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Biological and Medical Applications of Materials and Interfaces

Hyaluronic Acid and Polyethylene Glycol Hybrid Hydrogel Encapsulating Nanogel with Hemostasis and Sustainable Antibacterial Property for Wound Healing Jie Zhu, Faxue Li, Xueli Wang, Jianyong Yu, and DeQun Wu ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b18927 • Publication Date (Web): 02 Apr 2018 Downloaded from http://pubs.acs.org on April 2, 2018

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ACS Applied Materials & Interfaces

Hyaluronic Acid and Polyethylene Glycol Hybrid Hydrogel Encapsulating Nanogel with Hemostasis and Sustainable Antibacterial Property for Wound Healing Jie Zhu†, Faxue Li*,†, Xueli Wang‡, Jianyong Yu‡, and Dequn Wu*,† †

Key Laboratory of Textile Science and Technology, Ministry of Education, College of Textiles,

Donghua University, Songjiang District, Shanghai, 201620, China ‡

Modern Textile Institute, Donghua University, Changning District, Shanghai, 200051, China

KEYWORDS: hydrogel, nanogel, hemostasis, antibacterial, wound healing

ABSTRACT Immediate hemorrhage control and anti-infection play important roles in the wound management. Besides, a moist environment is also beneficial for wound healing. Hydrogels are promising materials in urgent hemosis and drug release. However, hydrogels have the disadvantage of rapid release profiles, leading the exposure to high drug concentrations. In this study, we constructed hybrid hydrogels with rapid hemostasis and sustainable antibacterial property combining aminoethyl methacrylate hyaluronic acid (HA-AEMA) and methacrylated

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methoxy polyethylene glycol (mPEG-MA) hybrid hydrogels and chlorhexidine diacetate (CHX) loaded nanogels. The CHX loaded nanogels (CLNs) were prepared by enzyme degradation of CHX loaded lysine based hydrogels. The HA-AEMA and mPEG-MA hybrid hydrogel loaded with CLNs (labelled as Gel@CLN) displayed a three dimensional microporous structure and exhibited excellent swelling, mechanical property and low cytotoxicity. The Gel@CLN hydrogel showed a prolonged release period of CHX over 240 hours and the antibacterial property over 10 days. The hemostasis and wound healing properties were evaluated in vivo using a mouse model. The results showed that hydrogel had the rapid hemostasis capacity and accelerated wound healing. In summary, CLN loaded hydrogels may be excellent candidates as hemostasis and anti-infection materials for the wound dressing application.

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1.

INTRODUCTION Wound healing is an intricate and dynamic process including hemostasis, inflammation,

proliferation, and remodeling.1-3 During the healing process, uncontrollable bleeding and wound infection are the most significant fatality issues that delay healing or even cause death, especially in some critical situations.4-6 On one hand, hemostasis is vital and should be achieved rapidly because bleed outs will occur within a short time.6 On the other hand, infection will result in serious tissue damages as microorganisms can compete with the host immune system and subsequently invade viable tissue, especially Staphylococcus aureus.7-10 Hemostatic agents are often used to control the bleeding including bandages, fibrin glues, liquids, powders, gels and dressings.11 Among the aforementioned materials, dressings can cover the injured site to prevent infection and provide physical support for wound healing. As previous studies proved, modern desirable wound dressings should have the characteristics of creating a moist wound healing environment, absorbing excess exudates and being removed easily as well as improving esthetic appearance of the wound site.12-14 Inspired by the concept of moist wound healing, dressings are designed and mainly classified by hydrocolloids, foams, alginates and hydrogels.12,13 Hydrogels have three dimensional (3D) polymeric networks similar to the natural extracellular matrix (ECM) and are capable of absorbing and retaining volumes of water, rehydrating dead tissues and enhancing autolytic debridement.15-19 Hydrogels also aid in rapid hemostasis by promoting the hemostatic plug formation and forming a physical barrier at the bleeding site.20 They can further be employed as carriers of antimicrobial agents and growth factors.21 Researchers have made lots of efforts so far 3



to realize the goals of hemostasis and anti-infection using hydrogels. For example, Behrens et al. applied the N-(3-aminopropyl) methacrylamide hydrogel particles for hemostasis application.22 Bu et al. designed an in situ hydrogel loaded vancomycin using 4-arm-PEG for anti-infection application.6 In general, one of the common methods for developing antibacterial wound dressings is incorporating antibacterial agents into the dressings.23 However, hydrogels are proved to have the disadvantage of relatively rapid release profiles, leading to a harmful side-effect for patients due to the exposure to the high drug concentrations.14,24 For the past few years, micro- or nano-devices as drug carriers which can control the release profile more precisely have gained increasing attention for various applications.14,25-28 Among those, polymer nanogels have the unique advantages of water-retaining and colloidal stability which can be applied in wound treatment.25,29-32 Besides, polymer nanogels with a controllable cross-linked network structure would allow the materials to encapsulate drugs more effectively, achieving a more sustainable and stable release effect. Hashimoto et al. developed a nanogel-crosslinked porous gel to trap proteins, liposomes, and cells.25 Lee et al. established a controlled insulin delivery system using a double-layered nanogel.33 The studies have shown that nanogels may contribute to the future development of novel controlled drug release. Hyaluronic acid (HA) is a linear anionic polysaccharide consisting of a repeated disaccharide of (1-3) and (1-4)-linked β-D-glucuronic acid and N-acetyl β-D-glucosamine monomer.34,35 HA is one component of the extracellular matrix and exhibits good biocompatibility. Moreover, HA plays a significant role in tissue regeneration and angiogenesis, and can promote dermal regeneration.36-38 Compared to hydrogels made from the natural HA, synthetic polymers display 4


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remarkable advantages such as convenient control on the material composition and hydrogel architecture.39 Methoxy polyethylene glycol (mPEG) based materials are used extensively in scaffolds for tissue engineering and drug release.40-42 Many FDA-approved, PEGylated therapeutic agents have been on the market.43 PEGylation can avoid uptake by the reticuloendothelial system, improve pharmacokinetics through stabilization and prolong circulation time, which are promising for enhancing in vivo stability.43-45 We can further control the hydrogel characteristics by changing the mPEG based polymer characters to fit the unique requirements, such as the molecular weight, and functional group amount.46 Meanwhile, antibiotics loaded lysine based nanogels with biocompatibility investigated in the previous study can be used as a sustained and controlled release system for antibacterial treatment.47,48 Here, we designed a hydrogel dressing with hemostasis and sustainable drug release rate formed by chlorhexidine diacetate (CHX) loaded nanogels (CLNs), aminoethyl methacrylate HA and methyl ether methacrylate mPEG as shown in Scheme 1. In detail, HA and mPEG based precursors and CLNs were first prepared by grafting and enzyme degradation, respectively. The HA-AEMA and mPEG-MA hybrid hydrogel loaded with CLNs (labelled as Gel@CLN) was then formed by photo crosslinking in the presence of a photo-initiator. The morphologies of freeze-dried hydrogels before and after loading CLNs were investigated. Swelling property, mechanical property, release kinetics, cytotoxicity and antibacterial property of hydrogels were characterized. Finally the bleeding control capacity and wound healing were evaluated in vivo and results showed the hydrogels had the capacity of rapid hemostasis and decreased the wound infection effectively. 5



Scheme 1. CLN loaded hybrid hydrogel formed by aminoethyl methacrylate HA, methyl ether methacrylate mPEG and CHX loaded nanogels (CLNs). The hydrogel can be used for controlling of bleeding and wound healing.

2. MATERIALS AND METHODS Materials. 1-[3-(dimethylamino) propyl]-3-ethyl-carbodiimide (EDC), N-hydroxysuccinimide (NHS), triethylamine (TEA), N, N’-methylene bisacrylamide (BIS), methoxy polyethylene glycol (mPEG, Mw of 750, 2000, 4000 g mol-1), methacryloyl chloride (MA) was purchased from Sinopharm Chemical Reagent Co., Ltd. The sodium salt of hyaluronic acid (Mw of 90-110 kDa) were purchased from Sigma-Aldrich Co., Ltd. 2-aminoethyl methacrylate hydrochloride 6


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(AEMA) was obtained from InnoChem Science & Technology Co., Ltd. Synthesis of HA-AEMA and mPEG-MA polymers. As previous study described,49 the sodium salt of HA (20 µmol) was first dissolved in 100 mL distilled deionic (DI) water. Then, EDC (15 mmol), NHS (15 mmol), AEMA (10 mmol) and EDC (10 mmol), NHS (10 mmol), AEMA (8 mmol) for different degrees of substitution (DS) were added to the HA solution and mixed for one day, respectively. The solution was dialyzed for several days (dialysis membrane cutoff Mw of 8000) and lyophilized in vacuo at -50 °C for 72 hours. The resultant product with higher degree of substitution was labelled as HA-AEMA-1 and the other one was labelled as HA-AEMA-2. In order to synthesize mPEG-MA, 10 g mPEG was first dried in vacuum and then dissolved in methylene chloride and the temperature was cooled down to 0 oC. MA (5.0 eq. to the molar of mPEG) and excess TEA (5.0 eq. to the molar of mPEG) were added to the mPEG solution and reaction lasted for 12 h. After the reaction, the solution was precipitated in diethyl ether and then dried in vacuum at room temperature to get the products. In order to purify the mPEG-MA, the products were dialyzed (dialysis membrane cutoff Mw of 300) and lyophilized in vacuo at -50 °C for 72 hours. Yield of HA-AEMA-1, HA-AEMA-2, mPEG-MA-1 (Mw of mPEG = 750 g mol-1), mPEG-MA-2 (Mw of mPEG = 2000 g mol-1) and mPEG-MA-3 (Mw of mPEG = 4000 g mol-1) were 61.34%, 52.97%, 91.39%, 98.16% and 96.71%, respectively. Preparation of CHX Loaded Nanogels (CLNs). The lysine based hydrogel was prepared according to the previous studies.47,48 Firstly, di-p-nitrophenyl adipate (NA, 0.76 g) and Lys-4 (1.00 g) were reacted at 80 oC for 30 min to form 4-Lys-4 hydrogel. Then 0.59 g TEA (4.5 eq. to Lys-4) was added to the solution. To obtain CLNs, 0.10 g CHX was added to the above solution 7



additionally before gelling. The freeze-dried 4-Lys-4 hydrogel was immersed in PBS buffer (pH 7.4, 0.1 M) with the trypsin (0.1 mg mL-1) at 37 oC for several days. After 6 days, the biodegradation solution was dialyzed for several days and lyophilized in vacuum at -50 oC for 72 hours. Formation of Hydrogels. In the process of preparing the pure HA-AEMA hydrogel, 0.50 g HA-AEMA and 20 mg APS were dissolved in 2 mL DI water. The hydrogel could be formed by irradiating by UV for 20 min. For the pure mPEG based hydrogel, 1.00 g mPEG-MA, 10 mg BIS and 20 mg APS were first dissolved in 2 mL DI water, and the solution was irradiated by UV for 20 min. The hybrid hydrogels (labelled as Gel-1, Gel-2 and Gel-3 for different Mw of mPEG) were prepared using 0.25 g HA-AEMA-1 and 0.50 g mPEG-MA in 2 mL DI water by photo-crosslinking. For CHX loaded hydrogels and CLN loaded Gel (Gel@CLN), 5 mg CHX and 10 mg CLNs were additionally preloaded in 2 mL DI water and ultrasonic dispersed before hydrogel formation, respectively. Hydrogel Morphology. The morphological structure of hydrogels and CLN loaded hydrogels was investigated by scanning electron microscopy (SEM, HITACHI, TM3000). Before test, the hydrogels were first swollen in DI water for several days to reach the swelling equilibrium. Then the hydrogels were freeze-dried for 72 hours in vacuum at -50 °C for the interior morphological investigation.50 Transmission electron microscopy (TEM, Hitachi, H-800) was conducted at an accelerating voltage of 10 kV to observe the nanogels. Swelling Ratio Characterization. The swelling capacity of hydrogels were characterized by measuring the gravimetric change at 37 oC over time.49-51 Before the swelling test, hydrogel 8


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samples were freeze-dried at -50 oC for 72 hours and the measuring process was described in previous studies using PBS buffer (pH 7.4, 0.1 M). The swelling ratio (Q) is calculated by Eq 1. Q =

Ws Wd Wd

× 100% (1)

In this equation, Ws is the weight of the swollen hydrogel at time t and Wd is the weight of the dry hydrogel at t = 0. Compression Test. Compressive modulus was measured by a DMA Q800 Dynamic Mechanical Analyzer (TA Instruments) in “controlled force” mode (the maximum force of 0.01 N and the rate of 0.0200 N min-1).49,50 Before test, hydrogel samples were swollen in DI water for several days and then cut into the same size (a diameter of 15 mm and height of 10 mm). The modulus (E) is calculated from the plot the compressional stress versus strain by Eq 2. E=

σ2 σ1 ε2 ε1

(2)

In this equation, σ is the stress and ε is the strain at a value of 5%-25%. In Vitro Release of CHX. The release profile of CHX from CHX loaded hydrogels and CLN loaded hydrogel samples were tested in the Ultraviolet-visible spectrophotometer (Hitachi U-4100). The measuring process was described in previous studies and the amount of CHX released was calculated according to the calibration curve by measuring the absorption at the wavelength of 254 nm.49,50 Cell Culture and Toxicity Test. The L929 Fibroblast cell attachment and proliferation on the hydrogel surface was evaluated by qualitative cell morphology and the cytotoxicity was evaluated by MTS assay. The hydrogels were pre-treated as described in previous studies.49,50 After that, the hydrogels were transferred into a 24-well culture plate, and fibroblast cells were 9



seeded (appropriate 30 000 cells per well) and incubated at 37 oC in 5% CO2 for 24 and 48 h, respectively. The MTS solution was added for 3 h incubation and then transferred into a 96-well culture plate for in vitro cytotoxicity and the detailed procedures were in accordance with the protocol. For in vivo toxicity test, 6 week-old male C57BL/6 mice with the weight of 22-25 g were injected subcutaneously with 100 µL hydrogel precursor solution without APS.52,53 The initial weight and weight of mice after being injected with hydrogel solutions was recorded at day 1, 3 and 7, respectively. The mice were sacrificed after one day, three days and one week, respectively. Major organs such as liver, spleen, kidney, heart and lung were fixed in 4% paraformaldehyde buffer solution for further hematoxylin-eosin (H&E) histological analysis. Antibacterial Activity Evaluation of Hydrogels. Antibacterial activity of HA-AEMA-1, mPEG-MA-1, CHX loaded Gel-1 and Gel-1@CLN hydrogels was performed by agar plate diffusion tests.50 The experiment process was described in previous studies, and Escherichia coli (E. coli) and Staphylococcus aureus (S. aureus) were used as the model strain for the test evaluation.49,50 The hydrogels were removed to another fresh inoculated agar plates daily. The test lasted until no inhibition was observed and the zones of the inhibition (ZOI) are estimated by the difference between the outer diameter of the inhibition and the diameter of hydrogel. Besides, the bacterial suspensions were prepared containing approximately 1 × 105 cfu/mL (Control group), then each sample was added and incubated at 37 oC for 24 h. The OD value of bacterial suspension for each sample was measured at 600 nm. The bacterial suspensions were changed daily and the percentage of bacteria survival (BS) at day 1, 2, 4, 7, 10 was determined by Eq 3: 10


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Percentage of BS =

Experimental group OD

× 100%

Control group OD

(3)

1 mL bacterial solution was spread onto the agar plates and incubated at 37 °C for 24 h to form viable colony units. Blood Clotting Index (BCI) Assay. To evaluate the hemostasis effect of different hydrogels, whole blood clotting test was investigated in vitro using HA-AEMA-1, mPEG-MA-1, