Injectable hydrogels based on the hyaluronic acid and poly (γ-glutamic acid) for controlled protein delivery

Carbohydrate Polymers 179 (2018) 100–109

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Research paper

Injectable hydrogels based on the hyaluronic acid and poly (γ-glutamic acid) for controlled protein delivery ⁎

Xuebin Maa, Tingting Xub, Wei Chenb, Hongye Qinb, Bo Chib, , Zhiwen Yea, a b

MARK

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School of Chemical Engineering, Nanjing University of Science and Technology, Nanjing, China State Key Laboratory of Materials-Oriented Chemical Engineering, Nanjing Tech University, Nanjing, China

A R T I C L E I N F O

A B S T R A C T

Keywords: Injectable hydrogel Hyaluronic acid Poly (γ-glutamic acid) Protein delivery

Injectable hydrogels have great potential in minimally invasive delivery. In this work, novel injectable hydrogels were prepared via self-crosslinking of aldehyde hyaluronic acid (HA-CHO) and hydrazide-modified poly (γglutamic acid) (γ-PGA-ADH) for proteins delivery. The HA/γ-PGA hydrogels could be formed in situ as fast as 9s with high swelling ratios. Rheological properties illustrated a wide processing range and good mechanical properties, which were reflected by broad linear viscoelastic region and higher threshold shear stress (σc) and storage modulus (G′). Meanwhile, the gelation time, swelling ratio, rheological properties, as well as the protein release behavior could be modulated conveniently. Bovine serum albumin (BSA) was designed as a model drug to study the release behavior. We found that the release mechanisms were either diffusion or Case-II relaxation depending on the different hydrogel components. The HA/γ-PGA hydrogels also showed good biocompatibility. Therefore, the HA/γ-PGA hydrogels have great potential as promising injectable biomaterials for controlled protein delivery.

1. Introduction In recent years, various native and genetically engineered proteins, including enzymes, antibodies, hormones, and cytokines, have been used as biopharmaceuticals. However, protein drugs often have poor bioavailability due to their short half-lives and rapid clearance from the body (Alves et al., 2017; Koyamatsu et al., 2014). In order to achieve therapeutic effects, a common strategy is bolus injection of proteins, but this approach often cause some side effects, such as hematoma, increased postoperative morbidity, risk of tumor formation and so on (Hariawala et al., 1996; James et al., 2016). One of the most effective methods to overcome this problem is pharmaceutical formulation of proteins. Injectable hydrogels are considered as an ideal material for protein delivery, because of remarkable advantages such as easy incorporation of therapeutic drugs via simple mixing, minimally invasive surgical procedure, convenience of filling irregular surgical defects completely, and the tunability properties (Burdick & Murphy, 2012). Previous studies have confirmed that the injectable hydrogels prepared by natural polysaccharides showed good biocompatibility than many synthetic polymers (Matsumura, Nakajima, Sugai, & Hyon, 2014). Hyaluronic acid (HA) is a naturally non-sulfated glycosaminoglycan, and possesses good biocompatibility, biodegradability, non-

⁎

immunogenicity, as well as excellent gel-forming properties. It has gained great attention and interest in drug delivery (Fang, Chen, Leu, & Hu, 2008; Zawko & Schmidt, 2010). Thermosensitive injectable HA hydrogels were concerned by many teams (Fang et al., 2008; Jung, Park, Park, Lee, & Na, 2017) because of its sensitivity to body temperature and superior biocompatibility, but low stability and poor mechanical properties were also observed due to the sensitivity of some non-covalent interactions to physiological environments. Other injectable HA hydrogels have also been prepared by various methods, such as UV crosslinking (Leach, Bivens, Collins, & Schmidt, 2004), “click” chemistry (Hu, Li, Zhou, & Gao, 2011) and enzymatic crosslinking (Kim et al., 2011). However, some potential cytotoxic molecules, such as photosensitizers, catalytic agents and oxidizing agents, which may cause protein degeneration during the crosslinking process, are inevitably introduced. Michael addition reaction is considered to be a suitable method to prepare injectable HA hydrogels. The resultant hydrogels exhibited good biocompatibility, whereas, the crosslinking reaction was too slow (∼30 min or longer) (Hahn, Oh, Miyamoto, & Shimobouji, 2006). In our work, the injectable hydrogels were prepared via Schiff base reactions because of the prominent advantages such as avoidance of crosslinking agents and easy control of reaction rate.

