissolved in 100 ml distilled water. Next, 2.2 ml triethylamine and 2.2 ml glycidyl methacrylate were added separately, thoroughly mixed for 1 h at 60 ◦ C, and the mixture was stirred overnight at room temperature. The reaction solution was precipitated in a 20-fold volumetric excess of acetone and dissolved twice in distilled water to remove excess reactants. The methacrylated HA macromer solution was lyophilized and stored desiccated at 4 ◦ C. Poloxamer macromer was synthesized as follows. Poloxamer (1.72 mmol) was dissolved in 180 ml THF, and then triethylamine (6.88 mmol) and acryloyl chloride (6.88 mmol) were added. After stirring for 6 h at 75 ◦ C, the reaction mixture was filtered to remove triethylamine-hydrochloride and the supernatant was precipitated in an excess of n-hexane. After precipitation, poloxamer macromer containing vinyl groups that were polymerizable through diacrylation of the end groups was obtained and dried at ambient temperature under vacuum for 24 h. The chemical structures of HA and poloxamer macromers were analyzed using 400 MHz 1 H nuclear magnetic resonance (NMR; JNM-AL400, JEOL, Tokyo, Japan) and Fourier transform infrared spectroscopy (FT-IR; FT/IR-460 PLUS, JASCO, Tokyo, Japan). 2.2.2. Preparation of HA/poloxamer hydrogels Crosslinked polymer networks composed of HA and poloxamer were synthesized as shown in Fig. 1. Different HA/poloxamer hydrogels were prepared by varying the weight ratio of HA and poloxamer macromers in the feed as described in Table 1. The HA:poloxamer feed weight ratios were adjusted to 1.0:5.0, 1.5:4.5,

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Fig. 1. Synthetic scheme for HA/poloxamer hydrogels. (A) Hyaluronic acid (HA), (B) poloxamer, (C) HA macromer, (D) poloxamer macromer, (E) HA network, (F) poloxamer network, and (G) HA/poloxamer hydrogel.

and 2.0:4.0. Dry nitrogen gas was bubbled through 0.45 g of a mixture of methacrylated HA and poloxamer in distilled water for 15 min to remove dissolved oxygen. After the nitrogen gas purge, 0.8 wt% (based on total polymer) of APS and 8.0 v/wt% (based on total polymer) of TEMED were added as the initiator and accelerator, respectively. The reaction mixture was stirred vigorously for 30 s and then held at room temperature for 24 h under white fluorescent light in a glass vial. After the crosslinking reaction, the resulting hydrogel products were washed three times for 3–5 min each in excess ultrapure water to remove unreacted compounds.

2.2.3. Biomineralization of HA/poloxamer hydrogels Biomineralization of the HA/poloxamer hydrogel was performed using a urea-mediated method to improve the interaction between the hydrogel and osteoblasts (Song et al., 2003). To prepare the mineralization solution, hydroxyapatite (2.92 mmol) was suspended in 100 ml deionized water with stirring, and 2 M HCl was added until the hydroxyapatite suspension was completely dissolved at a final pH of 2.5–3.0. Urea (0.2 mol) was dissolved in the solution to a concentration of 2 M. Each of the HA/poloxamer hydrogel samples (1.5 cm × 1.5 cm × 2 mm) was placed in 50 ml of

H.W. Huh et al. / Carbohydrate Polymers 126 (2015) 130–140 Table 1 Compositions of HA/poloxamer and biomineralized HA/poloxamer hydrogels. No. 1 2 3 4 5 6 a b

Hydrogel sample HA/poloxamer-1.0/5.0 HA/poloxamer-1.5/4.5 HA/poloxamer-2.0/4.0 B-HA/poloxamer-1.0/5.0b B-HA/poloxamer-1.5/4.5b B-HA/poloxamer-2.0/4.0b

remaining weight (%) of HA/poloxamer hydrogels was calculated using the following equation:

HA:poloxamer weight ratio in feeda

Biomineralization

1.0:5.0 1.5:4.5 2.0:4.0 1.0:5.0 1.5:4.5 2.0:4.0

– – – √ √ √

remaining weight (%) =

the acidic HA-urea stock solution. The mineralization solution was slowly heated to 90–95 ◦ C over 2 h at an average heating rate of 0.6 ◦ C/min without agitation and maintained at that temperature overnight. The final pH was approximately 8.0. After the reaction, mineralized HA/poloxamer hydrogel samples were repeatedly rinsed in ultrapure water to remove loosely attached minerals and soluble ions. 2.2.4. Characterization of HA/poloxamer and biomineralized HA/poloxamer hydrogels The chemical structure of the HA/poloxamer hydrogels before and after biomineralization was analyzed by FT-IR measurement. FT-IR spectra were recorded on a FT/IR-460 PLUS spectrometer (JASCO) between 4000 and 650 cm−1 , with a resolution of 2 cm−1 and 64 scans. The HA/poloxamer and biomineralized HA/poloxamer hydrogel samples were freeze-dried overnight, weighed upon removal from the freeze-dryer, and immersed in PBS for 24 h. The water content was calculated on the basis of the weight difference of the hydrogel samples before and after swelling: (Ws − Wd ) × 100, Ws

W  t

W0

× 100,

(2)

where W0 is the initial weight of the hydrogel, and Wt is the swollen weight of the hydrogel at a certain time point. Three specimens were tested for each sample.

All hydrogels were synthesized in distilled water as a reaction solvent. Biomineralized HA/poloxamer hydrogel samples.

water content (%) =

133

(1)

where Ws is the weight of the swollen gel, and Wd is the weight of the dry gel. The morphology of the HA/poloxamer and biomineralized HA/poloxamer hydrogels was examined using field emission scanning electron microscopy (FE-SEM) (JSM-7000F, JEOL, Tokyo, Japan) at 3 kV. The degree of biomineralization of the HA/poloxamer hydrogel was determined by thermal gravimetric analysis (TGA) using a Metter Toledo TGA/SDTA 851 (Columbus, OH, USA). Approximately 20 mg of dried hydrogel sample was placed in an alumina sample pan for TGA analysis and heated from room temperature to 1100 ◦ C at a rate of 5 ◦ C/min under nitrogen atmosphere. The degree of biomineralization of hydrogel was determined as the remaining weight percentage after heating (n = 3). 2.2.5. Enzymatic degradation kinetics of HA/poloxamer hydrogels Enzymatic degradation kinetics of HA/poloxamer hydrogels were investigated by monitoring the mass loss of the hydrogel samples as a function of time of exposure to an enzyme solution (Kim et al., 2005; Kim & Healy, 2003). The degradation behavior of the HA/poloxamer hydrogels according to the concentration of hyaluronidase solution and the composition of hydrogel was determined over 10 days. Each HA/poloxamer hydrogel sample (2.0 × 5.0 × 5.0 mm3 ) was placed in PBS (pH 7.4) at 37 ◦ C for 24 h, and the initial weight of the swollen samples was measured. The hydrogels were then placed in PBS containing hyaluronidase, 0.2 mg/ml sodium azide, and 1 mM CaCl2 and shaken gently in a shaking water bath at 37 ◦ C. The mass loss of the hydrogel samples was tracked over time relative to the initial swollen weight. The

