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Bioinspired pH- and Temperature-Responsive Injectable Adhesive Hydrogels with Polyplexes Promotes Skin Wound Healing Thai Minh Duy Le, Huu Thuy Trang Duong, Thavasyappan Thambi, V.H. Giang Phan, Ji Hoon Jeong, and Doo Sung Lee Biomacromolecules, Just Accepted Manuscript • DOI: 10.1021/acs.biomac.8b00819 • Publication Date (Web): 13 Jul 2018 Downloaded from http://pubs.acs.org on July 14, 2018

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(Revised Manuscript for Biomacromolecules)

Bioinspired pH- and Temperature-Responsive Injectable Adhesive Hydrogels with Polyplexes Promotes Skin Wound Healing

Thai Minh Duy Le,†,‡ Huu Thuy Trang Duong,†,‡ Thavasyappan Thambi,†,‡ V.H. Giang Phan,# Ji Hoon Jeong*,⊥ and Doo Sung Lee*,†

†

School of Chemical Engineering, Theranostic Macromolecules Research Center, Sungkyunkwan University, Suwon 440-746, Republic of Korea # Faculty of Applied Sciences, Ton Duc Thang University, Ho Chi Minh City 70000, Vietnam ⊥School of Pharmacy, Theranostic Macromolecules Research Center, Sungkyunkwan University, Suwon 440-746, Republic of Korea

‡

These authors contributed equally to this work.

*Corresponding authors: Doo Sung Lee, Ph.D. Tel.: +82-31-299-6851; Fax: +82-31-299-6857; e-mail: [email protected] Ji Hoon Jeong, Ph.D. Tel.: +82-31-290-7783; Fax: +82-31-292-8800; e-mail: [email protected]

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ABSTRACT Despite great potential, the delivery of genetic materials into cells or tissues of interest remains challenging owing to their susceptibility to nuclease degradation, lack of permeability to the cell membrane, and short in vivo half-life, which severely restrict their widespread use in therapeutics. To surmount these shortcomings, we developed a bioinspired in situ-forming pH- and temperature-sensitive injectable hydrogel depot that could control the delivery of DNA-bearing polyplexes for versatile biomedical applications. A series of multiblock copolymer, comprised of water

soluble

poly(ethylene

glycol)

(PEG)

and

pH-

and

temperature-responsive

poly(sulfamethazine ester urethane) (PSMEU), has been synthesized as in situ-forming injectable hydrogelators. The free-flowing PEG-PSMEU copolymer sols at high pH and room temperature (pH 8.5, 23 oC) were transformed to stable gel at the body condition (pH 7.4, 37 oC). Physical and mechanical properties of hydrogels, including their degradation rate and viscosity, are elegantly controlled by varying the composition of urethane ester units. Subcutaneous administration of free-flowing PEG-PSMEU copolymer sols to the dorsal region of SpragueDawley rats instantly formed hydrogel depot. The degradation of the hydrogel depot was slow at the beginning and found to be bioresorbable after two months. Cationic protein or DNA-bearing polyplex-loaded PEG-PSMEU copolymer sols formed stable gel and controlled its release over 10 days in vivo. Owing to the presence of urethane linkages, the PEG-PSMEU possesses excellent adhesion strength to wide range of surfaces including glass, plastic and fresh organs. More importantly, the hydrogels effectively adhered on human skin and peeled easily without eliciting an inflammatory response. Subcutaneous implantation of PEG-PSMEU copolymer sols effectively sealed the ruptured skin, which accelerated the wound healing process as observed by the skin appendage morphogenesis. The bioinspired in situ-forming pH- and temperature-


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sensitive injectable adhesive hydrogel may provide a promising platform for myriad biomedical applications as controlled delivery vehicle, adhesive and tissue regeneration.

Key words: Injectable hydrogel, polyplex, urethane ester, DNA, tough adhesive and wound healing.

INTRODUCTION Pharmaceutical and medical development in recent years has resulted in a flurry of innovation and drug approvals across several diseases. Among them, gene therapy has been widely researched, and has shown many promising results in treating many diseases, such as cancer, genetic disorders, and tissue regeneration.1-5 Despite these potential applications, clinical applications of gene therapy still suffer from several challenges, which include enzymatic susceptibility, short plasma half-life, and more importantly, reticuloendothelial system clearance results in low therapeutic concentrations at the target tissue.5-8 In recent years, the development of smart materials, which can control delivery of therapeutics in a spatial and temporal controlled pattern has received great attention.9, 10 There are multiple potential materials that can be utilized as a controlled delivery system, such as nanoparticles, polymeric micelles, dendrimers, and liposomes.11-13 These delivery systems possess individual properties, and demonstrate their advantages as well as limitations in different circumstances. However, owing to the hydrophilic, biomacromolecular and anionic characteristics of genetic material, it is challenging to cross the negatively charged cell membrane. To surmount this shortcoming, the surface of the nanocarriers has been modified with 2’-fluoro phosphorothiolate or 2’-O-methyl phosphorothiolate that could



effectively facilitate cell uptake.14,

15

Hence, the preparation of nanocarriers with appropriate

polymers, such as cationic polymers, may also facilitate the cell uptake. Generally, the development of DNA delivery vehicles has mainly focused on nanocarriers that could be systemically administered.16-18 Such delivery vehicles have been frequently prepared between cationic polymer and DNA, termed polyplex, which protects the genes from proteolytic and enzymatic degradation.19-21 Furthermore, the cationic polymer facilitates the cellular uptake of genetic materials.21, 22 However, because of their easy dispersion, systemically administered polyplexes have very short retention time and subsequently provide low therapeutic effect. Therefore, an optimal formulation could transmit high dose of genetic materials to target tissues in a sustained manner. Among promising systems, injectable hydrogels are presently considered as one of the most effective designs for controlled delivery application.23-26 Hydrogels are hydrophilic polymer matrixes that in their swollen state absorb a very high amount of water or biological fluids.4, 27, 28 Owing to the high water absorbing property of hydrogels, they can simulate the properties of biological tissues, resulting in superb biocompatibility.29 For controlled delivery, in addition to high biocompatibility, the hydrogels require the ability to encapsulate therapeutic agents with high loading content, and release therapeutic agents in a controllable manner, with tunable stability and biodegradability.30-32 Injectable hydrogels have the considerable advantages that they do not require surgical implantation, and the release of encapsulated proteins is regulated by the degradation pattern of hydrogel matrix, as well as by simple diffusion.33-37 Although injectable hydrogels for gene delivery has long been used, the genetic materials have been physically loaded through ionic interaction and release in a native form, which exhibited short in vivo half-life and lead to the low transfection efficiency.38 To utilize the properties of genetic materials without sacrificing


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their therapeutic benefits, a formulation could load genetic materials by protecting their surface properties and release them in a sustained manner. Such released materials should direct the genetic materials to reach the target site effectively. Therefore, the genetic materials need to be wrapped with appropriate non-toxic materials, and loaded into the formulation. In this study, for the complexation we first prepared a non-toxic cationic polymer, and prepared a polyplex by simple mixing with genetic materials. Since the cationic polymer has a proton buffering capacity, the polyplex can effectively transfuse the genetic materials into cells, and effectively deliver to maximize the therapeutic effect. Particularly, various in vivo gene transfer strategies have been employed to repair cutaneous wounds.39 Delivery of genes encoding specific antigenic determinant resulted in recruitment and processing of dermal cells, which induced biological response, stimulate tissue repair and have the potential to enhance tissue regeneration. Polymeric adhesives based on hydrogels can be exploited as surgical sealants for reconnecting ruptured tissues.40-42 Unlike conventional dressings, hydrogel-based adhesive dressings allow effective permeation of oxygen to the wound surface and cool the wound site, resulting pain relief for patient.43-45 Thus far, various wound dressings, such as sutures, staples, foam or electrospun, have widely been used to connect wound tissues and restore tissue functions. Particularly, adhesive hydrogel possess unique properties such as in situ encapsulation of genes and growth factors, ability to fill irregular shape wounds, and adhere to wounds that protects the wound site from external environment.46-48 It has been found that polymeric adhesive hydrogels often have low adhesion, inadequate mechanical strength, and formation of toxic degradation prevent clinical translation. Hence, engineering of in situ forming adhesive hydrogel-based wound dressing with high adhesive property similar to that of mussels, and biocompatibility is highly desirable. However, bioinspired injectable hydrogel adhesives that



could reconnect wound tissues for recovering tissue structure and function remains elusive. To test this hypothesis, we synthesized a new type of copolymers constitute of poly(ethylene glycol) (PEG) and poly(sulfamethazine-ester-urethane) (PSMEU), which could form bioinspired adhesive hydrogel that could be applied for versatile biomedical application such as controlled delivery system for polyplexes, bio adhesive and surgical sealants for wound healing (Scheme 1).

EXPERIMENTAL PROCEDURE Materials PEG (Mn = 2,050 g/mol), dibutyltin dilaurate (DBTL, 95.0%), ε-caprolactone (CL, 97.0%), stannous octoate (Sn(Oct)2, 95.0%), α-thioglycerol (TH, 97.0%), 1,6-hexamethylene diisocyanate (HDI, 98.0%), sulfamethazine (SM, 99.0%), L-glutamic acid γ-benzyl ester (BLG, ≥99.0%), butylamine (99.5%), diethylenetriamine (99.0%), branched polyethylenimine (bPEI, Mw=25,000 Da), thiazolyl blue tetrazolium bromide (MTT, 98.0%) and various anhydrous solvents used in the experiments, were obtained from the Sigma-Aldrich Co. (St. Louis, MO, USA). Acryloyl chloride, triphosgene, and fluorescein 5-isothiocyanate (isomer I) were bought from the Tokyo Chemical Industries (TCI, Tokyo, Japan). Penicillin-streptomycin solution, 4′,6diamidine-2′-phenylindole dihydrochloride (DAPI), dulbecco’s modified Eagle’s medium (DMEM), fetal bovine serum (FBS), and trypsin-EDTA cell culture reagents with high purity were bought from Invitrogen (Carlsbad, CA). The remaining reagents were of analytical grade and utilized as received.

Synthesis of PEG-PSMEU copolymers 6 ACS Paragon Plus Environment

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Figure S1 shows the four-step synthetic process of the PEG-PSMEU copolymer. Synthesis of thioglycerol-functionalized sulfamethazine (SSM). SSM was synthesized through Michael-addition reaction with acrylate and thiol groups of sulfamethazine acrylamide (SMA) and α-thioglycerol, respectively. In an equimolar mixture of DMF and MeOH (270 mL), SMA49 (16.52 g, 0.05 mol) and α -thioglycerol (5.95 g, 0.055 mol) were dissolved and simply heated the reactants for 36 h at 50 °C. Then, DMF and MeOH were evaporated and the crude residue was dissolved in CHCl3. Finally, SSM was precipitated by adding excess amount of cold diethyl ether, filtered and dried in a vacuum for two days. Synthesis of ε-caprolactone-terminated SSM (SSM-CL). The monomer SSM-CL was prepared by ring-opening reaction of CL with SSM. SSM (4.38 g, 0.01 mol) and Sn(Oct)2 (0.06 g, 1.5% SSM mol) were added to the flask and then changed with CL (1.05 g, 0.01 mol) under a nitrogen atmosphere. Then, to remove the moisture, the added reactants were vacuum dried for 1 h. Thereafter, dioxane (50 ml) was added to completely dissolve reactants and continued the reaction at 110 °C for one day. The synthesized SSM-CL was then precipitated by the addition of diethyl ether, followed by filtration and drying in vacuum for two days. Synthesis of PEG-PSMEU copolymer. Based on our previously reported procedure,49 a series of PEG-PSMEU copolymer was prepared. In brief, PEG (2.050 g, 1 mmol) and DBTL (0.0082 g) were transferred to a schlenk flask and vacuum was applied for 2 h at 100 °C to remove moisture. Thereafter, SSM-CL (2.76 g, 5 mmol) was added to the flask, and continued the drying for another 30 min. The flask was then filled with nitrogen and dissolved the reactants using anhydrous THF (60 ml). Finally, the polymerization was initiated by the addition of HDI (0.81 mL, 5 mmol) at 60 °C for 3 h with continuous magnetic stirring. The polymerized product was



precipitated using cold diethyl ether, followed by filtration and drying under vacuum until the solvents were removed completely.

Synthesis of cationic polymer (butyl-(PN2LG)20) The cationic butyl-(PN2LG)20 synthesis was shown in Figure S4A. Synthesis of butyl-[poly(γ-benzyl-L-glutamate)]20 (butyl-(PBLG)20). For the preparation of butyl-(PBLG)20, BLG-NCA was first synthesized using previously reported procedure.50 BLGNCA (3 g, 11 mmol) was firstly dissolved in CHCl3 (30 mL). Thereafter, butylamine (57 µL, 0.6 mmol) was added into the flask and stirred for 72 h under room temperature. The butyl-(PBLG)20 was precipitated in excessive cold diethyl ether, filtered, and vacuum dried. Synthesis of butyl-(PN2LG)20. For the preparation of butyl-(PN2LG)20, DMF (20 mL) was added to butyl-(PBLG)20 (1.94 g, 0.44 mmol) under nitrogen atmosphere at 55 °C. Thereafter, diethylenetriamine (9.4 mL, 87 mmol) was injected and continued the reaction for 72 h. Finally, butyl-(PN2LG)20 was obtained by precipitation method using an excess amount of cold diethyl ether. For purification, the obtained crude polymer was dispersed in water and dialyzed against water to remove the byproduct and diethylenetriamine, and the dialysate was lyophilized to obtain butyl-(PN2LG)20.

Characterization 1

H NMR: The molecular weight and composition of SMA, SSM, SSM-CL, PEG-PSMEU, and

cationic polymer butyl-(PN2LG)20 were determined using 1H NMR spectra (Varian Unity Inova


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500 instrument (500 MHz)). For the 1H NMR measurement, copolymer samples were prepared using CDCl3, DMSO-D6 or D2O. Gel permeation chromatography (GPC): GPC (Agilent 1100 system, Waldbroun, Germany) technique was used to measure average molecular weight (Mn) and polydispersity index (Ɖ) of synthesized copolymers. Fourier-transform infrared spectroscopy (FT-IR): FT-IR spectra were analyzed using FT-IR spectrometry (FTIR-4100 Type A, TGS, Jasco). Particle size and zeta potential of polyplexes: Particle size analysis and zeta potential measurement of polyplexes at different cation:anion ratio were performed using dynamic light scattering (DLS) (Zetasizer, model Nano-ZS90). Based on the mobility value, zeta potential values are calculated using Smouluchowski’s formula.51 SEM. Cross-sectional morphology of hydrogels was observed by SEM (JSM-6390, MA, USA).

Phase transition pattern The pH- and temperature-induced phase transition was recorded using the tube-inversion method.52, 53 For measurement, 20 wt% PEG-PSMEU copolymers were dispersed in PBS and the pH was raised to 11.0, and stirred for 30 min, to completely dissolve the copolymer. Next, 0.5 ml copolymer solution was pipetted out to 5 ml vials, and the pH was tuned (e.g., 6.8, 7.0, 7.2, and 7.4). The vials were then moved to temperature-tunable water bath and gently raised the temperature. The vials were taken out and the flowability was investigated by inverting vials.

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The viscosity of PEG-PSMEU copolymers with response to physiological stimuli, such as pH and temperature, were examined using Bohlin Rotational Rheometer. Briefly, samples were placed to the 100 mm diameter flat-bottom plate, and then the 20 mm diameter upper plate was brought down. The distance between the bottom and upper plate was fixed at 0.25 mm. Then, the viscosity result was obtained with low shear stress (0.4 Pa) under an oscillation frequency of 1 rad s-1.

In vitro cytotoxicity of synthesized copolymers 293T and RAW 264.7 cell lines were cultured in DMEM supplemented with 10% (v/v) FBS and 1% (w/v) penicillin-streptomycin antibiotic at 37 oC in a humidified 5% CO2. When the cells reach ~80% confluence, they were detached using trypsin-EDTA and seeded in a conventional flat-bottomed plates (96-well) with 1.5 x 104 cells/well. The cells were allowed to grow for 24 h with DMEM and washed two times using PBS (pH 7.4). Thereafter, cells were exposed with different concentrations of copolymers or polymers (PEG-PSMEU and cationic butyl-(PN2LG)20 copolymers). After one day growth, yellow MTT solution (20 µL, 5 mg/mL in PBS), known to be reduced by the mitochondria of living cells, was added to the cells and the incubation continued for another 2 h. Finally, the culture medium and unreacted MTT solution were removed by aspiration. The remaining purple formazan crystals were dissolved by the addition of DMSO (200 µL) and the cell viability was measured at 490 nm by microplate reader (Multiskan Go, Winooski, VT, USA).

Preparation of polyplexes and characterization


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Polyplex formulation and electrophoretic mobility shift assay: Butyl-(PN2LG)20/DNA polyplexes (10:1 (wt/wt)) were prepared by gentle stirring in PBS buffer. To demonstrate polyplex formation, the ionic complex of butyl-(PN2LG)20 and DNA was run on 1% agarose gel at 130 V for 45 min, and was observed by UV illumination. Transfection ability of cationic polymer butyl-(PN2LG)20: To examine the transfection ability of prepared cationic polymer, RAW 264.7 cells were seeded into 12-well plates at 15 x 104 cells/well. At 80% confluence, cells were incubated with 450 µL media without FBS and butyl(PN2LG)20/DNA complexes (50 µL, 1 µg DNA) for 4 h. The medium was replenished subsequently with fresh medium containing FBS, and cells were exposed for one more day. The transfection efficiency was then determined using Micro-BCA assay (Pierce, Rockford, IL, USA). Cellular uptake of polyplexes: FITC-labeled butyl-(PN2LG)20 (butyl-(PN2LG)20-FITC) polymer and DNA (10:1, wt/wt) were mixed to prepare polyplex and exposed to RAW 264.7 macrophages to examine the cellular uptake behavior. After the optimal cell growth in confocal dish, the cells were rinsed twice with PBS, 900 µL of DMEM (without FBS) and 100 µL of release solution obtained from polyplex-loaded hydrogel formulation incubated in PBS (pH 7.4) for 4 h (Seventh day release collection). The solution was then removed and 1 mL DMEM (with FBS) was added to dish and incubated for further 24 h. Finally, cell nuclei were stained using DAPI and fixed with 1.5 mL of 10% buffered formaldehyde solution. The cells were imaged under confocal laser scanning microscopy Zeiss LSM 510 (Carl Zeiss MicroImaging GmbH, Oberkochen, Germany).



In vivo imaging Live animal experiments were conducted in accordance with relevant regulation and institutional guidelines of Sungkyunkwan University. Animal experiments performed in this study were approved by the institutional committees of Sungkyunkwan University. Prior to imaging in vivo, 6-7 weeks old BALB/c female mice were fed with controlled diet food to minimize autofluorescence. For in vivo imaging, mice were injected subcutaneously with either polyplex solution or polyplex-loaded hydrogelators (200 µL), and fluorescence images of the mice were obtained at 0, 3, 7, 10 days using an Optix MX3, ART system.

In vivo gelation and biodegradation Sprague-Dawley (SD) rats with 220-250 g weight bought from Korea Research Institute of Bioscience and Biotechnology (KRIBB, Daejeon, Korea) were used to investigate the in vivo in situ gelation and biodegradation of PEG-PSMEU. The rats were acclimatized for one week in the animal facility (12 h dark/bright cycles) with free access to water and food. For in situ gelation, PEG-PSMEU sample solution (300 µL, 20 wt%, pH 8.8) was injected into the subcutaneous layer of SD rats. Rats were sacrificed at different time intervals and gels were withdrawn to investigate the state of gels and surrounding tissues.

Histocompatibility study Histocompatibility of hydrogels was further confirmed by histological study. For histocompatibility study, 20 wt% of PEG-PSMEU copolymer solution was administered to the dorsal region of SD rats to form gel. Rats were sacrificed at predetermined time points, and the 12 ACS Paragon Plus Environment

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hydrogels as well as their surrounding tissues were isolated. The specimens were fixed in 10% buffered formaldehyde solution and embedded with paraffin. The tissues were then sectioned with a thickness of about 4 µm. Finally, hematoxylin & eosin (H&E) was used to stain skin tissues and visualized under a microscope for histocompatibility examination.

Lysozyme release pattern SD rats were used for lysozyme release in vivo. The rats were randomly split into two groups (n=3). One group was subcutaneously injected with lysozyme solution alone (200 µL, 0.25 mg/mL) while other group was injected with lysozyme-loaded hydrogelators (200 µL, 20 wt%, pH 8.2). At designed time points, blood was collected from lateral rat tail vein and centrifuged to obtain plasma. Then the amount of lysozyme in the plasma was analyzed by lysozyme kit (Mercodia Lysozyme ELISA, Mercodia AB, Sweden).

Adhesion and mechanical property tests The adhesion property of PEG-PSMEU hydrogels was tested by applying them to the different surfaces. For this test, fresh organs such as heart, spleen, lung, and kidney were chosen as hydrophilic surface, whereas polyethylene and titanium were chosen as hydrophobic and metal surfaces, respectively. The adhesion test was conducted immediately after the preparation of in vitro or in vivo gels retrieved after subcutaneous administration. The adhesive property of the gels was also examined by sticking the piece of hydrogels on a human arm. Furthermore, adhesive property of the hydrogels was further demonstrated by placing the solid hydrogels between the two pieces of skins collected from SD rats. Then pressed them for one minute to



adhere effectively, and then lifted to ensure the adhesive property. A universal testing machine (UTM model 5565, LIoyd, Fareham, UK) was to test the mechanical properties, such as tensile and compression strength, of hydrogels.

In vivo wound healing To generate wounds, SD rats were anesthetized and the hair on the skin was removed for clear surgical excision. By using surgical knife, full-thickness incisions (1 cm) were created. The rats were split into 3 groups (n = 4) and applied with PEG-PSMEU hydrogel, PEG-PSMEU hydrogel+polyplex, and suture group in which the wounds were sutured by simple surgical stitching (Control group). The wound closure pattern was photographed at day 0, 1, 3, 5 and 7 post-wounding, and the extent of wound closure was examined. At day 7, the rats were sacrificed and the wound area was recovered, stained using H&E to observe the wound healing effect of hydrogels. Furthermore, the wound breaking strength after the healing was also measured using a universal testing machine (UTM model 5565, LIoyd, Fareham, UK) with a load of 250N.

Statistical analysis The one-way ANOVA was used to test statistical significant difference between free lysozyme solution group and lysozyme-loaded hydrogel group. If p < 0.05, the values are considered statistically significant.


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RESULTS AND DISCUSSION Synthesis and characterization of PEG-PSMEU copolymer Bioinspired PEG-PSMEU copolymers with pH- and temperature dual-sensitivity were synthesized using four-step process with controlled structures (Figure S1). Firstly, SMA was prepared by reacting SM with controlled addition acryloyl chloride. Triethylamine base was used to quench the HCl. Next, SSM was synthesized by the thiol-acrylate Michael-type addition of acryl-terminated SMA and α-thioglycerol. The presence of new peak at 3.36 ppm and the disappearance of acrylate peaks in 1H NMR spectrum of SSM illustrated the transformation of acrylate group into diol group (Figure S2A). Then, the ring-opening reaction was carried out to introduce ester functional group into the monomer; such functional groups may hydrolyze the hydrogel network at the physiological condition. 1H NMR spectra of SSM-CL exhibited new methylene characteristic peaks at 1.1 to 1.96 ppm, which suggested the presence of methylene protons corresponding to the CL (Figure S2B). Finally, PEG-PSMEU copolymers were prepared through the polyaddition reaction of SSM-CL and PEG with HDI cross-linker. The hydrophilic and hydrophobic balance of PEG-PSMEU copolymers were elegantly tuned by controlling feed ratio of cross-linker HDI to hydroxyl-terminals of PEG and SSM-CL (PEG:SSM-CL:HDI = 1:3:4, 1:3:4.5, 1:3:5) (Table 1). 1H NMR spectra of representative PEG-PSMEU1 copolymers, as shown in Figure S2C, shows the methylene characteristics proton signals of HDI at 3.31, 1.35 and 1.22 ppm, whereas the characteristic PEG peak appeared at 3.60 ppm. Furthermore, presence of aromatic peaks at 6.74 to 7.94 ppm, and the protons at 3.36 and 2.10 ppm are corresponds to the peaks of α to oxygen and carbonyl, respectively, suggesting the existence of pH-sensitive SM-TMC in the synthesized copolymer. 1H NMR spectra of PEG-PSMEU indicated the formation copolymers. The PEG-PSMEU copolymer formation was further confirmed using 15 ACS Paragon Plus Environment


GPC. Figure 1A shows the GPC trace of synthesized copolymers. As expected, the GPC trace appeared at a low retention time by increasing the molecular weight of the copolymers. Table 1 shows the Mn and Ɖ of PEG-PSMEU copolymers. The structure and characteristic functional groups in the PEG-PSMEU copolymer was further confirmed using FT-IR spectra. Figure 1B shows urethane N-H group stretching vibration at 3,360 cm-1; whereas, urethane -C=O group stretching vibration bands appeared at 1,730 cm-1. Stretching vibrations of methyl, methylene, and methane C-H bond groups in the PEG-PSMEU copolymer were found at 2,940 cm-1. The characteristic peak observed at 1,346 cm-1 was attributed for vibration band of organic sulfur compounds, -S-CH2-. The stretching vibration of ether (C-O-C) groups in PEG was noticed at 1107 cm-1. Thus, 1H NMR and FT-IR spectra collectively confirms successful preparation of multiblock PEG-PSMEU copolymer with controlled molecular weight.

Sol-gel transition of copolymers The sol (flow) to gel (non-flow) transition pattern of copolymers is a major characteristic that determine the applicability of the materials for in vivo implantation applications. The vialinversion technique was utilized to plot the phase transition pattern of PEG-PSMEU copolymers under different pH and temperature conditions. Aqueous solutions of PEG-PSMEU copolymers displayed sol-to-gel transitions by adjusting the pH and temperature (Figure 1C). The gel window of PEG-PSMEU copolymers covers the physiological condition (pH 7.4 and 37 oC); i.e., the free-flowing copolymer sols at room temperature readily transformed to gel at body temperature. At basic conditions, the sulfonamide groups are ionized, owing to the presence of acidic sulfonamide protons. As a consequence, PSMEU blocks become hydrophilic and the 16 ACS Paragon Plus Environment

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negatively charged blocks electrostatically repel each other, which allowed the easily flow of PEG-PSMEU copolymers solution at high pH. Interestingly, the sulfonamide groups were deionized at the physiological condition and the copolymer turned to hydrophobic. Furthermore, at the physiological temperature (37 °C), hydrophobic balance of PSMEU blocks was significantly increased due to the dehydration and form gel through hydrophobic interaction. More importantly, urethane linkages in copolymers induced the formation hydrogen bonding between the hydrogel networks, and enhanced the strength of the gel. Generally, larger molecular weight molecules have higher viscosity and have strong intermolecular force, which inhibits molecular flow. Thus, it is necessary to examine the appropriate molecular weight of copolymer that can allow exhibit easy flow and injection during subcutaneous administration. Three different PEG-PSMEU copolymers, Mn ranging from 2.9 KDa to 8.8 KDa, were examined (Table 1). By increasing the molecular weight, the sol-gel diagrams of these samples shift down along the temperature axis. This indicates that the larger molecular weight copolymers have strong intermolecular interactions, resulting high viscosity and restricts the molecular flow. Furthermore, hydrophobicity of PSMEU blocks also increased with higher Mn, as more hydrogen bonds from the urethane groups exert strong binding. At high temperature, the copolymer network squeezed out the water molecules more easily, and led to aggregation. Therefore, the PEG-PSMEU1 copolymer showed wide coverage in the gel window. The PEG-PSMEU2 gel region was higher than that of PEG-PSMEU3 along the temperature axis. The rapid solubility and easy molecular flow at high pH and stable gel at physiological pH of PEG-PSMEU1 were used as a representative copolymer for further experiments such as rheology, in vitro toxicity test, and in vivo studies such as biodegradation, adhesion, and wound healing. 17 ACS Paragon Plus Environment


Dynamic rheological measurement To examine the effects of pH and temperature on the viscoelastic property of PEG-PSMEU copolymers rheological analysis was performed (Figure 1D). The change in viscoelastic property in response to stimuli may provide useful information for preparing injectable hydrogelator solution. Viscosity of copolymers was significantly low at high pH and room temperature, which indicated that copolymer solution can easily transferred to syringe. Such low viscous copolymer solution could be administered using small gauge hypodermic needles. Furthermore, such sols are used to load proteins or drugs with simple mixing. It is noteworthy that when the temperature rises to physiological conditions, viscosity of copolymers significantly increases. At the physiological condition, the ionized sulfonamide groups are deionized, and subsequently increase the hydrophobicity of the hydrogel networks. The sol-to-gel transition property of hydrogels, which effectively covers the physiological window, could be utilized in the delivery of various therapeutic agents that need in situ-forming gel depot.

Cytocompatibility of PEG-PSMEU copolymer A material used for in situ-forming injectable hydrogel in vivo application should be biocompatible. To examine the cytocompatibility of PEG-PSMEU copolymer, it was exposed to 293T and RAW 264.7 cells, and the presence of viable cells was assessed by MTT assay (Figure 2A). To determine the optimal viability, a range of concentrations of copolymer, 50-1,000 µg/mL, were incubated for one day in standard DMEM. The cytocompatibility results indicated that even at high copolymer concentrations, PEG-PSMEU copolymer showed no noticeable cytotoxicity.


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Regardless of the copolymer concentrations, higher than 85% cells were viable. This result demonstrates that PEG-PSMEU copolymer is biocompatible. Hence, the PEG-PSMEU copolymer can be utilized in implantation based controlled delivery applications.

In vivo in situ gelation and controlled biodegradation In situ gelation of PEG-PSMEU copolymer solutions was investigated on SD rats by injection of copolymer solutions (300 µL, 20 wt%, pH 8.8) to the back of SD rats. Gels were harvested from SD rats at predetermined time points and the degradation rate of the hydrogel was examined. Irrespective of the copolymer formulation, the gelation results indicate that all three PEGPSMEU copolymer solutions exhibit stable gel formation at the subcutaneous tissue, which is identical to that of sol-gel phase diagram in vitro. Nevertheless, gel properties varied between the three samples. In spite of the fact that PEG-PSMEU1 copolymer formed gel, the mechanical property of the gels was not good due to the low viscosity. In contrast, the PEG-PSMEU3 copolymer with large molecular weight has strong intermolecular interactions and found to be hydrophobic, which exhibited high viscosity and difficult for injection. It was fascinating to observe that the PEG-PSMEU2 copolymer solution at a range of concentrations from 15 wt% to 30 wt%, possessed an appropriate viscosity, and flowed easily in 26G hypodermic needle, and formed stable viscoelastic gel in SD rats (Figure 2B). SEM micrographs of the retrieved gels demonstrated a uniform porous structure with smooth surface morphology (Figure S3). Such porous structure allows the encapsulation of therapeutics. The interconnected porous network structure of hydrogels benefits the exchange of nutrients and infiltration of the cells into the gel. Such properties can be exploited in wound healing, because the optimal wound dressings need to



exchange water and nutrients, and provide moist environment to wounds. Owing to the versatility of PEG-PSMEU2 copolymer, it was used in further experiments. To examine the bioresorbable property of hydrogels in vivo, the PEG-PSMEU2 copolymer solutions (20 wt%) were injected subcutaneously into SD rats and were sacrificed at different time intervals (10 min, 7, 14, 21, 42, 49 days). The gels collected for visual inspection and weight analysis. Figure 2C shows that after three weeks, the hydrogel had lost 21% of its weight via erosion degradation. After 6 weeks, more than 95% of gels were reduced (Figure 2D). At 7 week, no signs of the gel were detected. In the time of the experimental period (~7 weeks), no toxicity was observed, which suggested that the PEG-PSMEU copolymer is bioresorbable and could be utilized in biomaterials as implantable device (Figure 2E).

In vivo release of lysozyme The potential application of PEG-PSMEU copolymer as an in vivo implantable depot to control the release of cationic macromolecules was first evaluated by loading cationic lysozyme (Figure 3). The control samples were free lysozyme solutions. In the case of free lysozyme samples, burst release was observed at the initial time. After 72 h, lysozyme concentration was significantly lower and was disappeared completely. Remarkably, the hydrogel formulation showed significant initial burst inhibition ability, compared to free lysozyme. It is noteworthy that lysozyme release in the bloodstream was sustained for over 7 days. The controlled biodegradation of PEG-PSMEU hydrogels remarkably inhibited burst release and sustained lysozyme release. The controlled release behavior of bioinspired pH- and temperature-responsive PEG-PSMEU copolymer demonstrates that it could be a useful formulation for the localized and long-term delivery of cationic macromolecules. 20 ACS Paragon Plus Environment

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Characterization of polyplexes The major role of PEG-PSMEU copolymer is the ability to sustain local delivery of cationic macromolecules. It should be noted that the therapeutic application of various genetic materials, including DNA or RNA, has been severely limited, owing to their liver accumulation and reticuloendothelial clearance, when administered systemically. As a result, sub-therapeutic concentration of genetic materials is found at the target tissues. To attain optimal therapeutic effect, a high concentration of materials needs to be frequently administered, which resulted in off-target toxicity, and limited the application of genetic materials in gene therapy. To surmount these shortcomings, the local implantation of a hydrogel depot that can control the release of encapsulated agents may be advantageous. Based on the results, the PEG-PSMEU copolymer can be an optimal candidate that can control the release of therapeutic agents for various therapeutic applications. Although various nanoparticles based approaches have been proposed for genetic materials delivery, the rapid delivery of the materials at the blood circulation suffers from various limitations, including loss of bioactivity, enzyme degradation and liver accumulation. Therefore, the genetic materials need to be wrapped with appropriate non-toxic materials, and loaded into the formulation. In this study, for the complexation we first prepared a non-toxic cationic polymer, and prepared a polyplex by simple mixing with genetic materials. Since the cationic polymer has a proton buffering capacity, the polyplex can effectively transfuse the genetic materials into cells, and effectively deliver to maximize the therapeutic effect. For polyplex formation, the cationic polymer was first synthesized (Figure S4A). Based on the 1H NMR spectrum, the comparison of integration peaks between benzylic methylene proton at 5.07 ppm with methyl protons of n-butylamine at 0.69 ppm represents 20 units of BLG 21 ACS Paragon Plus Environment


in the butyl-(PBLG)20 polymer (Figure S4B). Then, the amine groups were introduced by reacting butyl-(PN2LG)20 with diethylenetriamine via aminolysis reaction, which brings positive charge to the polymer. Prior to complexation, the compatibility of this cationic polymer and polyplexes was demonstrated by MTT assay (Figure 4A and 4C), with higher than 80% live RAW 264.7 cells for all concentrations of polymer from 10-2,000 µg/mL and of all ratios of polyplexes, leading to this cationic polymer being deemed suitable for use in in vivo experiments. The electrophoresis image in Figure 4B shows that polyplexes were formed at ratios of butyl-(PN2LG)20/DNA = 3, 4, 5, 10 or 20. The butyl-(PN2LG)20/DNA 10 ratio was chosen as the polyplex model for in vivo experiments, because its transfection-ability is relatively high compared to the other ratios (Figure 4D). Besides, the particle size of the butyl-(PN2LG)20/DNA 10 ratio is less than 90 nm, which exhibits easy escape from filtration by liver and spleen after injection and releasing from hydrogel (Figure 4E). Moreover, the zeta potential value of the butyl-(PN2LG)20/DNA 10 ratio shows positive charge that increases the uptake into the cells to deliver DNA effectively (Figure 4F). These results demonstrate that cationic polymer butyl(PN2LG)20 is appropriate for combination with DNA to form polyplexes, as well as to protect naked-DNAs in blood circulation. Such cationic polymer allows the effective complex formation with anionic macromolecules, including DNA, and can effectively transfuse the genetic materials into cells (Figure 4G).

In vivo imaging To evaluate the sustained release of polyplex-loaded hydrogel, it was subcutaneously administered into the back of mouse, and polyplex solution was used as control group. FITClabeled butyl-(PN2LG)20 was synthesized in two steps, as shown in Figure 5A, and was used to 22 ACS Paragon Plus Environment

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observe the distribution of polyplex after injection. Prior to in vivo imaging, polyplexes released from hydrogels were exposed to the RAW 264.7 cells to examine the cellular uptake. As shown in Figure 5B, FITC signal was mainly detected at the cytoplasm of macrophage cells. The strong fluorescence in the cells indicated high transfection efficiency of butyl-(PN2LG)20. In addition, the high positive charge density of polyplex can easily take up by the cells through electrostatic interaction. Confocal z-stack images strongly suggest that the polyplex was mostly transfected to the cytoplasm of cells. The fluorescent intensity of mouse that was injected with polyplex solution, and reveals rapid diffusion, and that the intensity decreased with time (Figure 5C). The fluorescent intensity was at very low level on the third day, and almost clear within 7 days. Interestingly, the polyplex loaded hydrogel group can retain the signals until 10 days, because the anionic hydrogel interacts with cationic polyplex by electrostatic interaction, moreover, from the slow biodegradation of this hydrogel, polyplexes were released gradually. This result demonstrates the suitability of PEG-PSMEU hydrogel for use as a DNA therapeutics sustained release depot.

Tissue adhesion properties The adhesive property of PEG-PSMEU hydrogels was examined by sticking to the wide range of surfaces, including hydrophilic, hydrophobic, and metal substrates. The hydrogels exhibited strong adhesion to all substrates, indicating that the hydrogels mimicked the adhesion of mussels to range of substrates (Figure 6A and Figure S5). Interestingly, the hydrogel effectively adhere on human skin without causing any irritation and can be easily peeled from skin without any allergies or damage to skin. It is noteworthy that hydrogels firmly adhere on skin despite arm swinging, which implied that hydrogels exhibited strong adhesion during movement (Figure 6B); 23 ACS Paragon Plus Environment


such biomaterials can also have application in soft robotics and soft machines. The adhesion strength of hydrogels with different formulation and composition was also quantified by tensile adhesion test. Irrespective of the formulation, increasing the weight of copolymers increased the adhesion strength (Figure 6C). Particularly, at 20 wt% copolymer concentration the adhesive strength of all copolymers was significantly high. The adhesive property of the hydrogels was further applied to reconnect tissues. Such adhesive function was investigated by simple attachment of rat skins. As expected, control group without any adhesive agents failed to connect the skins. Interestingly, hydrogel group with or without polyplexes exhibited bridging effect and effectively joined the skins (Figure 6D). The tensile strength test on Figure 6E shows that the adhesion strength of hydrogels was significantly high, and it was not affected by mixing therapeutic agents. The displacement test revealed that polyplex-loaded hydrogels showed more displacement when compared with free hydrogels, indicating that presence of multiple amine and phosphate groups in the polyplexes played a major role in displacement test (Figure 6F). Finally, the adhesion property of the hydrogels was further verified on various major organs including kidney, heart, spleen, and lung. The result shows that PEG-PSMEU hydrogels strongly adhere onto the surface of organs, which might be due to the presence of urethane esters that generate bridge between the materials and tissues (Figure 6G and Figure 6H).

Hydrogels for cutaneous wound healing The clinical application of PEG-PSMEU-based adhesive hydrogels as surgical sealants or wound dresser to repair in vivo longitudinal cutaneous wounds has been examined (Figure 7A). PEGPSMEU hydrogels with optimized adhesiveness, mechanical strength and bioresorbable with 24 ACS Paragon Plus Environment

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polyplex were employed as experimental group. On the other hand, pristine hydrogels were chosen as positive control; whereas, the animal group treated using by simple surgical suture was chosen as negative control. After implantation of adhesive hydrogels on cutaneous wounds, the wounds were rapidly healed when compared with suture group (Figure 7B). The extent of wound closure was significantly high and completely closed at 7 days post-surgery for adhesive PEGPSMEU hydrogels releasing polyplexes and had full hair coverage. As expected, the untreated negative control group not completely closed and had blood scabs at day 7. Quantitative demonstration of wound closure rate indicated that at first day differences between the groups was not significant (Figure 7C). In contrast, at day 3, the adhesive hydrogel groups facilitated the wound healing than the negative control group. Remarkably, wound treated with polyplex releasing adhesive hydrogel group closed completely; more importantly, wound marks were completely disappeared and the skin was similar to that normal skin. To confirm the effective wound healing, cutaneous tissues that received treatment were harvested and were applied to measure the wound breaking strength (Figure 7D). The wound breaking strength of polyplex releasing adhesive hydrogel group was closure to that of uncut skin. Although the wound breaking strength for adhesive hydrogels was slightly lower when compared with uncut skin, it was significantly more than that suture group. These results highlighted that the porous adhesive hydrogels could act as sealants and develop tissue microenvironment for neo-tissue formation. Furthermore, histological analysis was carried out to examine the internal structure and condition of treated wounds. The adhesive hydrogel treated wounds was effectively covered by regenerated epidermis with healthy epithelial tissues. H&E stained histological images suggests that the regenerated skin tissue was similar to that of host tissue, indicating the effecting the merging of regenerated tissue with normal tissue (Figure 7E). However, the interfacial difference



was clearly visible for control group treated by simple suture. When the cutaneous wounds treated with polyplex releasing hydrogel, the wounds were effectively closed and the newly formed tissues characteristics mimicked the natural skin, which suggested that localized and controlled delivery of polyplexes from adhesive hydrogels accelerated re-epithelization. In addition, invasion of inflammatory cells can be found in the control group, whereas very less number of inflammatory cells can be found at the hydrogel group, indicating that hydrogels did not provoke inflammatory response. The collagen deposition on the treated wounds was examined by Masson’s trichrome staining (Figure 7F). Hydrogel group showed that huge collagen fiber deposition and alignment in the wounded area. On the other hand, control group loosely aligned with collagen fibers. These results suggested that porous adhesive hydrogels effectively homes cells and facilitated cell migration, resulting tissue regeneration.

CONCLUSIONS In summary, we developed bioinspired in situ-forming pH- and temperature-sensitive injectable adhesive hydrogels with superior adhesive, mechanical, and bioresorbable properties, which allowed the utilization of adhesive hydrogels for multiple biomedical applications. The freeflowing polymer sols at low temperature transformed to gels at physiological conditions. The PEG-PSMEU copolymer hydrogelator could form gel at the back of SD rats, and was found to be bioresorbable after two months. During the implantation period, PEG-PSMEU copolymer showed no sign of inflammation around the injection site. As a consequence, this smart hydrogel shows promising potential for the sustained delivery of therapeutic genetic materials, including DNA. PEG-PSMEU copolymer based physically cross-linked adhesive hydrogels showed excellent adhesion to various hydrophilic, hydrophobic and metal substrates, which makes 26 ACS Paragon Plus Environment

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adhesive hydrogels universal biomaterials that can be used for range of biomedical applications. The in vivo application of polyplex-loaded PEG-PSMEU copolymer adhesive hydrogels effectively sealed the cutaneous wounds, absorb wound exudates, and promote the tissue regeneration in the wounded area. Overall, the bioresorbable pH- and temperature-sensitive PEG-PSMEU copolymer has potential application in various biomedical applications. By simple mixing of therapeutic agents with PEG-PSMEU copolymer hydrogelator, the hydrogel formulation can be applied to other biomedical applications, including controlled delivery, tissue adhesive and engineering.