Corresponding author at: State Key Laboratory of Materials-Oriented Chemical Engineering, Nanjing Tech University, No. 30 South Puzhu Road, Nanjing 211800, Jiangsu, China. Corresponding author at: School of Chemical Engineering, Nanjing University of Science and Technology, Xiaolingwei 200, Nanjing 210094, Jiangsu, China. E-mail addresses: [email protected] (B. Chi), [email protected] (Z. Ye).

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http://dx.doi.org/10.1016/j.carbpol.2017.09.071 Received 18 July 2017; Received in revised form 9 September 2017; Accepted 22 September 2017 Available online 23 September 2017 0144-8617/ © 2017 Elsevier Ltd. All rights reserved.

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Another biomaterials widely used in protein delivery systems are polypeptides, which are poly (amino acid)s linked by peptide bonds with structures mimicking natural proteins (Deming, 2007). Previous studies showed that the polypeptide-based drug carriers could enhance protein drugs stability, bioactivity and improve drugs bioavailability (Lee et al., 2009; Sun, Huang, Shi, Chen, & Jing, 2009). However, polypeptides are usually expensive and their synthesis procedures are too complex. In recent years, poly (γ-glutamic acid) (γ-PGA) has attracted considerable attention due to its good biocompatibility and commercial scale availability through microbial fermentation (Zhang, Feng, Zhou, Zhang, & Xu, 2012). It contains lots of carboxyl side group that can be functionalized easily (Gentilini et al., 2012) and is watersoluble, biodegradable, edible and nontoxic towards human (Shih, Wu, & Shieh, 2005). Moreover, the amine group in the N-terminal γglutamic unit can be identified by γ-glutamyl transpeptidase in the cell membrane and thus promoting drug delivery (Peng et al., 2011; Sung, Sonaje, Liao, Hsu, & Chuang, 2012). Therefore, γ-PGA is a desirable natural polymer for proteins delivery. Here, the HA/γ-PGA hydrogels prepared by Schiff base reaction can be formed fastly and possess good biocompatibility because of the excellent biological properties of HA and γ-PGA. Moreover the gelation time and mechanical properties as well as the protein release behavior can be tailored. In this work, the modification of HA and γ-PGA were described. The gelation time, swelling behavior, rheological properties were investigated at various conditions. The release kinetics and mechanism of protein drugs from HA/γ-PGA hydrogels were also analyzed using different mathematical models. Finally, the cell biocompatibility of HA/γ-PGA hydrogels was examined.

Table 1 Compositions and rheological parameters of the γ-PGA/HA hydrogel samples. Sample

DS of γPGAADH/ (%)

OD of HACHO/ (%)

Solid content/ (%)

Molar ratio of −NH2/ −CHO

σc/(Pa)

tan δ

Change the solid content

A1 A2 a A3 A4

36.2 36.2 36.2 36.2

19.2 19.2 19.2 19.2

1 3 5 7

1:1 1:1 1:1 1:1

– 450 660 700

– 0.020 0.025 0.025

Change the molar ration of −NH2/ −CHO

B1 a B2 B3 B4

36.2 36.2 36.2 36.2

19.2 19.2 19.2 19.2

5 5 5 5

2:1 1:1 1:2 1:3

1080 660 619 405

0.0076 0.025 0.029 0.053

Change the DS

C1 C2 a C3 C4

24.3 30.9 36.2 42.4

19.2 19.2 19.2 19.2

5 5 5 5

1:1 1:1 1:1 1:1

790 630 660 380

0.025 0.015 0.025 0.017

Change the OD

D1 D2 a D3 D4

36.2 36.2 36.2 36.2

11.5 15.7 19.2 22.1

5 5 5 5

1:1 1:1 1:1 1:1

234 314 660 670

0.017 0.020 0.025 0.022

a

The samples A3, B2, C3 and D3 are the same sample.

