An Injectable Hydrogel with Slow Degradability Composed of Gelatin

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An Injectable Hydrogel with Slow Degradability Composed of Gelatin and Hyaluronic Acid Crosslinked by Schiff’s Base Formation Takuro Hozumi, Tatsuto Kageyama, Seiichi Ohta, Junji Fukuda, and Taichi Ito Biomacromolecules, Just Accepted Manuscript • DOI: 10.1021/acs.biomac.7b01133 • Publication Date (Web): 29 Dec 2017 Downloaded from http://pubs.acs.org on December 31, 2017

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An Injectable Hydrogel with Slow Degradability Composed of Gelatin and Hyaluronic Acid Crosslinked by Schiff’s Base Formation Takuro Hozumi,† Tatsuto Kageyama,‡ Seiichi Ohta,§ Junji Fukuda‡ and Taichi Ito*,†,§

†

Department of Chemical System Engineering, The University of Tokyo, 7-3-1 Hongo,

Bunkyo-ku, Tokyo, 113-8656, Japan

‡

Faculty of Engineering, Yokohama National University, 79-5 Tokiwadai, Hodogaya-ku,

Yokohama, Kanagawa, 240-8501, Japan

§

Center for Disease Biology and Integrative Medicine, The University of Tokyo, 7-3-1

Hongo, Bunkyo-ku, Tokyo, 113-8655, Japan

KEYWORDS Schiff’s base, gelatin, hyaluronic acid, hydrolysis

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ABSTRACT

We developed an injectable gelatin/hyaluronic acid hydrogel with slow degradability, which consisted of carbohydrazide-modified gelatin (Gel-CDH) and hyaluronic acid mono-aldehyde (HA-mCHO). Gel-CDH/HA-mCHO hydrogels were degraded much more slowly in phosphate-buffered saline than the other Schiff’s base crosslinked gelatin/hyaluronic acid hydrogels that were comprised of native gelatin, adipic acid dihydrazide-modified gelatin, or hyaluronic acid di-aldehyde because of stable Schiff’s base formation between aldehyde and carbohydrazide groups, and suppression of ring-opening oxidation by mono-aldehyde modification. This prolonged degradation would be suitable for inducing angiogenesis. Therefore, the Gel-CDH/HA-mCHO hydrogels were sufficiently stable during the angiogenesis process. In addition, the hydrogel had a pore size of 15–55 µm and a shear storage modulus of 0.1–1 kPa, which were appropriate for scaffold application. Ex vivo rat aortic-ring assay demonstrated the concentration dependency of microvascular extension in the Gel-CDH/HA-mCHO hydrogel. These results demonstrated the potential usefulness of Gel-CDH/HA-mCHO hydrogel for tissue-engineering scaffolds.


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INTRODUCTION In situ crosslinkable hydrogels1,2 are useful for tissue-engineering scaffolds, because they can be injected into molds for in vitro culture matrix fabrication and into tissue defects for in vivo regeneration. After scaffolds have adapted to the transplanted site, infiltration of parenchymal cells into the scaffolds and angiogenesis are necessary. To achieve efficient tissue regeneration and mechanical stability, hydrogel scaffolds with controlled functions and properties are essential. Extracellular matrix (ECM)-derived polymers are widely used for hydrogels, because they can offer a similar environment to living tissues,2 which is suitable for angiogenesis and other cell recruitment. Hyaluronic acid (HA) and gelatin (Gel) are examples of such ECM-derived polymers. HA has many biological functions3, including stimulation of cytokine production4 and angiogenesis.5 Gel,6 another example, is also a biodegradable and low immunogenic polymer that allows cell attachment. Because of these functions of HA and Gel, several injectable hydrogels composed of Gel and HA have been developed for tissue-engineering scaffolds. Hydrogel scaffolds composed of thiolated HA, thiolated Gel and poly(ethylene glycol) diacrylate were developed for a tubular tissue construct7 and as a filler to repair osteochondral defects.8 Cloyd et al. prepared hydrogels by mixing partially oxidized HA and Gel with appropriate material properties in unconfined compression for nucleus



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pulposus support.9 Kageyama et al. fabricated hydrazide-modified Gel and oxidized HA-based hydrogel to support perfusable vascular networks in vitro.10 To use these ECM-based injectable hydrogels as scaffolds, it is important to control the hydrogel properties, including degradation kinetics and mechanical strength. For example, ideal hydrogel scaffolds should degrade at an optimum rate to be gradually replaced by regenerated tissue, and finally excreted from the body. This optimum degradation rate would depend on various factors, including tissue type and size. For example, in rats, nerve myelination needed approximately 2–8 weeks,11 whereas acellular bladder matrix took 12 weeks to recruit a comparable amount of smooth muscle compared with normal bladder.12 In addition, recipient liver volume doubled 1–2 months after living right-lobe transplantation.13 Therefore, a methodology for controlling degradation kinetics of injectable hydrogel scaffolds over a few weeks to a few months is required. To control the degradation rate, various chemical crosslinking methods,14,15 including Michael addition,16 click reaction17 and Schiff’s base formation18,19 have been investigated. Among them, Schiff’s base formation has been widely explored, because of rapid crosslinking and excellent biocompatibility. It has been used to fabricate hydrogels with various

polymers,

including

Gel/carboxyl

methyl

cellulose,20

chitosan/HA21

and

Gel/alginate.22 Despite numerous previous studies, control of the degradation rate has remained a challenge in Schiff’s base crosslinked hydrogels.23 Recent studies have shown 4 ACS Paragon Plus Environment

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that the hydrolytic stability of Schiff’s base is enhanced by reducing the electrophilicity of carbonyl carbon.24 An increase of nucleophilicity of the amine derivatives by electron donation to C=N and π-π conjugation25 contributed to the hydrolytic stability. Hydrazone- or oxime-crosslinked hydrogels using benzoic aldehyde26 as carbonyls, while hydroxylamines and hydrazides18,27 were included as counterparts, were stable to hydrolysis at physiological pH. To control the degradation rate of injectable hydrogels, the effect of the crosslinking structure requires further study. In the present study, we synthesized a novel Gel derivative, carbohydrazide (CDH)-modified Gel. Because the imine structure formed by CDH generates resonance structure, CDH modified Gel would contribute to more stable Schiff’s base formation than native Gel. As a counterpart, mono-aldehyde modification of the HA side chain was employed. Mono-aldehyde HA would contribute to a greater stability of hydrogels formed with this HA form against hydrolysis than hydrogels containing conventional di-aldehyde HA oxidized simply by periodate, because the ring-opened structure of polysaccharides was reported to be susceptible to hydrolysis.28,29 Using native Gel and a Gel derivative as well as an aldehyde-modified HA for comparison, the effect of the crosslinking structure on the stability of Gel/HA hydrogels was investigated. Furthermore, we tested the potential of the hydrogel as a scaffold by in vitro cell-viability assay and ex vivo rat aortic-ring assay.



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MATERIALS AND METHODS Materials. Gel from porcine skin type A was kindly donated by Nitta Gelatin (Osaka, Japan). HA was kindly provided by Denka (Tokyo, Japan). Adipic acid dihydrazide (ADH), CDH, sodium chloride (NaCl), ethanol, glycine, 2,4,6-trinitrobenzenesulfonic acid sodium salt dihydrate (TNBS), polyethylene glycol (PEG) of approximate molecular weight 1, 2, 15–25 and 300–500 kDa (used as gel permeation chromatography (GPC) standards), sodium periodate (NaIO4), ethylene glycol, tert-butyl carbazate (t-BC) and dimethyl sulfoxide (DMSO) formalin solution were purchased from Wako Pure Chemical Industries, Ltd. (Osaka, Japan). 1-Hydroxybenzotriazole monohydrate (HOBt) was obtained from Tokyo Chemical Industry (Tokyo, Japan). 1-Ethyl-3-(3-dimethylaminopropyl) carbodiimide monohydrochloride (WSCD) was purchased from the Peptide Institute (Osaka, Japan). (±)-3-Amino-1,2-propanediol, Triton-X100 and BS1 Lectin-FITC (Lectin from Bandeiraea simplicifolia FITC Conjugate) (cat. no. L9381) were obtained from Sigma Aldrich (St. Louis, USA). Sodium cyanoborohydride (NaBH3CN) was purchased from MERCK (Darmstadt, Germany). Dextran of approximate molecular weight of 8, 15, 10 and 500 kDa (GPC standards) were obtained from Extra Synthese (Genay, France). Ninety-six-well plates (IWAKI 3860-096) were obtained from AGC Techno Glass (Shizuoka, Japan). Human umbilical vein endothelial cell (HUVEC, C2517A) and EGM-2 Bullet Kit (CC-3162) was obtained

from

Lonza

(Bazel,

Switzerland).

Cell

Counting

Kit-8

and

DAPI 6

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(4′,6-diamidino-2-phenylindole) solution were purchased from Dojindo Laboratories (Kumamoto, Japan).

Synthesis of hydrazide-modified Gels. For CDH-modified Gel (Gel-CDH) preparation, 3 g Gel was dissolved in 300 mL pure water in a 500 mL eggplant flask and 2.2 g CDH was added with stirring. Then, 0.45 g HOBt dissolved in 10 mL DMSO and 0.46 g WSCD dissolved in 10 mL H2O were added dropwise with stirring. The pH of the solution was adjusted to 5 and the amidation reaction was performed at room temperature overnight. The product was dialyzed in approximately 0.3 M NaCl solution, 25%(v/v) ethanol solution and pure water for 2, 1 and 4 days, respectively, using a dialysis membrane with a 6,000–8,000 Da molecular weight cut off and lyophilized. ADH-modified Gel (Gel-ADH)10,

20

was

synthesized following the same protocol, using ADH instead of CDH and was used as a control material. (Figure 1)



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Figure 1. Fabrication of Gel/HA-based hydrogel scaffolds.

Synthesis of aldehyde-modified HA. HA was modified with a mono-aldehyde group on the side chain (HA-mCHO) following a previously reported protocol19 with minor modification. HA was first modified with diol (HA-diol) as follows: 1 g HA, 2.5 mmol disaccharide unit,was dissolved in 60 mL pure water and 5.0 mmol (±)-3-amino-1,2-propanediol was added dropwise with stirring. Next, 2.5 mmol HOBt dissolved in 5 mL DMSO was added to the mixture and the pH was adjusted to 6 using 1 M hydrochloric acid solution, then 0.75 8 ACS Paragon Plus Environment

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mmol WSCD dissolved in 5 mL pure water was added to the mixture. After overnight reaction, HA-diol was purified by dialysis in 0.1 M NaCl solution and pure water for 2 and 1 days, respectively, using a dialysis membrane with 6,000–8,000 Da molecular-weight cut off and lyophilized. The conjugated diol groups were oxidized to obtain HA-mCHO. A total of 200 mg HA-diol were dissolved in 25 mL pure water and 0.5 mmol NaIO4 dissolved in 0.5 mL pure water were added dropwise. The oxidation reaction was performed for 5 min in the dark with stirring, and subsequently 10 mmol ethylene glycol was added to quench the reaction. HA-mCHO was obtained by dialysis in pure water for 1 day using a dialysis membrane with a 6,000–8,000 Da molecular-weight cut off and lyophilized. Di-aldehyde-modified HA on the main chain (HA-dCHO) was synthesized as reported previously.30,31 In brief, 200 mg HA was dissolved in 25 mL pure water and 0.5 mmol NaIO4 in 0.5 mL pure water were added dropwise. The oxidation reaction was performed for 1 h in the dark with stirring, and subsequently 10 mmol ethylene glycol was added. HA-dCHO was obtained by dialysis in pure water for 1 day using a dialysis membrane with a 6,000–8,000 Da molecular-weight cut off and lyophilized. The modification ratio of the aldehyde group was determined by reaction with t-BC followed by reduction with NaBH3CN.32 The aldehyde-modified HA was dissolved in pure water at a concentration of 10 mg/mL and 0.5 M t-BC solution, in a 10-fold excess over the NaIO4 used in the aldehyde modification, was added. The mixture was stirred for 1 h at room 9 ACS Paragon Plus Environment


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temperature. A 0.5 M NaBH3CN aqueous solution, equimolar to t-BC, was added and allowed to react for 24 h. The samples were purified by dialysis in a 1,000 molecular-weight cut off membrane and lyophilized. The aldehyde content was evaluated by 1H NMR measurement using a αJEOL JNM-LA400 spectrometer (JEOL, Tokyo, Japan). The degree of aldehyde modification was calculated by comparing the signal of tert-butyl substituent (1.4 ppm, 9H) with that of HA acetamide (1.9 ppm, 3H).

Characterization of precursor polymers. Molecular weights of the synthesized precursor polymers were determined by GPC using a TSK-GEL GMPWXL column equipped with a Shimadzu pump (LC-10AD) and a refractive index detector (830-RI; Jasco, Tokyo, Japan). PEG standards and dextran standards were used for Gel derivatives and HA derivatives, respectively. The mobile phases were 0.2 M phosphate buffer (pH 6.8) for Gel derivatives and 50 mM phosphate buffer with addition of 0.2 M NaCl (pH 6.7) for HA derivatives at a flow rate of 0.5 mL/min. In addition, the degradation kinetics of HA derivatives was measured. HA, HA-mCHO and HA-dCHO were dissolved each in 50 mM phosphate buffer with addition of 0.2 M NaCl (pH 6.7) at a concentration of 1mg/mL and incubated at 37°C for 8 days. Then, the molecular weight of incubated HA derivatives was determined by GPC.


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Fourier-transform infrared (FT-IR) spectroscopy of Gel and HA derivatives was performed using an FT-IR spectrophotometer (FT/IR-4200ST; Jasco, Tokyo, Japan) to confirm the modification. Samples were mixed with potassium bromide and formed into tablets, and the measurement was performed at 4,000–400 cm−1 at a resolution of 4 cm−1.

Characterization of Gel/HA hydrogels. Gel- and HA-derivative solutions were prepared in phosphate-buffered saline (PBS). The hydrogels were fabricated using a double-barreled syringe fibrin glue applicator (Baxter, Deerfield, USA), in which one syringe was filled with the Gel-derivative solution and the other with an equal volume of HA-derivative solution. The applicator was fitted with an 18G needle. We prepared Gel/HA hydrogels with three different concentrations: 2.5% Gel derivatives and 1.0% HA derivatives, 5.0% Gel derivatives and 2.0% HA derivatives, and 7.5% Gel derivatives and 3.0% HA derivatives. These

are

denoted

as

Low,

Mid

and

High,

respectively.

Gelation

time

of

Gel-CDH/HA-mCHO hydrogel was measured by referring to the previous research30. The swelling and degradation behavior of the hydrogels were evaluated by immersing 0.4 ml hydrogels in 20 ml PBS at 37°C. The hydrogels were weighed at various time intervals until the samples were totally degraded. The swelling ratio of the test sample (Qt) was defined as

Qt = Wt W 0 , 11 ACS Paragon Plus Environment


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where Wt is the weight of the swollen hydrogel and W0 is the initial weight of the hydrogel. The time till the complete hydrogel degradation was defined to be the characteristic time of degradation, when Qt = 0. Morphology of the hydrogel was observed using Scanning Electron Microscopy (SEM) on an FE-SEM S-900 (Hitachi, Tokyo, Japan). The hydrogel was embedded in O.C.T. Compound (Sakura Finetek USA, Inc., Torrance, USA) and frozen in liquid nitrogen. The samples were sectioned using a scalpel and cross-sections were coated with platinum. Hydrogel pore size was determined based on four images each hydrogel using ImageJ software (National Institutes of Health, Bethesda, USA). Dynamic viscoelastic measurement was performed to determine the mechanical strength of the hydrogels. Linear viscoelastic region was confirmed at 5% strain (data not shown). Hydrogel discs of 25-mm diameter and 0.5-mm height were prepared by mixing equal volumes of Gel-CDH and HA-mCHO solutions at various concentrations, followed by overnight incubation at 37 °C to ensure complete gelation. A frequency sweep test was performed at 10–0.01 Hz with a 5% strain at 37 °C.

Cell-viability test. In vitro cell viability in the presence of Gel-CDH or HA-mCHO was determined using a Cell Counting Kit-8 following the standard protocol for a 96-well plate. HUVECs were maintained in EGM-2 Bullet Kit medium at 37 °C under 5% CO2 and used at 12 ACS Paragon Plus Environment

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passage six. HUVECs were seeded on a 96-well plate at 4.8 × 103 cells/well with 100 µL of the medium and incubated at 37 °C under 5% CO2 overnight. The culture medium was replaced with 100 µL EGM-2 Bullet Kit containing different concentrations of Gel-CDH or HA-mCHO. Forty-eight hours after the medium change, 10 µL Cell counting Kit solution was added to each well and incubated a further 2 h. The absorbance at 450 nm was determined using a plate reader (2030 ARVO V3; PerkinElmer, Waltham, MA, USA) and normalized to that of 8 control wells where only EGM-2 was added.

Ex vivo angiogenesis assay. The aortic-ring assay33 was performed to characterize the Gel-CDH/HA-mCHO hydrogel as a scaffold. All animal experiments were performed in accordance with the Guidelines of Animal Experiments of The University of Tokyo, and the protocols were approved by the Animal Care Committee of The University of Tokyo. Gel-CDH and HA-mCHO were first sterilized for 2 h under UV irradiation and dissolved in sterile

PBS.

Seven-week-old male

Sprague-Dawley rats were maintained in a

temperature-controlled room under a 12 h light-dark cycle. At age 8 week, the rats were euthanized by CO2 inhalation. The chest cavity was opened and an aorta was dissected from the thoracic cavity. The aorta was washed with EGM-2 basal medium, followed by the removal of surrounding tissues. The dissected aorta was cut in round slices of 1-mm width, placed in each well of a 96-well plate and covered by 100 µL of Gel-CDH/HA-mCHO 13 ACS Paragon Plus Environment


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hydrogel formed in situ. After 1 h incubation at 37 °C under 5% CO2 to complete the gelation, 100 µL EGM-2 was added (day 0). The medium was replaced daily and cell migration from the embedded aorta was observed under a microscope (IMT-2, OLYMPUS, Tokyo, Japan) equipped with a 5× objective lens. The maximum distance between the migrating cells and the aorta was determined using ImageJ software. At day 7, the aortas were stained with DAPI and BS1 Lectin-FITC. The constructs were fixed using 4% formalin and permeabilized with 0.1% Triton-X100. Immunostaining with DAPI for nuclei and BS1 Lectin-FITC for endothelial cells was performed according to the manufacture’s instruction to observe the cell migration. The fluorescence images were obtained using a fluorescence microscope (BZ-X700, Keyence, Osaka, Japan).

RESULTS AND DISCUSSION Synthesis and characterization of precursor polymers. The new Gel derivative, Gel-CDH, was synthesized by carbodiimide chemistry. For a comparison, we also used ADH-modified Gel generated according to our previous reports.10,20 Figure 2(A) represents FT-IR spectra of synthesized Gel derivatives. The hydrazide modification to Gel was indicated by the increase of the peaks at 1240 cm−1 (Amide III), 1543 cm−1 (Amide II), 1655 cm−1 (Amide I) and 3322 cm−1 (Amide A). The modification ratio of hydrazide on Gel was determined to be 26% for 14 ACS Paragon Plus Environment