ASSOCIATED CONTENT Supporting Information The supporting Information is available free of charge on the ACS Publication website. Synthetic route, 1H NMR, SEM, and adhesion strength.

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). This research was also supported by the National Research Foundation of Korea (NRF) funded by The Ministry of Science, ICT & Future Planning (NRF-2017R1D1A1B03028061).



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Table 1. Physicochemical characteristics of PEG-PSMEU copolymers.

Name

Mn (PEG)

PEG:SM-CL:HDI

PEG-PSMEU1

2050

1:3:4

2970

PEG-PSMEU2

2050

1:3:4.5

PEG-PSMEU3

2050

1:3:5

Mole ratio

Mn

Ɖ

(26 G)

In vivo gel properties

1.23

Easy

Weak

5876

1.13

Easy

Good

8798

1.66

Hard

Good


Injectability

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Biomacromolecules

Scheme 1. Schematic concept of sol-to-gel phase transition of PEG-PSMEU copolymers and their versatile biomedical application. (i) Subcutaneous administration of polyplex-loaded PEGPSMEU copolymers sols formed a viscoelastic gel and released the polyplex via diffusion controlled mechanism. (ii) The PEG-PSMEU copolymer hydrogel effectively absorbs onto the surface of skin similar to that mussel inspired adhesives. (iii) The tissue adhesive property of the hydrogels utilized in wound healing to bring the wounds closure and restore the tissue structure and function.



Figure 1. (A, B) GPC trace and FT-IR spectra of PEG-PSMEU copolymers. (C) Phase diagram of PEG-PSMEU copolymers with different molecular weight. The phase transition behavior was measured using the tube-inverting method. Inset photographical images show the sol-to-gel phase transition properties of PEG-PSMEU2 copolymers at different pH and temperature. (D) Viscosity of PEG-PSMEU2 copolymer hydrogel as a function of pH and temperature. The heating rate during the measurement was maintained at 0.5 °C/min.


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Biomacromolecules

Figure 2. (A) In vitro viability of 293T and RAW 264.7 cell lines co-cultured with PEGPSMEU2 copolymer at the indicated concentrations. The error bars represent standard deviations (n = 6). (B) In vivo in situ gelation of PEG-PSMEU2 copolymer in the back of SD rats. The photograph was taken 10 min after subcutaneous administration of the 20 wt% copolymer. (C) Optical images of in vivo biodegradation of hydrogels at the indicated time point after subcutaneous injection aqueous solution of PEG-PSMEU2 copolymers (20 wt%). (D) Extent of degradation quantification of (C) calculated using moss loss method. (E) H&E staining of tissues surrounding the hydrogel implants after subcutaneous injection of the copolymer solution into SD rats.



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* Figure 3. Cationic model protein, lysozyme, concentration in the blood of SD rats after the subcutaneous administration of 200 mL of free lysozyme solution and lysozyme-encapsulated PEG-PSMEU hydrogel formulation (n=5). * indicates statistical difference (p < 0.05) between free lysozyme solution and lysozyme-loaded hydrogel formulation.


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Biomacromolecules

Figure 4. (A) In vitro toxicity of RAW 264.7 cells after exposing with various concentrations of butyl-(PN2LG)20. The error bars represent standard deviations (n=8). (B) Agarose gel electrophoretic band mobility shift assays in of butyl-(PN2LG)20/DNA polyplexes at various 37 ACS Paragon Plus Environment


cation:anion ratios (w/w, butyl-(PN2LG)20/DNA). (C) In vitro toxicity of RAW 264.7 cells after incubation with different concentrations of polyplexes. The error bars represent standard deviations (n=8). (D) Transfection efficiency of polyplexes in RAW 264.7 cells at different cation:anion ratios (w/w, butyl-(PN2LG)20/DNA). (E) Particle size of polyplexes at different cation:anion ratios (w/w, butyl-(PN2LG)20/DNA). (F) Zeta potential of polyplexes at different cation:anion ratios (w/w, butyl-(PN2LG)20/DNA). (G) Schematic illustration of internalization of polyplexes and their subsequent release of DNA at the intracellular level.


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Biomacromolecules

Figure 5. (A) Synthesis route for the preparation of FITC-labeled butyl-(PN2LG)20 polymers. (B) In vitro cell uptake of FITC-labeled polyplexes released form PEG-PSMEU hydrogel. Sample was incubated in RAW 264.7 cell lines in a serum-free culture medium for 1 h. (C) Biodistribution of free polyplex solution and polyplex-loaded PEG-PSMEU hydrogel formulation at various time intervals after subcutaneous administration.



Figure 6. Adhesion behavior of PEG-PSMEU hydrogels. (A) In vitro or in vivo hydrogels prepared using PEG-PSMEU copolymers exhibited excellent adhesion to glass, steel and plastic substrates. (B) The hydrogel also showed excellent adhesion to skin and can be completely peeled off from the skin without causing any harm to skin. (C) The adhesion strength of hydrogels to glass, polyethylene and rat skin under different copolymer composition. (D) Adhesive property of the hydrogels that acts as efficient adhesive to join skin together. (E, F) Quantitative adhesive strength and normalized force displacement curve of hydrogels on (D) was measured using a universal testing machine. (G) Adhesion of hydrogels on biological tissues such as heart, spleen, lung and kidney. (H) Graphic illustration of adhesion chemistry of PEGPSMEU hydrogels on tissue surface.


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Biomacromolecules

Figure 7. Bioinspired PEG-PSMEU copolymer hydrogel as wound dressing materials. (a) Schematic illustration of wound healing and tissue regeneration using bioinspired PEG-PSMEU copolymer-based adhesive hydrogels. (B) Representative wound photographs after 0, 1, 3, 5, and 7 days of healing. (C) Quantification of wound closure kinetics of untreated, PEG-PSMEU, and PEG-PSMEU+Polyplex groups, which was expressed as percentage in comparison with initial wound. (D) Wound breaking strength of healed wounds. (E) H&E staining of wound tissues harvested on day 7. The interface between regenerated tissue (NT) and host tissue (HT) was separated for visualization. (F) Masson’s trichrome staining of wound tissues harvested on day 7. Green arrow represents skin appendages.



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Bioinspired pH- and Temperature-Responsive ... - ACS Publications

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