Development of an Injectable Tissue Adhesive Hybrid Hydrogel for

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Development of Injectable Tissue-Adhesive Hybrid Hydrogel for Growth Factor-Free Tissue Integration in Advanced Wound Regeneration Md. Hasan Turabee, Thavasyappan Thambi, and Doo Sung Lee ACS Appl. Bio Mater., Just Accepted Manuscript • DOI: 10.1021/acsabm.9b00204 • Publication Date (Web): 15 May 2019 Downloaded from http://pubs.acs.org on May 15, 2019

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ACS Applied Bio Materials

(Revised Manuscript for the ACS Applied Bio Materials)

Development of Injectable Tissue-Adhesive Hybrid Hydrogel for Growth Factor-Free Tissue Integration in Advanced Wound Regeneration

Md. Hasan Turabee,† Thavasyappan Thambi,† and Doo Sung Lee*

School of Chemical Engineering, Theranostic Macromolecules Research Center, Sungkyunkwan University, Suwon 16419, Republic of Korea

†These

authors contributed equally to this work.

*Corresponding author: Doo Sung Lee, Ph.D. Tel.: +82-31-299-6851; Fax: +82-31-299-6857; e-mail: [email protected]

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ABSTRACT Development of injectable hydrogels with tunable multifunctional properties, for example, adhesive, elastic, swelling, biodegradable, and wound-healing properties that mimic the dynamic-healing process of skin regeneration, are currently a major challenge in tissue regeneration. Here, we report a fillable topical formulation of injectable gelatins (IGs) for the integration of advanced excisional wounds. Bioresorbable temperature-sensitive block copolymer was conjugated to the backbone of gelatins to form a dynamic coordinative network of IGs. The flexible IG precursors transformed to viscoelastic hydrogel at the subcutaneous tissues with cell affinity and tissue adhesive properties. The micro porous dynamic network of IGs allows access to nutrients, and recruits immune cells that accelerate neovascularization within hydrogel network. Combined with adhesive and neovascularization properties, the IGs alone exhibit accelerated wound healing in open wounds featured by skin appendages without scar formation. More remarkably, in an excisional wound (1 cm × 1 cm) animal model, the IGs promoted the wound healing as observed by the skin appendages. Furthermore, histological analysis demonstrated that IGs not only accelerates the rate of wound healing, but also promoted the quality of wound healing observed by collagen deposition and neovascularization. The in situ forming IGs with superior adhesive property can be considered as promising wound dressing materials for the treatment of multiple wounds, without the need for the encapsulation of biofactors or antibacterial metals. The IGs prepared in this study can also be employed to repair tissues or organs using minimal invasive administration. Key words: Injectable gelatins, biresorbable, poly(esters), wound healing, neovascularization.


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INTRODUCTION Hydrogels are three-dimensional porous materials that can be biologically engineered to mimic extracellular matrix-like biological soft tissues that have excellent potential in tissue repair and engineering applications.1-4 The interconnected porous biopolymer networks of hydrogels with high water content provide a suitable microenvironment to control biochemical and biophysical cues in tuning tissue reconstruction.5-8 Furthermore, the soft and pliable nature of hydrogels serves as matrices to match the defected area;9 subsequent transportation of nutrients through the porous materials accelerates the vascularization and formation of tissue constructs.2,

10-11

Traumatic injuries of soft tissues, including skin and muscle, are commonly encountered sport injuries, which can be treated with a water-swollen network of hydrogel.12-14 Although various hydrogels have been used in the regeneration of soft tissues, the high water content in the hydrogels makes them brittle materials.15 Therefore, there is a strong demand for hydrogels with superior mechanical properties. To improve the mechanical properties of hydrogels, tough hydrogels, such as interpenetrating network hydrogels and nanocomposite hydrogels with ductility, have been developed.16-20 However, the inability of tough hydrogels to mimic natural tissues severely limits their application in endogenous tissue-repair. On the other hand, fabrications of synthetic hydrogels that mimic biological tissue exhibit poor self-healing or adhesive properties, because when the hydrogels are subject to deformation or cyclic loading, the sacrificial covalent bonds in the hydrogel network cannot easily reintegrate.21-23 Hence, it is desirable to develop ideal tissue adhesives that employ natural or synthetic biocompatible and biodegradable polymeric materials for rapid tissue adhesion and wound closure, and protection against infections. In general, hydrogels applied for tissue repair or regeneration should possess excellent 3 ACS Paragon Plus Environment