2.2.6. In vitro drug release of HA/poloxamer and biomineralized HA/poloxamer hydrogels For the in vitro drug release test, we used cefazoline as an ionic drug and theophylline as a nonionic drug. These two model drugs were loaded into each HA/poloxamer hydrogel sample using the swelling-loading technique. Briefly, dried HA/poloxamer samples were soaked in each aqueous drug solution for two days at 25 ◦ C and allowed to swell to an equilibrium state in order to achieve a high loading content in the hydrogel. The fully swollen hydrogel samples were removed from the drug solution, blotted with filter paper to eliminate surface water, and dried to a constant weight. The amount of drug loaded into the HA/poloxamer hydrogel was determined from the weight difference of hydrogel between the fully swollen state in drug aqueous solution and the completely dry state. Dried, drug-loaded HA/poloxamer hydrogels were placed in 25-ml sealed vials containing PBS (0.1 M, pH 7.4). All release studies were conducted in a shaking water bath at 37 ± 0.5 ◦ C with gentle shaking at 50 rpm. At predetermined time intervals, 3-ml aliquots of aqueous solution were withdrawn from the release medium and replaced with the same volume of fresh buffer solution. The medium was analyzed using UV–vis spectrophotometry (Shimadzu Model UV 2101PC, Kyoto, Japan), and the UV absorbance of cefazoline and theophylline was measured at 274 and 270 nm, respectively. The cumulative amount of released drug was determined using the standard calibration curve, and all release experiments were carried out in triplicate. 3. Results and discussion 3.1. Synthesis of HA and poloxamer macromers An HA macromer with polymerizable residues was synthesized through a methacrylation reaction, as shown in Fig. 1(C). Fig. 2(A) shows the 1 H NMR spectra of native HA that served as a control (a) and the HA macromer with methacrylated functional groups (b). The 1 H NMR spectrum confirmed the substitution of methacrylate groups on the HA backbone, showing peaks at approximately 1.9 ppm and 3.0–4.4 ppm corresponding to the methyl protons and the protons on the ring of original HA, respectively, as well as two distinctive peaks at 5.6 and 6.0 ppm that were attributed to the two protons of the double bond of the methacrylate group and a peak at approximately 1.8 ppm ascribed to the methyl group adjacent to the double bond, which were not present in the original HA. The FT-IR spectra of HA and the HA macromer are shown in Fig. 2(C). HA exhibited a characteristic peak in the region of 3419 cm−1 due to hydrogen bonded O H stretching and N H stretching vibration of the N-acetyl side chain. The peaks at 1637 and 1415 cm−1 were assigned to the amide I group of C O carboxyl and aromatic amine CN stretching, respectively. A peak at 2911 cm−1 caused by the methyl C H stretch corresponding to glucuronic acid and a peak at 1036 cm−1 for the primary alcohol C O stretch were also observed (Fig. 2(C)(a)). Compared with the original HA, methacrylated HA showed a sharp peak at ∼1556 cm−1 (C C) and a new peak at ∼1459 cm−1 (CH2 ), representing the double bond of the methacrylate group (Fig. 2(C)(b)). These results confirmed that HA was derivatized with methacrylated groups.

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(A)

(C)

(B)

(D)

Fig. 2. (A) 1 H NMR spectra of (a) HA and (b) HA macromer; (B) 1 H NMR spectra of (a) poloxamer and (b) poloxamer macromer; (C) FT-IR spectra of (a) HA and (b) HA macromer; (D) FT-IR spectra of (a) poloxamer and (b) poloxamer macromer.

In addition, poloxamer macromer was synthesized by reaction between the terminal hydroxyl group of poloxamer and acryloyl chloride (Fig. 1(D)). Fig. 2(B) shows the 1 H NMR spectra of poloxamer and poloxamer macromer. Comparison of the peak areas of the acryl protons at 6.0–6.5 ppm and the methyl protons from propylene oxide at 1.2 ppm demonstrated that approximately 95.5 mol% hydroxy groups reacted with acryloyl chloride. The reaction was also confirmed from the FT-IR spectrum of poloxamer macromer, in which the peak for stretching vibration in the acrylate group appeared at 1727 cm−1 , as shown in (b) of Fig. 2(D). These results indicate the introduction of vinyl groups that could be further polymerized. 3.2. Characterization of biomineralized HA/poloxamer hydrogels As shown in Table 1, HA/poloxamer hydrogels were prepared by varying the feed ratio of HA and poloxamer. The synthesis of HA/poloxamer hydrogels was confirmed by FT-IR spectroscopy. Fig. 3(A) shows FT-IR spectra of (a) poloxamer macromer, (b) HA macromer, (c) HA/poloxamer-2.0/4.0, (d) HA/poloxamer1.5/4.5, and (e) HA/poloxamer-1.0/5.0 hydrogels. The HA macromer

with methacrylated side groups showed a characteristic peak at ∼1556 cm−1 (C C) that can be attributed to the double bond of the methacrylate group (Fig. 3(A)(b)). The peak for stretching vibration in the acrylate group at 1727 cm−1 was observed in the FT-IR spectrum of poloxamer macromer (Fig. 3(A)(a)). However, these peaks were reduced in the FT-IR spectra of HA/poloxamer hydrogels (Fig. 3(A); (c) HA/poloxamer-2.0/4.0, (d) HA/poloxamer-1.5/4.5, and (e) HA/poloxamer-1.0/5.0). In addition, the FT-IR spectra of HA/poloxamer hydrogels included characteristic absorptions of both HA and poloxamer, clearly indicating the successful preparation of HA/poloxamer hydrogels. After biomineralization, the physical appearance of the HA/poloxamer hydrogels changed from transparent to white. The FT-IR spectra of the untreated HA/poloxamer hydrogels and the biomineralized HA/poloxamer hydrogels are compared in Fig. 3(B). After biomineralization, a strong band appeared at 1024 cm−1 (Fig. 3(B); (e) B-HA/poloxamer-2.0/4.0, (f) B-HA/poloxamer-1.5/4.5, and (g) B-HA/poloxamer-1.0/5.0). This band has been attributed to phosphate stretching vibrations. As shown in Fig. 3(B)(a), hydroxyapatite can be identified on FT-IR by its characteristic bands at 633 cm−1 corresponding to OH vibrations and at 961–1106 cm−1

H.W. Huh et al. / Carbohydrate Polymers 126 (2015) 130–140

135

(A)

(B) Fig. 3. (A) FT-IR spectra of HA/poloxamer hydrogels: (a) poloxamer macromer, (b) HA macromer, (c) HA/poloxamer-2.0/4.0, (d) HA/poloxamer-1.5/4.5, and (e) HA/poloxamer1.0/5.0; (B) FT-IR spectra of biomineralized HA/poloxamer hydrogels: (a) hydroxyapatite, (b) HA/poloxamer-2.0/4.0, (c) HA/poloxamer-1.5/4.5, (d) HA/poloxamer-1.0/5.0, (e) B-HA/poloxamer-2.0/4.0, (f) B-HA/poloxamer-1.5/4.5, and (g) B-HA/poloxamer-1.0/5.0.

resulting from phosphate stretching vibrations. These results demonstrate the presence of hydroxyapatite in the HA/poloxamer hydrogel. The cross-sectional morphologies of HA/poloxamer hydrogels before and after biomineralization were observed by FE-SEM. Images of the initial HA/poloxamer hydrogels (HA/poloxamer2.0/4.0) show surfaces free of any foreign particles or formations (Fig. 4(A) and (B)). After the biomineralization procedure, a crystalline formation was observed within and on the surfaces

of biomineralized HA/poloxamer hydrogels (B-HA/poloxamer2.0/4.0), as shown in Fig. 4(C)–(F). To quantitatively determine the incorporation of hydroxyapatite within the biomineralized HA/poloxamer hydrogels, we performed TGA measurements. Biomineralized HA/poloxamer hydrogel samples were completely dried for 48 h, and TGA measurements were performed from room temperature to 1100 ◦ C. The remaining weight on the TGA curve was exclusively ascribed to inorganic substances, i.e., hydroxyapatite incorporated within

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Fig. 4. FE-SEM images of (A) HA/poloxamer-2.0/4.0 (scale bar = 1.00 mm), (B) HA/poloxamer-2.0/4.0 (scale bar = 50.0 ␮m), (C) B-HA/poloxamer-2.0/4.0 (scale bar = 1.00 mm), (D) B-HA/poloxamer-2.0/4.0 (scale bar = 100.0 ␮m), (E) B-HA/poloxamer-2.0/4.0 (scale bar = 50.0 ␮m), and (F) B-HA/poloxamer-2.0/4.0 (scale bar = 10.0 ␮m).

the HA/poloxamer hydrogel. Typical TGA curves are shown in Fig. 5. The weight (%) of the residue, the inorganic component of the biomineralized HA/poloxamer hydrogel, gradually increased with increasing content of HA in the hydrogels. From the results of the TGA measurements of biomineralized HA/poloxamer hydrogels, the percentage of incorporated hydroxyapatite within B-HA/poloxamer-1.0/5.0, B-HA/poloxamer-1.5/4.5, and B-HA/poloxamer-2.0/4.0 was 11.3, 12.6, and 14.3%, respectively. This indicated that an increase in HA content in the HA/poloxamer hydrogel enhanced intermolecular chelation with calcium ions, leading to an increase in nucleation and growth of calcium phosphate in the HA/poloxamer hydrogel (Amosi et al., 2012; Li et al., 2014; Liu et al., 2014; Ma & Feng, 2011). These results were in good agreement with the results of FT-IR measurement. 3.3. Water content of biomineralized HA/poloxamer hydrogels The swelling kinetics of HA/poloxamer hydrogels in PBS are plotted in Fig. 6. All of the hydrogels swelled rapidly, reaching

equilibrium within 24 h. The water content of HA/poloxamer and biomineralized HA/poloxamer hydrogels in PBS is described in Table 2. All of the HA/poloxamer hydrogel samples exhibited a relatively high water content of >90% at 25 ◦ C regardless of composition. The water content gradually increased as the amount of HA (pKa = 3.0) in the hydrogel increased. HA contains an ionizable group, COO− . The presence of ionized COO− groups within the HA/poloxamer hydrogel causes repulsion between them, resulting in an increase in free volume in the polymeric matrix and thus an increase in swelling of the hydrogel. The water content of biomineralized HA/poloxamer hydrogels was lower than that of untreated HA/poloxamer hydrogels (Table 2). This suggests that the nucleation and growth of calcium phosphate in the hydrogels during biomineralization inhibits interaction between ionic groups in the hydrogels and water molecules, leading to a reduction in water content (Li et al., 2014; Liu et al., 2014). HA/poloxamer hydrogels were injectable through a syringe without demonstrating appreciable macroscopic fracture at room temperature (Fig. 6). When the temperature was increased from

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Fig. 5. TGA curves of (a) B-HA/poloxamer-1.0/5.0, (b) B-HA/poloxamer-1.5/4.5, and (c) B-HA/poloxamer-2.0/4.0.

content was enhanced by increasing poloxamer content in the HA/poloxamer hydrogels. These results suggest that in situ stabilization of the HA/poloxamer hydrogels at body temperature (37 ◦ C) may support controlled release of drugs.

3.4. Enzymatic degradation behavior of biomineralized HA/poloxamer hydrogels The enzymatic degradation of HA/poloxamer and biomineralized HA/poloxamer hydrogels was investigated by monitoring mass loss of the hydrogel samples as a function of time of exposure to hyaluronidase solution (Fig. 7). The degradation of hydrogel in solution is generally associated with several network parameters such as the number of cross-links per backbone chain, the number of vinyl groups on the cross-linking molecule, molecular weight of the backbone, and the proportion of degradable groups in the main and side chain (Drury & Mooney, 2003). We therefore determined the in vitro degradation kinetics of the HA/poloxamer hydrogels with respect to the HA/poloxamer composition within the hydrogel and the concentration of hyaluronidase solution (1–100 U/ml). Fig. 7(A) shows the degradation profiles of HA/poloxamer and biomineralized HA/poloxamer hydrogels prepared using different feed ratios of HA/poloxamer (as shown in Table 1) in 100 U/ml hyaluronidase solution. The degradation rate of HA/poloxamer hydrogel gradually increased with increasing content of biodegradable natural polymer, HA, in the hydrogels. In addition, the degradation of biomineralized HA/poloxamer hydrogels was

Fig. 6. Swelling kinetics of HA/poloxamer hydrogels according toHA/poloxamer composition and injectability of HA/poloxamer hydrogel at room temperature.

25 to 50 ◦ C, the HA/poloxamer hydrogel became more opaque and adopted a more rigid structure as a result of the increase in hydrophobicity of poloxamer composed of PEO–PPO–PEO block copolymers. The effect of temperature on the water content of HA/poloxamer hydrogels was also determined. As the temperature increased, the water content gradually decreased, as shown in Table 2. In addition, the temperature-dependent decrease in water

Table 2 Water content and drug loading content of HA/poloxamer and biomineralized HA/poloxamer hydrogels. No.

Hydrogel sample

Water contenta ◦

25 C 1 2 3 4 5 6

HA/poloxamer-1.0/5.0 HA/poloxamer-1.5/4.5 HA/poloxamer-2.0/4.0 B-HA/poloxamer-1.0/5.0 B-HA/poloxamer-1.5/4.5 B-HA/poloxamer-2.0/4.0

90.2 92.6 93.4 83.2 85.5 86.8

DLC (%)b ◦

◦

37 C ± ± ± ± ± ±

1.1 0.6 0.6 0.7 0.7 0.5

87.8 91.5 92.8 81.9 84.5 86.5

50 C ± ± ± ± ± ±

0.9 0.7 0.6 0.9 0.7 0.5

86.5 90.3 92.2 79.8 83.6 86.0

Cefazoline ± ± ± ± ± ±

1.0 0.6 0.8 1.0 0.6 0.1

16.0 21.9 23.5 11.4 13.7 16.1

± ± ± ± ± ±

0.2 0.7 0.3 0.2 0.3 0.1

Theophylline 18.1 23.1 24.8 12.5 14.9 16.7

± ± ± ± ± ±

0.5 0.2 0.5 0.4 0.5 0.7

a Water content of hydrogel = (Ws − Wd )/Ws × 100, where Ws is weight of swollen gel, and Wd is weight of dry gel. Each water content value represents the average of four samples. b Drug loading content (DLC) = (mass of drug/mass of drug − loaded hydrogel) × 100 = [drug/(drug + polymer)] × 100.

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(A)

(B)

(C)

Fig. 7. Enzymatic degradation of HA/poloxamer hydrogels. (A) Degradation of HA/poloxamer and biomineralized HA/poloxamer hydrogels in 100 U/ml hyaluronidase solution; (B) degradation of HA/poloxamer with respect to the concentration of hyaluronidase solution; (C) degradation of biomineralized HA/poloxamer hydrogels with respect to the concentration of hyaluronidase solution. Each point represents the mean ± S.D. of three samples.

slower than that of HA/poloxamer hydrogels; the biomineralized B-HA/poloxamer-2.0/4.0 showed 64.4% degradation whereas HA/poloxamer-2.0/4.0 showed 80.7% degradation after 10 days in 100 U/ml hyaluronidase solution (Fig. 7(A)). This indicates that the nucleation and growth of calcium phosphate in the hydrogels during biomineralization inhibit effective interaction between the hydrogel network and enzyme molecules, leading to a reduction in degradation rate. Furthermore, HA/poloxamer and biomineralized hydrogels exhibited significant weight loss in a hyaluronidase concentrationdependent manner (Fig. 7(B) and (C)). As the concentration of hyaluronidase solution increased from 1 to 100 U/ml, the degradation rate of HA/poloxamer-2.0/4.0 gradually increased. 3.5. Drug loading and release from biomineralized HA/poloxamer hydrogels The drug loading content (DLC) of cefazoline and theophylline in HA/poloxamer and biomineralized HA/poloxamer hydrogels with various compositions is described in Table 2. The DLC of HA/poloxamer hydrogels gradually increased as the HA content increased and the poloxamer content decreased

(HA/poloxamer-1.0/5.0: 16.0 ± 0.2% cefazoline, 18.1 ± 0.5% theophylline; HA/poloxamer-1.5/4.5: 21.9 ± 0.7% cefazoline, 23.1 ± 0.2% theophylline; HA/poloxamer-2.0/4.0: 23.5 ± 0.3% cefazoline, 24.8 ± 0.5% theophylline). The trend in DLC of HA/poloxamer hydrogels is consistent with the water content, which also increased as the amount of HA (pKa = 3.0) containing ionizable groups ( COO− ) increased. This indicates that repulsion between ionized COO− groups within the HA/poloxamer results in an increase in free volume in the polymeric matrix and a corresponding increase in the loading content of cefazoline and theophylline in the hydrogel (Kono, Oeda, & Nakamura, 2013; Yan, Jin, Gao, & Chen, 2014). Theophylline, a nonionic drug, exhibited a higher DLC than cefazoline, an ionic drug (Table 2). In addition, the DLC of the biomineralized HA/poloxamer hydrogels was lower than that of untreated HA/poloxamer hydrogels (B-HA/poloxamer-1.0/5.0: 11.4 ± 0.2% cefazoline, 12.5 ± 0.4% theophylline; B-HA/poloxamer-1.5/4.5: 13.7 ± 0.3% cefazoline, 14.9 ± 0.5% theophylline; B-HA/poloxamer-2.0/4.0: 16.1 ± 0.1% cefazoline, 16.7 ± 0.7% theophylline). This result might reflect the decrease in water content of the hydrogels after biomineralization. In vitro release tests of the model drugs from HA/poloxamer and biomineralized HA/poloxamer hydrogels were performed