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Gel-ADH and 33% for Gel-CDH by the TNBS assay34. Including amine groups contained in the Gel backbone, the amount of NH2 was calculated to be 1.1 mmol/g for both Gel-ADH and Gel-CDH. In addition, we determined the molecular weight distribution of native Gel, Gel-ADH and Gel-CDH by GPC (Figure 2(C)). The resulting mean molecular weights of native Gel, Gel-ADH and Gel-CDH were 107, 40.6 and 36.7 kDa, respectively (Table 1). Although the mean molecular weight was decreased after the hydrazide modification because of the degradation of the Gel backbone, the difference was not significant. Mono-aldehyde-modified HA, HA-mCHO, was synthesized according to a previous report.19 Di-aldehyde-modified HA, HA-dCHO, was also synthesized according to our previous study.31 While diol compound conjugated to HA was oxidized in HA-mCHO, hydroxyl groups on C2-C3 of the six-membered ring were oxidized in HA-dCHO. The synthesized polymers were characterized by FT-IR (Figure 2(B)). The FT-IR spectra of HA, HA-dCHO and HA-mCHO were similar, making it difficult to distinguish the signal corresponding to the aldehyde groups35 at 1730 cm−1. This may result from the low degree of oxidation and the formation of hemiacetals. Rather than by FTIR, the number of aldehyde groups was characterized by reaction with tBC followed by reduction with NaBH3CN.32 The modification ratio of HA-mCHO and HA-dCHO was 12% and 4%, respectively. The molecular weight distribution of HA derivatives were determined by GPC (Figure 2(D)). The weight-average molecular weights of native HA, HA-mCHO and HA-dCHO were 19.1, 11.8 15 ACS Paragon Plus Environment


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and 8.06 MDa, respectively (Table 1). In addition, Table 2 showed that the molecular weight of HA-dCHO decreased dramatically compared with native HA and HA-mCHO after 8 days of incubation in the phosphate buffer. This result indicated that HA-mCHO was much more stable against hydrolysis than HA-dCHO. Because synthesized Gel derivatives, Gel-CDH and Gel-ADH, and HA derivatives, HA-mCHO and HA-dCHO, had approximately the same modification rate and molecular weight, these polymers would be suitable for exploring the effect of crosslinking structure, as reported in the following experiments.


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Figure 2. Characterization of gelatin (Gel) and hyaluronic acid (HA) derivatives. FT-IR spectra of (A) Gel derivatives and (B) HA derivatives. Specific peaks derived from hydrazide or aldehyde modification are indicated by dashed lines. Molecular weight distribution of (C) Gel and (D) HA derivatives.

Table 1. Summary of number- and weight-average molecular weight (Mn and Mw) and polydispersity index (PDI) of gelatin (Gel) and hyaluronic acid (HA) derivatives measured by gel permeation chromatography Mn [kDa]

Mw [kDa]

PDI

Mn [MDa]

Mw [MDa]

PDI

Gelatin

27.7

107

3.86

HA

1.23

19.1

15.5

Gel-ADH

15.4

40.6

2.63

HA-dCHO

0.455

8.06

17.7

Gel-CDH

15.2

36.7

2.42

HA-mCHO

0.550

11.8

21.5

Gel-CDH/HA-mCHO hydrogel degraded much more slowly than Gel-CDH/HA-dCHO, Gel-ADH/HA-dCHO and Gel/HA-dCHO hydrogels. Synthesized Gel-CDH immediately formed hydrogel by mixing with both HA-mCHO and HA-dCHO (Figure 3(A)). Gelation time of Low Gel-CDH/HA-mCHO was 6 seconds and those of Mid and High Gel-CDH/HA-mCHO were 1 second, which is sufficiently rapid to manipulate the



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Gel-CDH/HA-mCHO in a clinical application. Gel-CDH/HA-dCHO, Gel-ADH/HA-dCHO10 and native Gel/HA-dCHO hydrogels were prepared as shown in Figure 3(A). Hydrazide modification dramatically altered the degradation time:

An Injectable Hydrogel with Slow Degradability Composed of Gelatin

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