2.4. Preparation and characterization of hydrogels HA-CHO was dissolved in PBS solution (0.01 mol/L, pH = 7.4) overnight in the 4 °C refrigerator and γ-PGA-ADH was dissolved in PBS at room temperature. Then, the two solutions were rapidly mixed by the oscillator (1500 rpm for 5–10 s) to prepare hydrogels. The compositions of all hydrogels are listed in Table 1. Gelation time was measured by the vial tilting method (Wang et al., 2016). According to the compositions in Table 1, HA-CHO and γ-PGAADH were mixed in a vial rapidly. The time until the mixtures don’t flow was considered as the gelation time. This test was carried out in triplicate. Microporous morphologies of hydrogels were observed by scanning electron microscopy (SEM, Philips-FEI, Holland). Firstly, the HA/γ-PGA hydrogels were freeze-dried for 3 days, and then the cross-sectional were treated with gold coating and viewed by SEM. Rheological measurements were carried out with an Anton Paar rheometer (MCR302) using a 25 mm plate–plate sensor with a proper gap. The hydrogel samples were prepared directly on the plate at room temperature. The linear viscoelastic region of all hydrogel samples were determined by the strain amplitude sweep within the range of 0.1–1000% at 1.0 Hz. Frequency sweep tests were carried out at a constant strain obtained from the linear viscoelastic domain. The frequency was varied from 0.6 to 100 Hz. These two tests were performed at 37 °C. For the swelling test, the hydrogel samples with different components were immersed in PBS at 37 °C for 48 h until the hydrogels reach steady state. The swollen hydrogels were taken out and removed the surface water gently, and then weighed (Ws) immediately. After that, the hydrogels were lyophilized for 3 days and weighed (Wd). This test was carried out in triplicate and the swelling ratio was defined as: (Ws − Wd)/Wd

2. Materials and methods 2.1. Materials HA (1000 kDa), Sodium periodate (NaIO4), Hydroxylamine hydrochloride and Ethylene glycol were obtained from Sinopharm Chemical Reagent Co. Ltd (Shanghai, China). γ-PGA (700 kDa) was obtained from Shineking Biotechnology Co., Ltd. (Nanjing, China). Fluorescein isothiocyanate conjugated bovine serum albumin (FITC-BSA), Acridine orange (AO) and Ethidium bromide (EB) were purchased from Solarbio Life Sciences (Beijing, China). 1-Ethyl-3-(3-dimethylaminopropyl)-carbodiimide (EDC), Adipic acid dihydrazide (ADH) and N-hydroxy-succinimide (NHS) were obtained from Aladdin Industrial Corporation (Shanghai, China). All chemicals are analytical and not further purified.

2.2. Synthesis of ploymers HA-CHO with different oxidation degree (OD) were synthesized by oxidizing the proximal −OH groups of HA using NaIO4 (Ito et al., 2007). γ-PGA-ADH with different degree of substitution (DS) were synthesized via a facile carbodiimide coupling reaction of γ-PGA and ADH using EDC and NHS (Yan et al., 2014). The details were in the supplementary data.

2.3. Characterization of polymers The modification of HA and γ-PGA were confirmed by 1H NMR (Bruck Avance 400 MHz) and FTIR spectrophotometer (Nicolet-6700 spectrometer from Thermo Electron). 1H NMR spectra were used to characterize the DS of ADH to the γ-PGA side chains. The DS was defined as the number of substituents per 100 carboxyl groups in γ-PGA. The functional groups of γ-PGA, ADH, γ-PGA-ADH, HA and HA-CHO were identified in the region of 2000–700 cm−1 by FTIR spectroscopy at room temperature.

2.5. Preparation of BSA-loaded hydrogels and BSA release study An in situ polymerization method was used to incorporate FITC-BSA molecules into hydrogel (Leacha & Schmidt, 2005). Briefly, FITC-BSA was initially dissolved in PBS to get a 0.02 wt% solution, and then HACHO and γ-PGA-ADH were dissolved in 0.02 wt% FITC-BSA solution, respectively. After that, these two solutions were mixed to prepare the BSA-loaded hydrogels at room temperature. 101

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(4, 5-dimethylthiazol-2-yl)-2, 5-diphenyltetrazolium bromide (MTT) assay and live/dead assay using NIH 3T3 mouse fibroblast. All sample solutions were sterilized by filtrating via 0.22 μm syringe filters in advance. Briefly, 100 μL 7 wt% HA/γ-PGA hydrogel was prepared at 37 °C in each 96-well cell culture plates, and then washed with PBS three times. Afterwards, NIH 3T3 cells were seeded in 200.0 μL medium which contains 10 vol% fetal bovine serum (FBS) and 1.0 wt% penicillin-streptomycin, and then cultured at 37 °C in an incubator with a 5% CO2. Then, on the 1st, 3rd, and 5th day, the cells were stained with AO and EB, and visualized by an inversed fluorescent microscope (Olympus TH4-200). The living cells were green and the dead cells were tangerine. After that, the optical density values were measured by a microplate reader. MTT assay was carried out in triplicate.