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the biomineralized HA/poloxamer hydrogels with relatively low water content and DLC was much lower than that of the nonbiomineralized hydrogels. Therefore, the release behavior of drug molecules from the HA/poloxamer hydrogel can be controlled by the composition, DLC, biomineralization of the hydrogel, and ionic properties of the drug. 4. Conclusions

(A)

(B) Fig. 8. In vitro release kinetics of cefazoline and theophylline from hydrogels at 37 ◦ C. (A) Drug release behavior from HA/poloxamer hydrogels, (B) drug release behavior from biomineralized HA/poloxamer hydrogels. Each point represents the mean ± S.D. of three samples.

Novel biomineralized HA/poloxamer hydrogels were designed as a biomimetic thermo-responsive injectable hydrogel system to improve bone regeneration through local drug delivery. Crosslinked hydrogels composed of the natural polymer HA with acidic functional groups, which serves as a binding site for calcium ions during the mineralization procedure, and poloxamer, which exhibits temperature-dependent sol–gel transition properties, were synthesized. Biomineralized organic/inorganic hybrid HA/poloxamer hydrogels were prepared using a urea-mediated method. After biomineralization, a crystalline formation was observed within and on the surfaces of hydrogels. HA/poloxamer hydrogels exhibited relatively high water content (>90%) and were injectable through a syringe without demonstrating appreciable macroscopic fracture at room temperature. However, when the temperature increased, HA/poloxamer hydrogels became more opaque, had a more rigid structure, and the water content gradually decreased because of the increasing hydrophobicity of poloxamer. The enzymatic degradation behavior of hydrogels was primarily dependent on the concentration of hyaluronidase, HA/poloxamer composition, and biomineralization. The DLC of the HA/poloxamer hydrogel gradually increased with the amount of HA within the hydrogel. This indicates that repulsion between ionized COO− groups within the HA/poloxamer results in an increase in free volume in the polymeric matrix and a corresponding increase in DLC in the hydrogel. The release behavior of model drugs from the HA/poloxamer hydrogels was significantly influenced by the drug loading content, water content, biomineralization of the hydrogels, and ionic properties of the drug. Our results suggest that the biomineralized HA/poloxamer hydrogel system could be applied to create a thermoresponsive injectable hydrogel system as a novel type of bone filling material that has the capability of controlled drug release for bone repair and regeneration. Acknowledgement

over 10 days. The release kinetics of cefazoline and theophylline from HA/poloxamer hydrogels with various compositions showed sustained release behavior and were influenced by the composition of the hydrogels. As shown in Fig. 8(A) and Table 2, the release rate of drugs from HA/poloxamer hydrogels gradually increased as the water content and DLC of the hydrogels increased (HA/poloxamer-1.0/5.0 < HA/poloxamer1.5/4.5 < HA/poloxamer-2.0/4.0). This suggests that an increase in the driving forces for diffusion of drug molecules, such as the concentration difference and swelling of the hydrogels, influence the drug release behavior (Lin & Metters, 2006). In addition, the release of ionic cefazoline was slower than that of the nonionic drug theophylline. These results indicate that drug release behavior from the HA/poloxamer hydrogel is significantly influenced by the interaction between ionic groups in the hydrogel and ionic drug molecules, as well as by swelling of the hydrogel (Colinet, Dulong, Mocanu, Picton, & Le Cerf, 2009; Park, Seo, & Song, 2013). The release behavior of cefazoline and theophylline from the biomineralized HA/poloxamer hydrogels exhibited a similar trend to that of HA/poloxamer hydrogels (Fig. 8(B)). The release of drugs gradually accelerated as the water content and DLC of hydrogels increased (B-HA/poloxamer-1.0/5.0 < B-HA/poloxamer-1.5/4.5 < BHA/poloxamer-2.0/4.0). As expected, the release rate of drugs from

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