tissue adhesion that allows open wounds or ruptured tissues to attach and accelerate tissue regeneration after hydrogel implantation.13-14, 24-26 Despite the biocompatibility of most hydrogels, poor cell affinity and lack of tissue adhesion properties cannot seal the wound tissues during surgical operation. In recent years, polydopamine-based mussel-inspired tissue adhesive hydrogels with structural similarity to marine mollusks have excellent adhesiveness to numerous surfaces, and outstanding cell affinity.16, 21, 27-29 Since polydopamine-based adhesive hydrogels were prepared through metal chelation of its highly reactive catechol groups, the toxicity of metals in these hydrogels might limit their biomedical applications.30 Recently, tremendous interest has been shown in natural protein-polymer conjugatebased hydrogels, owing to their excellent biological functions, such as cell interaction and cell signaling.31-36 The combination of proteins and synthetic polymers opens opportunities to produce smart hybrid materials, because of their highly tunable physiochemical characteristics.37 These biohybrid materials can combine the benefits from both components, and surmount the drawbacks of individual components. Proteins possess unique properties, such as specific binding to drugs or other biomolecules; whereas, synthetic polymers exhibit high thermal stabilities, controllable physiochemical and mechanical properties, and less immunogenicity.1 The fusion of appropriate hybrid polymers that undergo a sol-to-gel phase transition, upon exposure to various external stimuli, including temperature, pH, and enzymes, provide a minimally invasive injectable hydrogel platform that may be useful to enhance tissue repair and regeneration.34, 38-39 Inspired by the excellent transparency, biocompatibility, and extracellular matrix mimicking properties of gelatin,40-42 herein, we describe fillable topical formulation of injectable gelatins (IGs), containing an interconnected microporous structure, as a promising network with 4 ACS Paragon Plus Environment

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cell affinity, tissue adhesiveness, and the ability to heal excisional wounds without any exogenous growth factors. The limitation of temperature-sensitivity was endowed by the conjugation of copolymers that can spontaneously self-assemble in response to temperature. Particularly, temperature-sensitive gelatin-based hydrogel has intensively been studied because of their positive attributes for the delivery therapeutics such as minimal invasiveness.43-45 However, these polymers are almost always based on non-degradable poly(acrylates), which limits their biomedical applications. To surmount the issue of non-degradability, the temperature-sensitive in situ-forming hydrogel based on triblock copolymer poly(ε-caprolactoneco-lactide)-b-poly(ethylene glycol)-b-poly(ε-caprolactone-co-lactide) (PCLA-b-PEG-b-PCLA, called in short PCLA) has been extensively studied by our group.46 Controlling the molecular weight of PCLA blocks by tuning the CL/LA feed ratio, the thermal responsiveness, gelation window, rheological property, and biodegradability can be finely tuned. In this study, we synthesize PCLA-bearing gelatin (Gel-PCLA) conjugates that form spontaneously assembling hydrogel onto the excisional wounds, and seal the wounds for external contamination (Scheme 1). The extent of wound healing was tested in vivo.

EXPERIMENTAL PROCEDURES Materials PEGs (Mn=1,620 and 1,500 g/mol) were purchased from ID Biochem, Inc (Seongnam, Korea). DL-lactide

(LA), ε-caprolactone (CL), 4-dimethyl aminopyridine (DMAP), stannous octoate

(Sn(Oct)2), sodium hydrogen carbonate, methyl thiazolyl tetrazolium (MTT), gelatin from cold water fish skin (60,000 Da), and anhydrous solvents, including 1,4-dioxane, acetone, and chloroform, were supplied by Sigma-Aldrich (St. Louis, MO, USA). The N,N’-disuccinimidyl 5 ACS Paragon Plus Environment


carbonate was obtained from TCI (Tokyo, Japan). Phosphate-buffered saline (PBS) (1X, pH 7.4) was purchased from Life Technologies Limited (Paisley, UK). Preparation of Gel-PCLA conjugates The Gel-PCLA conjugates were prepared using a three-step process, as shown in Scheme 2. Preparation of PCLA block copolymer: The triblock PCLA copolymer prepared according to the previous report with slight modification.17 The block copolymer was prepared through ring opening polymerization of LA and CL with Sn(Oct)2 catalyst and PEG macroinitiator. In brief, PEG (10 g, 6.1 mmol) and Sn(Oct)2 catalyst (0.1 g, 0.25 mmol) were placed into a flask, and dried under high vacuum for 2 h at 110 °C. The temperature was then slowly reduced to 60 °C; subsequently, CL (19.5 mL, 166 mmol) and LA (7.5 g, 52 mmol) were slowly added, and dried for a further 1 h. Thereafter, to initiate polymerization, the temperature was gradually increased to 130 °C, and the polymerization continued under inert atmosphere. After 1 day, the reaction mixture was diluted using chloroform, and precipitated using n-hexane and diethyl ether (1:1, v/v) mixture. The precipitated block copolymer was filtered, and dried in a vacuum oven under reduced pressure at 25 °C for 2 days. Preparation of succinimidyl-activated PCLA (PCLA-NHS): The chain-end of triblock copolymer was activated for further conjugation with gelatin. To synthesize PCLA-NHS, PCLA (4.53 g, 1 mmol) was dissolved in 1,4-dioxane (40 mL) and acetone (40 mL). Thereafter, DMAP (0.26 g, 2.1 mmol) and N,N’-disuccinimidyl carbonate (0.77 g, 3 mmol) were added, and vigorously stirred at room temperature (RT) for 1 day. Then, the reaction mixture was concentrated using rotary evaporator to remove half of the solvent, and precipitated using excess diethyl ether. Precipitates of PCLA-NHS was then filtered and dried under vacuum. 6 ACS Paragon Plus Environment

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Preparation of Gel-PCLA hybrid conjugates: Gel-PCLA conjugates were synthesized by chemical conjugation of PCLA-NHS to the backbone of amine containing gelatin through amide bond formation. A series of Gel-PCLA conjugates were synthesized by simply regulating feed ratio of PCLA-NHS to gelatin. For the preparation of Gel-PCLA conjugates, gelatin (1 g, 0.016 mmol) was dissolved at 0 °C using sodium hydrogen carbonate buffer (pH 8.5), and stirred in an ice bath. Then different mole ratio (30, 40 and 50) of PCLA-NHS was added to the solution, and stirred under ice cold condition for 2 days. The concentration of PCLA-NHS was kept at 8 mg/mL. The crude mixture was then transferred to dialysis membrane tubes (MWCO: 12,000 Da), and dialyzed against excess distilled water at 0 °C for 2 days. Then, the purified conjugate solution was lyophilized to obtain Gel-PCLA conjugates. Characterization 1H

NMR spectra: The chemical composition of copolymers and Gel-PCLA conjugates were

determined by 1H NMR. 1H NMR spectra were obtained from Varian Unity Inova 500NB with 1 wt% sample concentration in CDCl3, D2O, or a mixture of D2O/DMSO-D6 (1:1, v/v).

Gel permeation chromatography (GPC): The GPC analyses (Agilent 1100 GPC system) were performed to determine molecular weight (Mn) and polydispersity index (Đ) of the PCLA copolymer. The GPC was performed with PCLA copolymer (1 wt%) sample solution using linear PEG standards. Chloroform (1.0 mL/min flow rate) was utilized as the mobile phase at 35 °C. FT-IR spectra: The FT-IR spectra were obtained using the KBr pellet technique. For the preparation of KBr pellets, copolymers mixed with KBr powder, and the pellets were made 7 ACS Paragon Plus Environment


under a pressure of 0.01 torr. Then, the FT-IR spectra were recorded in IFS-66/S FT-IR spectrometer, Bruker, USA. Scanning electron microscope (SEM): The porous structural property of the hydrogel was examined using SEM (SEM, JEOL JSM-6390). For SEM analysis, the hydrogel were recovered from the subcutaneous tissue of SD rats at various time points and freeze-dried. The solid hydrogels were coated with platinum to check the SEM image. Quantification of Gel-PCLA conjugation ratio The PCLA conjugation ratio in the Gel-PCLA conjugates was evaluated using 1H NMR spectra and 2,4,6-trinitrobenzene-sulfonic acid (TNBS) assay. In 1HNMR, phenylalanine was used as the internal reference to normalize the free amine signal. The conjugation ratio was calculated using the following formula:

Conjugation ratio = (1-

Lysine integration signal of Gel-PCLA conjugate ) 100 Lysine integration signal of Gel

The TNBS assay was also used to determine conjugation ratio of PCLA in the conjugates, based on the procedure developed by Habeeb using TNBS.47 The conjugation ratio was calculated by the ratio of free amine group in Gel-PCLA and gelatin. Each sample was checked three times. Sol-gel phase diagram In aqueous solutions, the flow (sol) and non-flow (gel) phase transition property of PCLA and Gel-PCLA was observed by tube inversion technique.48 Aqueous solutions of copolymers or conjugates at different concentrations were stirred vigorously at 0 °C for 12 h. The temperature of the water bath was raised from 0 to 65 °C with 4 °C interval, and the equilibration time for each point was kept for 20 min. The phase transition was examined by tilting sample vials, and 8 ACS Paragon Plus Environment

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observed the movements of the gel. Mechanical property measurement The change in viscosity with temperature was used to determine the rheological property of the copolymer and conjugate. A dynamic mechanical analyzer was used with 0.4 Pa controlled stress and 1 rad/s frequency in the oscillation mode (Bohlin Rotational Rheometer). The copolymer and conjugates solutions were placed between the 100 mm lower and 20 mm upper plate with 250 μm gap size with the temperature change from 5 to 70 °C. The complex viscosity was measured

with 2 °C min-1 of heating rate. In vitro cytotoxicity Human embryonic kidney 293T cells were used to examine the cytotoxicity of PCLA copolymers and Gel-PCLA conjugates. Fresh DMEM was used to culture the human embryonic kidney 293T cells with fetal bovine serum (FBS, 10% (v/v)) and penicillin-streptomycin (1%, (w/v)) in a humidified 5% CO2˗95% air at 37 °C. The 293T cells were detached, and then 1×104 cells/well were seeded in a flat-bottomed 96-well plate. After one day, PCLA copolymers and Gel-PCLA conjugates at different concentrations (from 50 to 2000 µg/mL) were exposed to human embryonic kidney 293T cells, and incubated at 37 °C for two days. Then, 20 μL of MTT solution (5 mg/mL) was added to the cells, and incubated for another 3 h. The culture medium was then discarded, the purple formazan crystals were dissolved using DMSO, and the absorbance at 490 nm was determined with the help of Microplate reader (BioTek, Korea). In vivo gelation and biodegradation Four weeks old male Sprague-Dawley (SD) rats purchased from the KRIBB (Daejeon, Korea) were employed in this study to examine in situ gelation, biodegradation, and bioresorbable 9 ACS Paragon Plus Environment


properties of PCLA copolymer and Gel-PCLA conjugates. The live animal experiments were approved and performed according to the guidelines of Sungkyunkwan University. To evaluate in vivo gelation property, PCLA copolymer or Gel-PCLA conjugate solutions (300 µL, 20 wt%) were injected into the dorsal region of SD rats. The SD rats were sacrificed at pre-defined time interval to collect the gel, and were photographed to determine the state of hydrogels. The hydrogels were flash-frozen using liquid nitrogen, and lyophilized to check the microporous structure of hydrogels. Furthermore, the remaining hydrogels were weighed, and the biodegradation rate was calculated using the mass loss method. Adhesion tests Adhesion properties of the polymeric hydrogels were measured by lap-shear test. Skins collected from SD rats were used as a substrate to examine the adhesive strength of the hydrogels. A piece of skin (dimension of 3 cm × 1 cm × 0.85 mm) was stuck on another skin with the help of 20 wt% (20 µL) of PCLA copolymer or Gel-PCLA conjugate hydrogel formulation. Control sample without hydrogel was used to compare the adhesion of hydrogels. The adhesive strength was determined using a universal testing machine (UTM model 5565, UK) under 250 N loads with a pulling rate of 1 mm min-1. Wound healing Two types of wounds were developed on the dorsal region of rats to investigate the wound healing property of Gel-PCLA conjugate hydrogel. Firstly, 1 cm long open wounds were created, after they were anesthetized with pentobarbital. The rats were splits into three different groups (n=3), (i) control group, (ii) PCLA hydrogel, and (iii) Gel-PCLA hydrogel. In the control group, the wounds were kept untreated and covered by simple surgical suturing. On the other hand, in 10 ACS Paragon Plus Environment

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the hydrogel groups, the hydrogel precursor was transferred to syringe (26 G), and injected to the wound area. The extent of wound healing was measured and photographed at different time points for seven days. Thereafter, the rats were sacrificed, and skins were harvested from wound sites; the harvested skins were fixed on glass slides, and stained for histology analysis. A full-thickness cutaneous wound model was created as another type of wound model. To create the excisional wound, the rats were anesthetized, and the dorsal region of the rats was shaved and disinfected with 70% aqueous ethanol. At the dorsal region, a 1 cm × 1 cm excisional wound was created using surgical instruments. The control group was kept untreated, and wrapped with transparent plastic mold. Other group animal wounds were filled with hydrogels, and wrapped with transparent plastic mold, to prevent interaction with dirt and dust particles. At different time intervals, the wounds were photographed, and wound contraction was measured until complete re-epithelialization. The wound healing progress was investigated by histological analysis. The skin tissue collected from the rats was fixed using 10% formalin, and embedded in paraffin. Skin tissues embedded in paraffin were then sliced with thickness of 4 µm, and used to stain hematoxylin & eosin (H&E) and Masson's trichrome (MT). The slides were then visualized with the help of microscope. To examine the integrity of wounds treated with hydrogels or untreated control group, the rats were sacrificed after complete healing, and the dorsal skin of the healed wounds was harvested with the dimensions of 3 cm × 1 cm × 0.85 mm. The collected skins were kept in PBS, prior to measurement. The UTM with pulling rate of 1 mm min-1 was used to examine wound breaking strength under a load of 250 N.



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RESULTS AND DISCUSSION Preparation and characterization of Gel-PCLA conjugates Scheme 2 shows the preparation of Gel-PCLA copolymer conjugates. First, PCLA block copolymer was prepared through the one-pot ring-opening polymerization of LA and CL in the presence of PEG. Figure S1A shows the 1HNMR spectra of PCLA copolymers, which reveal the major characteristic peaks of PEG, PCL, and PLA. Furthermore, 1H NMR spectra used to calculate degree of polymerization of PCL and PLA blocks by comparing the integration of methylene and methine protons of PCL and PLA at 4.07 ppm and 5.15 ppm, respectively, with methylene protons of PEG at 3.64 ppm. The composition of PCLA copolymer was varied by regulating the feed ratio of CL and LA to PEG. As expected, the copolymer weight of PCLA significantly increased as the feed ratio of CL and LA increased (Table S1). The Mn and Đ of PCLA block copolymers were calculated using GPC (Figure 1A). It is noteworthy that the Mn and Đ obtained from GPC were close to those of