The protein release tests were performed by immersing BSA-loaded hydrogels in 10 mL PBS with occasionally shaking at 37 °C. At different times, the release medium (4 mL) was taken out and replaced with an equal volume of PBS. BSA concentration was quantified by spectrofluorometer (Fluormmax-4, HORIBA Instruments Incorporated) at 517 nm. The total amount of BSA was calculated by a standard curve of FITC-BSA in PBS versus fluorescence intensity (Fig. S1 in supplementary data). All release experiments were performed in triplicate. The kinetics of protein release from HA/γ-PGA hydrogels was studied by several mathematical models. The mechanistic models are described below: Zero-order:

Mt = k0 t M∞

(1)

2.7. Statistical analysis

First-order:

Mt = 1 − e−K1 t M∞

Experimental data were expressed as the mean ± standard deviation, with at least three samples. Statistical analyses were performed using one-way ANOVA followed by post hoc Tukey honestly significant difference (HSD) test, and statistical significance was accepted at P < 0.05.

(2)

Where Mt/M∞ is the fraction of drug released at time t. K0 and K1 represent the zero-order and first-order release kinetic constant, respectively. Higuchi model (Costa & Lobo, 2001):

Mt = Kh t 1/2 M∞ Where Kh represent the Higuchi release kinetic constant. Korsmeyer-Peppas model (Korsmeyer, Gurny, Buri, & Peppas, 1983):

Mt = Kt n M∞

3. Results and discussion 3.1. Synthesis and characterization

(3)

The synthesis scheme of HA-CHO is presented in Fig. 1a. In the HACHO FTIR spectrum (Fig. 1c), a new peak was observed at 1731 cm−1, which was ascribed to the C]O stretch of HA-CHO (Kim et al., 2007). The OD of HA was determined by the hydroxylamine hydrochloride titration method (Li et al., 2014), and the details were in the supplementary data. With the increase in the molar ratio of NaIO4 to HA from 0.7 to 2, the OD increased from 11.5% to 22.1%. Fig. 1b shows the synthesis scheme of γ-PGA-ADH. From the FTIR spectra in Fig. 1c we can see that two new peaks at 1641 and 1517 cm−1 were observed, which were due to the superposition of vibrations of C]O, amide I and II groups both in γ-PGA and ADH (Tsao et al., 2010; Yan et al., 2014). The representative 1H NMR spectra of γPGA, ADH, and γ-PGA-ADH are showed in Fig. 1d. In the γ-PGA-ADH spectrum, a new peak at 1.46 ppm which corresponding to the methylene protons of ADH was observed, indicating that γ-PGA has been hydrated by ADH successfully (Yan et al., 2014). The DS of γ-PGA was determined by comparing the integral areas of signals at δ 1.5-1.75 ppm (ADH, methylene protons) to δ 4-4.5 ppm (γ-PGA, α-protons). With increasing the molar ratio of EDC: −COOH (of γ-PGA) from 0.3 to 2, the DS of γ-PGA increased gradually from 24.3% to 42.4%.

Doelker,

(4)

Where K is a constant incorporating structural and geometric characteristics of the drug dosage form, and n is the release exponent, indicative of the drug release mechanism: n = 0.5 suggests Fickian diffusion with a negligible relaxation coefficient; n = 1 refers to a nonFickian with the characteristic of zero order release; if 0.5 < n < 1, the transport process is anomalous with a comparable structural relaxation to diffusion; n < 0.5 indicates a pseudo-Fickian diffusion with a relative slow release that positively relied on the n value (Li, Fu, & Zhang, 2014). Eq. (4) was modified to take account of a lag period before releasing (Ford et al., 1991): Modified Korsmeyer-Peppas model:

Mt = K '(t − l)n M∞

(5)

Where l is the lag time. Peppas and Sahlin derived Eq. (6) by introducing a second term into Eq. (4): Peppas-Sahlin model:

Mt = K1 t m + K2 t 2m M∞

3.2. Hydrogel preparation and characterization The HA/γ-PGA hydrogels were prepared by the Schiff base reaction between aldehyde groups of HA-CHO and hydrazide groups of γ-PGAADH (Fig. 2a). Both HA-CHO and γ-PGA-ADH solutions were flowable liquids, as shown in Fig. 2b, and a stable hydrogel was formed after mixing. The rheology test reflected that the individual polymer solution was viscous, and an increase in moduli of several orders of magnitude was observed after mixing (Fig. 2c), which is consistent with the result of Fig. 2b and indicates the crosslinking occurred. Fig. 2d shows the FTIR spectroscopy of γ-PGA-ADH, HA-CHO and HA/γ-PGA hydrogel. The curve of HA/γ-PGA hydrogel demonstrated characteristic HA peaks at 1037, 1078 cm−1 for CeOeC and CeO stretching (Kim et al., 2007), and the peaks at 1542, 1600, 1639 cm−1 due to the superposition of the amide I and II groups, C]O and the newly-formed C]N groups (Vieira, Cestari, Airoldi, & Loh, 2008). The microporous morphologies of hydrogels were observed by SEM. As shown in Fig. 2e, the HA/γ-PGA hydrogels showed porous structures and could be severed as a reservoir for lots of water or drugs, which were very suitable for application in

(6)

Where the first term of the right-hand side is the Fickian diffusion contribution, the second term being the Case-II relaxational contribution (Peppas & Sahlin, 1989). Considering the lag times introduced in Eq. (5), Eq. (6) can similarly be rewritten as Eq. (7) (Ford et al., 1991): Modified Peppas-Sahlin model:

Mt = K1 '(t − l)m + K2 '(t − l)2m M∞

(7)

2.6. In vitro cytotoxicity study In vitro cytotoxicity of HA/γ-PGA hydrogels were assessed by the 3102

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Fig. 1. Schematic representations of (a) HA oxidation and (b) γ-PGA modification. (c) FTIR spectra of HA, HA-CHO, γ-PGA, ADH and γ-PGA-ADH. (d) 1H NMR spectra of γ-PGA, ADH and γ-PGA-ADH. The 1H NMR spectra of γ-PGA-ADH obtained from different molar ratio of EDC to −COOH of γ-PGA (0.3, 0.8, 1.2 and 2).

gelation time are demonstrated in Fig. 3c and d. The gelation time gradually decreased along with the decrease of DS or increase of OD. As shown in Table 1, the content of γ-PGA-ADH in hydrogels (C4 → C1) increased with the decrease of DS. Similarly, the content of γ-PGA-ADH in hydrogels (D1 → D4) also increased with the increase of OD. It is clearly that the gelation time decreased with increasing the γ-PGA-ADH content. Increasing the γ-PGA-ADH content maybe has two effects on hydrogel: more carboxyl groups were provided and more intermolecular hydrogen bonds were formed with HA; the hydrogen bonds enhanced the intermolecular entanglement. The influence of these two aspects will increase the crosslinking degree of the HA/γ-PGA hydrogels, as depicted in Fig. 3i. The swelling ratios of the hydrogels are showed in Fig. 3e–h. All hydrogels showed high swelling ratios, and the highest swelling ratio could reach 70. Fig. 3e showed that when the solid content was 1 wt%, the hydrogel was soluble in PBS. With increasing the solid content, the swelling ratio decreased, which may be due to the high crosslinking density (Yan et al., 2016).The influence of −NH2/−CHO molar ratio on the swelling ratio is presented in Fig. 3f. As shown in Fig. 3f, when the −NH2/−CHO molar ratio was 1:2, the swelling ration was about 2 fold higher compared to the 2:1 hydrogel, and when −NH2/−CHO was 1:3, the hydrogel was soluble in PBS because of the low crosslinking

drug delivery.

3.3. Gelation time and swelling behavior Gelation time plays an important role in the clinical application, and this parameter should be fast to prevent hydrogel from premature dissolution upon administration and controllable conveniently to expand application range. Obviously, solid content has a remarkable influence on the gelation time (Fig. 3a). The gelation time was reduced drastically with an increase in solid content. The gelation time of the hydrogel with 1 wt% solid content was more than 3 h, whereas, the gelation time reduced to 14 s drastically in 7 wt% solid content. This may be due to the increase in the reactive groups per unit volume leads to the accelerated gelation and short gelation time (Yan et al., 2016). Fig. 3b shows the effect of the molar ratio of −NH2/−CHO on the gelation time. With the increase in the molar ratio of −NH2/−CHO, the gelation time decreased from 117 s to 16 s because of the increase of the −NH2 group. Meanwhile, we noticed an interesting phenomenon. When the molar ratio of −NH2/−CHO were 1:1 and 2:1, there was no significant difference in the gelation time (p > 0.05), suggesting that the γ-PGA-ADH may be involved in the crosslinking process by winding (Fig. 3i). The influences of DS of ADH to γ-PGA and OD of HA on 103

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Fig. 2. (a) Schematic of in situ hydrogel crosslinking by Shiff base reaction. (b) Photographs for the hydrogel formation when 3 wt% HA-CHO solution mixed with 3 wt% γ-PGA-ADH solution. (c) Oscillatory time sweeps of individual polymer solutions and hydrogel; storage modulus (G', filled symbols) and loss modulus (G″, empty symbols) at 1 Hz, 1.0% strain. (d) FTIR spectra of HA-CHO, γ-PGA-ADH and HA/γ-PGA hydrogel. (e) SEM images of the hydrogel at 5 wt%.

3.4.1. Amplitude sweep test Linear viscoelastic domain which can reflect the processing range of hydrogels was obtained from amplitude sweep tests. The results of the storage modulus (G') and loss modulus (G″) as a function of the strain are showed in Figs. 4a and S2 . All the HA/γ-PGA hydrogels had a broad linear viscoelastic domain, which indicated that the hydrogels had a wide processing range. As shown in Fig. 4a, when the strain was less than 40%, the values of G′ and G″ remained almost constant, which indicated the structure of hydrogels remained undamaged under relatively large deformations (Jia & Zhu, 2015). Table 1 reports the threshold shear stress (σc) obtained from Figs. 4b and S3, and the

density. The swelling behavior was also influenced by the DS and OD (Fig. 3g and h), and the swelling ratio was reduced with the decrease of DS or increase of OD, which may be ascribed to the increase of crosslinking density (Yan et al., 2016).

3.4. Rheological properties of hydrogels Rheological measurements were carried out to assess the working ability of HA/γ-PGA hydrogels in drug delivery and were conducted to investigate amplitude sweep and frequency sweep tests. 104

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Fig. 3. Gelation time of HA/γ-PGA hydrogels with different (a) solid contents, (b) −NH2/−CHO molar ratio, (c) DS of ADH to γ-PGA and (d) OD of HA; Swelling ratio of the hydrogels as a function of (e) solid contents, (f) −NH2/−CHO molar ratio, (g) DS of ADH to γ-PGA and (h) OD of HA; (i) Schematic representation of chemical junction (left) and a multifold junction (right). (n = 3, *p < 0.05).

approximately an equal number of −NH2 and −CHO groups, but sample B1 with the −NH2/−CHO molar ratio of 2:1 showed the highest σc. This implies that HA/γ-PGA hydrogels not only have chemical crosslinking from Shiff base reaction, but also have the intermolecular hydrogen bond and physical entanglement as hypothesized above (Fig. 3i). It is clearly in Table 1 that with the decrease of DS or increase of OD, the σc increased, which indicated that the antishear ability was improved due to the increased crosslinking density.

tangent phase angle (tan δ) expressed as G″/G' (Shui, Guo, Chen, Xu, & Zheng, 2005). From Table 1 we can see that the values of tan δ were far below 1, which indicated that the hydrogels were stiff and the elastic property was predominant (Jia & Zhu, 2015; Wang and Zhou, 2009). It should be noticed that the σc increased with the value of solid content, which revealed that the higher solid content could improve the shear resistance. This phenomenon may be due to the increase of crosslinking density. The σc also increased with increasing the molar ratio of −NH2/ −CHO. A higher crosslinking degree could be obtained at 105

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Fig. 4. (a) The G' (filled symbols), G″ (empty symbols) as a function of the strain and (b) the applied stress versus the strain for hydrogel A3. The G' and G″ as a function of frequency with different (c) solid contents, (d) −NH2/−CHO molar ratio, (e) DS of ADH to γ-PGA and (f) OD of HA, respectively.

structure (Galatanu, Chronakis, Anghel, & Khan, 2000). With decreasing the DS or increasing the OD, G′ increased, as demonstrated in Fig. 4e and f, which suggested that increasing the γ-PGA-ADH content could make the internal structure of hydrogel become more stable (Galatanu et al., 2000).

3.4.2. Frequency weep test The frequency versus G′ and G″ are showed in Fig. 4c–f. All samples showed the elastic behavior was dominant rather than the viscous behavior, which suggested that characteristic of a “strong hydrogel” (Bai, Sheng, Zhang, Li, & Shi, 2011; Tan, Rubin, & Marra, 2011). G′ increased with solid content, as shown in Fig. 4c, which indicated that the hydrogel was stiffer at higher solid content (Yan et al., 2016). This phenomenon is also consistent with the result of threshold shear stress. With increasing the −NH2/−CHO molar ratio, the G′ increased (Fig. 4d), and when the −NH2/−CHO molar ratio was 2:1, the G′ was about 2 fold higher compared to the 1:1 hydrogels, which displayed an intensive connection of the network and a more stable internal

3.5. In vitro drug release study BSA was used as a model protein drug to study the drug release behavior. Fig. 5 shows the BSA-loaded hydrogels release kinetics. Compared with the release time of drugs from injectable HA hydrogels prepared by physical mixing (about a few hours) (Chen, Leu, Fang, 106

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Fig. 5. In vitro release profiles of FITC-BSA from HA/γ-PGA hydrogels with different (a) solid contents, (b) −NH2/−CHO molar ratio, (c) DS of ADH to γ-PGA and (d) OD of HA at 37 °C in 0.01 mol/L pH = 7.4 PBS.

generalization of Fickian diffusion and Case-II relaxation (Sinclair & Peppas, 1984). Therefore, a better fitting was obtained. But Eq. (4) is based on the assumption that release occurs as soon as the hydrogel contact the release medium (Ford et al., 1991) and ignores a lag period. So the introduction of a lag period (Eq. (5)) resulted in better fitting than Eq. (4). All values of n were less than 0.5 and l values were positive, which indicated that all samples showed pseudo-Fickian diffusion with a lag period (Li et al., 2014). Eqs. (4) and (5) are regardless of the form of the constitutive equation and the type of coupling of relaxation and diffusion (Sinclair & Peppas, 1984). For further research, Eqs. (6) and (7) were used to distinguish the roles of relaxation and diffusion. The curve-fitting data of Eq. (6) showed that although values of K2 exceeded K1, no trends were found for the effects of solid contents, −NH2/−CHO molar ratio, DS and OD on their values.

Chen, & Fang, 2011; Ha, Lee, Chong, & Lee, 2006), the release of BSA from HA/γ-PGA hydrogels with various compositions showed sustained release behavior (> 30 h). The crosslinking density is one of the factors that affect drug release (Kong, Kim, & Park, 2016), so the factors influencing the hydrogel crosslinking degree also affect the drug release. As shown in Fig. 5, the release rate of BSA gradually decreased as the crosslinking degree increased. The release curves were fitted by the software Origin 8.0 according to several kinetic models. Correlation coefficient (r2) values and kinetic constants are summarized in Tables 2 and S1. When hydrogels are used as drug carriers, the release mechanism is complicated due to drug diffusion and self-swelling (Caccavo, Cascone, Lamberti, & Barba, 2015). As shown in Table S1, the zero-order, first-order and Higuchi functions showed poor fitting. Korsmeyer-Peppas model (Eq. (4)) is a Table 2 Drug release kinetics and correlation coefficient values from different kinetic models. Sample

A2 a A3 A4 B1 a B2 B3 C2 a C3 C4 D1 D2 a D3 D4 a

Korsmeyer-Peppas

Modified Korsmeyer-Peppas

Peppas-Sahlin

r2

n

l

r2

K1

K2

r2

K 1′

K 2′

l

r2

0.9308 0.9653 0.9764 0.9682 0.9653 0.9499 0.9596 0.9653 0.9913 0.9860 0.9825 0.9653 0.9527

0.1960 0.2228 0.2581 0.2248 0.2228 0.1751 0.2191 0.2228 0.2642 0.2758 0.2531 0.2228 0.2138

0.7847 0.8827 0.7969 0.9162 0.8827 0.9443 0.5494 0.8827 0.2139 0.4056 0.3823 0.8827 0.8643

0.9436 0.9861 0.9863 0.9961 0.9861 0.9928 0.9619 0.9861 0.9910 0.9868 0.9830 0.9861 0.9730

−265.0 −9.261 −1.888 −22.41 −9.261 −440.7 −3.810 −9.261 −0.3520 −0.7823 −0.9663 −9.261 −60.00