1H

NMR, indicating the uniform

polymerization of PCLA block copolymers. FT-IR spectra of copolymerized PCLA shows the CO-C ether stretch, C=O stretch, and alkane stretching vibrations at 1,100 cm-1, 1,740 cm-1 and 2,860-3,000 cm-1, respectively (Figure 1B). After chain-end activation, N-O stretching was found at 1,660 cm-1 confirming NHS activation. Then, the NHS-activated PCLA copolymer was prepared for conjugation with gelatin. The esterification reaction of PCLA with N,N’-disuccinimidyl carbonate may activate either mono or both termini of hydroxyl groups. However, the controlled addition of N,N’disuccinimidyl carbonate to PCLA mainly provides mono-esterified PCLA. This was clearly observed from the 1H NMR spectrum (Figure S1B). The chain end activation of the PCLA 12 ACS Paragon Plus Environment

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copolymers was estimated using integral values of a PEG peak at 3.64 ppm to that of succinimide peak at 2.82 ppm. Finally, the Gel-PCLA conjugates were synthesized by the chemical conjugation of PCLA-NHS to the backbone of gelatin through amide bond formation. The conjugation of GelPCLA was confirmed using 1H NMR spectra (Figure S1C and D). In addition to the gelatin peaks, the new characteristic peak at 3.66 ppm indicated the effective conjugation of PCLA to gelatin. Furthermore, new characteristics peaks at 1.45, 3.65 and 5.1 ppm corresponding to CL, PEG, and LA also appeared. A series of conjugates with different composition of PCLA and gelatin were prepared by regulating feed ratio of gelatin and PCLA (Table S2). The 1H NMR spectra result suggests that PCLA-NHS was successfully conjugated to gelatin. The successful synthesis of Gel-PCLA conjugates was further investigated using water GPC (Figure 1C). As expected, the retention time of Gel-PCLA conjugate is shorter than that of free gelatin. This is mainly due to the conjugation of PCLA to gelatin increasing Mn and allowing early elution. FT-IR spectra were used to further characterize the characteristics peaks of Gel-PCLA (Figure 1B). The characteristic absorption peaks of gelatin at 1,640, 1,538, and 1,175 cm-1 corresponds to amide-I, amide-II, and amide-III, and a weak band appeared at 2,850-3,000 cm-1. For the Gel-PCLA conjugate, the new peak appeared at 1,100 cm-1 for the C-O-C bond, and, 1,740 cm-1 for C=O stretching. Furthermore, the band at 2,850-3,000 cm-1 became strong, indicating the successful conjugation of PCLA with gelatin. Based on the Mn and easy solubility, we decided to use PCLA-1 for the conjugation with gelatin. Table S2 shows the quantification of Gel-PCLA. The PCLA compositions in the conjugates calculated using 1H NMR spectra, and TNBS assay, were identical. Sol-to-gel phase transition of IGs 13 ACS Paragon Plus Environment


Sol-to-gel phase transition pattern of PCLA block copolymers with varying hydrophobic and hydrophilic balance was examined. Regardless of the hydrophobic and hydrophilic balance, all the PCLA copolymers showed temperature-sensitivity and the freely flowing PCLA copolymer sols at low temperature was transform to stable hydrogel. Although the PCLA-2 forms stable gel, the gelation time is long, and not suitable for in situ-formation of injectable gel based biomedical application. On the other hand, due to the higher hydrophobic balance the solubility of PCLA-3 was poor, and took a long time for the aqueous dissolution. Therefore, for the hybrid copolymer conjugate preparation, the PCLA-1 copolymer with a hydrophilic and hydrophobic ratio of 1:1.8 was used. A series of Gel-PCLA conjugates were synthesized by varying the feed of gelatin to PCLA. The Gel-PCLA conjugates with low PCLA content could not form stable IG in vitro, which indicated the higher hydrophilicity of conjugates. In particular, the solubility of gelatin at high temperature and poor hydrophobic segments in the conjugates restrict the gelation. Therefore, Gel-PCLA1 and Gel-PCLA2 conjugates could not be gelled at the physiological condition. Interestingly, the Gel-PCLA3 could form stable gel at body temperature as low as 18 wt% concentration in PBS (Figure 1D). The higher number of PCLA block polymer in the conjugates creates a hydrophobic balance to maintain gel state at the body temperature (37 °C). In addition, the existence of functional groups in Gel-PCLA conjugate can make hydrogen bonding that contributes to the improvement of gel strength. Overall, the Gel-PCLA conjugate (shortly IG conjugate) hydrogel region tended to shift towards a lower temperature region than the PCLA gel region. This is mainly due to the solubility of gelatin at high temperature and formation of gel at low temperature, which lowers the sol-gel window. Figure 1E clearly shows that the gel boundary of IG is retained within the body condition, which implies the formation of 14 ACS Paragon Plus Environment

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gel in vivo. Viscosity measurements An appropriate change in the mechanical properties of copolymers is an important prerequisite for the in situ formation of IGs. For IGs, the change in viscosity of copolymer in response to the physiological stimuli changes is a big challenge. It is important that the copolymers should be flowable at low pH or temperature that could form gel, while being injected into warm-blooded animals. The PCLA copolymers and IG precursor possess inherent responsiveness to temperature. At low temperature from 0 to 23 °C, both aqueous copolymers exhibited good flowability; the viscosity of PCLA copolymer and IG precursor were recorded at about 0.6 Pa.s and 3.9 Pa.s, respectively (Figure 1F), which allowed the injection of copolymers using small gauge needles. At sol state, the low viscous property of the copolymers allows easy mixing of therapeutic agents or cells and could be used as matrices to culture and grow cells. As the temperature rose, the viscosity of the copolymer dramatically increased over 2,000 Pa.s. The viscosity of IG is slightly lower at the high temperature than the pristine PCLA hydrogel. This is mainly due to the solubility of gelatin that caused decrease in viscosity. The rheological properties of copolymer were similar to that of sol-to-gel phase transition obtained at different temperature, as observed in Figure 1E. At high temperature, the copolymer makes hydrogel through hydrophobic interaction, and molecular entanglement between copolymer is responsible for gelation. The free-flowing IG conjugates at the low temperature and the gelation at the physiological condition could be a candidate material for in vivo applications. In vitro biocompatibility To assess the potential biomedical application of the prepared PCLA copolymers and IG 15 ACS Paragon Plus Environment


conjugates, the biocompatibility of the samples was firstly tested using human embryonic kidney 293T cells. As shown in Figure 2A, the 293T cells cultured with different concentration of either PCLA copolymers or IG conjugates exhibited high cell viability. In particular, the cell viability exceeded 95% even at 2,000 µg/mL of IG conjugate exposure, which confirmed the superb biocompatibility of the prepared conjugates. The high viability of the cells cultured with IG conjugate group was caused by the presence of gelatin polymer, a natural translucent polymer obtained by the hydrolysis of collagen.49 This result indicates the biocompatibility of IG conjugates, which warrants the usage of the materials in tissue engineering applications. In vivo gelation and biodegradation To determine the gelation and the biodegradable properties of hydrogels in vivo, the PCLA block copolymer and IG conjugate copolymer precursor were subjected to subcutaneous administration into the dorsal region of the SD rats. Regardless of formulations, nodules were formed right after the injection, which indicated the in situ gelation similar to that of in vitro gelation observed in Figure 2B. Ten minutes after post-injection of formulations, rats were sacrificed, and the nodules were retrieved to examine the stability of the hydrogel. The PCLA copolymer and IG conjugate formed stable gel with a well distinguished shape, without inducing any inflammatory response. The recovered hydrogels were subjected to SEM analysis, to examine the porous property of hydrogels. As shown in Figure 2C, it is clear that the porous properties of hydrogels were retained, even after applying shear stress. However, the pore sizes varied, depending on the formulation. The pore size of the PCLA hydrogel was smaller than that of the IG hydrogel. This might be due to the higher hydrophobic balance of PCLA copolymer making a strong network, and making the mesh size smaller. On the other hand, the presence of hydrophilic gelatin in the IG hydrogel restricted the strong inter-/intramolecular interaction, resulting in higher mesh size. 16 ACS Paragon Plus Environment

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The pore size of the copolymers was increased one week after post-injection. In particular, IG hydrogels were found to degrade faster than the PCLA copolymer hydrogel, which indicated that the natural polymer containing hybrid hydrogel network exhibits good biodegradability in vivo. Furthermore, the extent of biodegradability, inflammation, and necrosis behavior of hydrogels was examined at different time points after the post-injection of the formulation. The hydrogel was withdrawn from rats at different time intervals, and biological response and degradation rate were examined. As shown in Figure 2D, the degradation of hydrogels started from the first week; it should be noted that the degradation rate of IG hydrogels was slightly faster than that of the synthetic PCLA hydrogels. At the second week, the remaining weight of PCLA copolymer hydrogel and IG conjugate hydrogel were found to be 76% and 64%, respectively (Figure 2E). Thereafter, the degradation rate was accelerated, indicating that the hydrogel networks loosened their integrity after a certain amount of network degradation. At the fourth week, less than 25% of IG hydrogels remained, and had lost their integrity; and the remaining hydrogel was completely covered with blood vessels, and showed a tissue-like network. It is noteworthy that subcutaneous implantation of IG hydrogels exhibited no signs of inflammation or infection. These results prove the in vivo biosafety of the hydrogels, and their potential application of IGs as implantable biomaterials. Tissue adhesiveness In general, the three-dimensional polymer network of hydrogels with high water content was reported to have poor cell affinity and adhesive properties. Recently, tough hydrogels, such as mussel-inspired hydrogels, have been developed to improve the adhesion properties in various surfaces, including biological tissues and hydrophobic surfaces. The tissue adhesive property of PCLA copolymer hydrogel and IG conjugate hydrogel was examined by attaching two separate 17 ACS Paragon Plus Environment


skins of rats (Figure 3A). The hydrogels resembling the property of biological tissues showed a bridging effect between the two skins, and the adhesiveness could be maintained by applying mechanical stress. However, the control group with or without PBS failed to join the skins. The excellent adhesive property of hydrogels without any external stimuli has been ascribed to the interfacial hydrogen bonding between hydrogels and skin tissues. As shown in Figure 3B, the adhesive strength of Gel-PCLA conjugate hydrogel was higher than those of PCLA copolymer hydrogel, which indicated that the presence of natural gelatins influence tissue adhesiveness in hydrogels. Furthermore, the effective adhesiveness of IG conjugate hydrogel also influenced the displacement property, which was larger than those of PCLA hydrogel, due to the presence of free amine groups having effective interaction with skin tissues (Figure 3C). The adhered hydrogel on skin could be easily peeled off, without leaving any residual hydrogel on the rat skin. The free amine groups on the surface of IG conjugate hydrogels could form ionic interaction with contact surfaces. In vivo wound healing Wound regeneration is a highly specific biological process of growth and generation of new tissues, which progressed through a series of interdependent and overlapping stages in which cellular and matrix components organize together for integration of damaged skins.50-52 In general, the wound regeneration process comprised of five overlapping complex biochemical and cellular phases, which are described as haemostasis, inflammation, migration, proliferation, and maturation phases.53 Numerous cells (fibroblast, macrophages, and endothelial cells) and cytokines (interleukin-1 and transforming growth factor-β1) have involved in wound healing process through the complex signaling pathways.54-55 However, various factors disrupt this healing process and wound persists for long-term. Therefore, in situ-forming hydrogels have 18 ACS Paragon Plus Environment

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been developed as wound dressing materials to improve the cutaneous wound healing. The tissue adhesion effect of IG conjugate hydrogel was further examined using cutaneous wound model in vivo. Firstly, the adhesive and sealant properties of hydrogels were examined using the simple cutaneous wound model. The simple liner open wound (1 cm) was treated with PBS, PCLA copolymer hydrogel, and IG conjugate hydrogel, and the wound closure rate was monitored (Figure 4A). As observed in Figure 4B, the animal groups treated with hydrogels showed better wound closure, compared with the PBS control group. In particular, the adhesive property of hydrogels effectively sealed open wounds, and the healing started right after in situ injectable gel administration. One-time administration of hydrogel formulations shows statistically significant wound contraction, when compared with PBS control (Figure 4C). It is noteworthy that IG conjugate hydrogel was able to heal the wound effectively, and achieved a smooth appearance that was similar to that of normal skin. In contrast, wounds treated with PCLA copolymer hydrogel and PBS control group showed a reddish surface. For hydrogel groups, wound contraction started to appear on the first day. Within 7 days, the IG conjugate hydrogel showed complete healing of wounds. The histological examination of H&E stained skin tissues treated using IG conjugate hydrogel showed complete re-epithelialization (Figure 4D). During the skin remodeling, IG conjugate hydrogel significantly improved the quality of wound healing with higher density of epithelial cells. Furthermore, the Masson’s trichrome stained skin tissue of the hydrogel group exhibited well-organized collagen fibers (Figure 4E). The enhanced wound healing property of our formulation encourages us to investigate its effectiveness in the excisional wound model. To investigate the efficiency of IG conjugate hydrogel in full-thickness wound healing, a 1cm × 1 cm excisional wound model was developed on the dorsal region of rats (Figure 5A). To 19 ACS Paragon Plus Environment


validate our hypothesis, in situ-forming hydrogel formulations were intradermally injected, and PBS was used as control. Excisional wounds treated with IG conjugate hydrogel exhibited accelerated wound closure compared with the PCLA hydrogel and PBS treated wounds (Figure 5B and C). The wound contraction rate of IG conjugate hydrogel at day 7 was 1.4- and 2-fold higher than that of the PCLA hydrogel and control groups. The H&E staining section shows that IG conjugate hydrogel filled wounds accelerated healing rate compared with PCLA hydrogel and control groups, as observed by the formation of granular tissues (Figure 5D). Furthermore, Masson’s trichrome assay shows the dermis deposition, which initially supports cell migration (Figure 5E). In particular, the IG conjugate hydrogel filled wounds regenerated thicker dermis layer than the PBS-treated wounds. Additionally, collagen remodeling is more homogenous in the IG conjugate hydrogel filled wounds. To further confirm the wound healing effectiveness, the healed wounds were recovered and subjected to breaking analysis (Figure S2). The wound breaking strength of IG conjugate hydrogel was significantly higher than that of control and PCLA hydrogel groups. These results suggest that IGs prepared in this study could be a promising fillable topical formulation for excisional wounds and develop microenvironment for neovascularization.

CONCLUSIONS In summary, we synthesized an in situ-forming IG hydrogel platform for adhesiveness and excisional wound healing. The unique properties of IG hydrogels, including their porous structure, biocompatibility, biodegradability, and in situ gelling ability, could be exploited as dressing materials or wound sealant for excisional wounds. The IG hydrogels not only adhered on skin tissues, but also effectively sealed the wounds, accelerated cutaneous wound healing, and 20 ACS Paragon Plus Environment

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promoted tissue regeneration in the wound area, without any exogenous growth factors or toxic metals. The fillable IG hydrogels used in this study is not only limited to cutaneous wound healing, but is also applicable in other wound healing or tissue regeneration applications, where in situ gel formation with tissue adhesive ability matters.

ASSOCIATED CONTENT Supporting Information The supporting Information is available free of charge on the ACS Publication website. 1H

NMR spectra, wound breaking strength, and Table S1 and S2.

Notes The authors declare no competing financial interest.

ACKNOWLEDGMENTS This research was supported by the Basic Science Research Program through a National Research Foundation of Korea grant funded by the Korean Government (MEST) (20100027955), and the National Research Foundation of Korea (NRF) funded by The Ministry of Science, ICT & Future Planning (NRF-2017R1D1A1B03028061).

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Gelatin

PEG

PCLA

Macrophages

Fibroblasts

Scheme 1. Schematic illustration of formation of IG hydrogels that could fill the full-thickness wound by simple extrusion of IG gelatin precursors through hypodermic needles. The controlled degradation of IG hydrogel on the wound site subsequently accelerates the wound healing cascade and promotes tissue regeneration.



PEG

LA

CL

PCLA N,N’-disuccinimidyl carbonate

PCLA-NHS

Gel-PCLA 26 ACS Paragon Plus Environment

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Scheme 2. Synthetic route to prepare Gel-PCLA conjugates.

A

D

B

PCLA 23 oC

Gel-PCLA 23 oC

37 oC

37 oC

C

E

F

Figure 1. (A) GPC traces of PCLA triblock copolymers. (B) The FT-IR spectrum of PCLA copolymers and Gel-PCLA conjugates. (C) HPLC of Gel-PCLA conjugates. Native gelatin was used as control. (D) Sol-to-gel phase transition photographs of PCLA copolymers and Gel-PCLA conjugates at different temperatures. (E) Sol-to-gel phase transition diagram of PCLA copolymers and Gel-PCLA conjugates by varying copolymer weights. (F) Viscosity of PCLA copolymers and Gel-PCLA conjugates with change in temperature.



A

10 min

1 week

2 weeks

PCLA

3 weeks

Gel-PCLA

C

4 weeks

PCLA

Gel-PCLA

E

PCLA

D

B

Gel-PCLA

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

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Figure 2. (A) Cytotoxicity of human embryonic kidney 293T cells co-cultured with PCLA copolymers and Gel-PCLA conjugates at different concentrations. (B) In vivo gelation of PCLA copolymers and Gel-PCLA conjugates into the back of SD rats. For this study, 300 µL (20 wt%) of copolymers was injected, and the gels were collected 10 min after post-injection. (C) SEM image of hydrogels recovered at 10 min and 2 weeks after post-injection. (D) Biodegradation pattern of hydrogels at different time points of post-injection. (E) The rate of degradation of hydrogels determined using the mass loss method. The error bars in the graph represent standard deviations (n=3).


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A

Control

PCLA

Gel-PCLA

B

C

Figure 3. (A) Aqueous solution of PCLA and IG hydrogels were used to adhere rat skins. PBS was used as control to compare the adhesive effect of hydrogels. The quantitative analysis of (B) force displacement curve and (C) adhesive strength of PCLA and IG hydrogels was estimated through universal testing machine.



A

B

PCLA

Gel-PCLA

3 Days

1 Day

0 Day

Control

5 Days

C

7 Days

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

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D

Control

PCLA

E

Gel-PCLA

Control

10 mm

PCLA

Gel-PCLA

Figure 4. IG-based adhesive hydrogel as fillable wound dressing materials. (A) Schematic of simple open wound healing using IGs. (B) Photograph of wounds at 0, 1, 3, 5, and 7 days post healing. (C) Wound closure kinetics quantification of PBS, untreated, PCLA hydrogel, and IG hydrogels. (D and E) H&E and Masson’s trichrome staining of healed wound tissues harvested at day 7; the bottom panels are corresponding magnified images. The scale bars in the top panels are for 100 µm, and the bottom panels are for 10 µm.


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Control

C

PCLA

Gel-PCLA

0 Day

A

1 Day

Hydrogel Macrophages

B

3 Days

ECM protein (collagen) Fibroblasts Neutrophils

PCLA

Gel-PCLA

5 Days

Control

7 Days

D

9 Days

E

14 Days

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


10 mm

Figure 5. IG-based hydrogel as fillable excisional wound dressing materials. (A) Schematics of excisional wound healing using IGs. (B) Photographs of excisional wounds after 0, 1, 3, 5, 7, 9 and 14 days of post healing. (C) Quantification of wound closure kinetics of untreated, PBS, PCLA hydrogel, and Gel-PCLA conjugate hydrogels. (D and E) H&E and Masson’s trichrome staining of healed wound tissues harvested at day 14; the bottom panels are corresponding magnified images. The scale bars in the top panels are for 100 µm, and the bottom panels are for 10 µm.



TOC

Gelatin

PEG

PCLA

Macrophages

Fibroblasts


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Development of an Injectable Tissue Adhesive Hybrid Hydrogel for

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