265.3 9.531 2.136 22.66 9.531 441.0 4.149 9.531 0.6655 1.051 1.274 9.531 60.29

0.9573 0.9852 0.9851 0.9947 0.9852 0.9848 0.9697 0.9852 0.9915 0.9889 0.9854 0.9852 0.9797

−296.9 −0.5552 0.05829 0.4010 −0.5552 0.5149 −193.8 −0.5552 −122.5 −284.3 −395.6 −0.5552 −144.9

297.3 0.9341 0.2874 −0.03011 0.9341 −0.04751 194.0 0.9341 122.5 284.3 395.7 0.9341 145.2

−0.1990 0.7615 0.8195 0.8288 0.7615 0.8953 −1.040 0.7615 −2.900 −2.405 −2.076 0.7615 −0.1527

0.9540 0.9851 0.9851 0.9969 0.9851 0.9926 0.9723 0.9851 0.9926 0.9915 0.9885 0.9851 0.9782

The samples A3, B2, C3 and D3 are the same sample.

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Fig. 6. (left) Percentage viability of NIH 3T3 cells seeded on the surface of HA/γ-PGA hydrogel. (right) Staining micrographs of the NIH 3T3 cells at different culture times. Bare cell culture plate surface as control (n = 3, *p < 0.05).

Case-II relaxation depending on the different hydrogel components. Furthermore, the hydrogels also possessed good biocompatible. Since the fast gelling ability, good mechanical properties and superior biocompatibility, the HA/γ-PGA hydrogels are expected to be a promising candidate as protein carrier and may have potential applications in cell carrier and cartilage tissue engineering.

The values of K1 were negative and K2 values were positive, which indicated that the Fickian diffusion release was an inhibited release and the release mechanism was followed Case-II relaxation (Ford et al., 1991). Considering the lag time (Eq. (7)), we found that the l values of samples (A2, C2, C4, D1, D2 and D4) were negative, indicating the lag time did not exist. Comparing the r2 values of different models fitting, we found the best fitting were obtained by Eq. (6), so these samples (A2, C2, C4, D1, D2 and D4) followed Case-II relaxation mechanism. The r2 values of sample A3, A4 and B3 from Eq. (5) fitting were higher compared to Eq. (6 and 7), indicating the drug release from sample A3, A4 and B3 followed pseudo-Fickian diffusion with a lag period. The samples B1 obtained the best fitting from Eq. (7), and the K2′ was negative, l value was positive, so B1 followed Fickian diffusion with a lag period.

Acknowledgments This work was supported by the National Natural Science Foundation of China (31771049, 31401588 and 51403103), the National Basic Research Program of China (973 Program) (2013CB733603), and State Key Laboratory of Materials-Oriented Chemical Engineering (ZK201606 and ZK201403).

3.6. In vitro cytotoxicity study Appendix A. Supplementary data In order to test the biocompatibility of the HA/γ-PGA hydrogels, they were synthesized directly on the bottom of 96-well cell culture plates. The cell viability was quantitatively determined via MTT assay and live/dead assay for 1, 3, and 5 days. The percentage viability of the cells cultured on the hydrogel surface is more than 85%, as shown in Fig. 6, and there is no significant difference was observed on culture time (p > 0.05). From the staining micrographs, we can see the cells displayed normal morphology, and could well adhere and proliferate on the HA/γ-PGA hydrogel surfaces. With increasing the culture time, the green points showed an obvious increase, and only a few tangerine points were observed, demonstrating a good cytocompatibility of HA/γPGA hydrogels.

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4. Conclusions In summary, we developed novel injectable hydrogels based on HA and γ-PGA to serve as protein carriers that can be rapidly formed in situ by Shiff base reaction. In this work, the HA-CHO and γ-PGA-ADH were synthesized successfully and confirmed by 1H NMR and FTIR. The gelation time, swelling ratio and rheological properties, as well as the drug release rate could be easily tailored by solid contents, −NH2/ −CHO molar ratio, DS of ADH to γ-PGA and OD of HA. Moreover, rheological measurements illustrated the hydrogels had wide processing ranges, good mechanical properties, and stable internal structure. The release of BSA from HA/γ-PGA hydrogels showed sustained release behavior (> 30 h). Different mathematical model fitting showed the release mechanisms of HA/γ-PGA hydrogels were either diffusion or 108

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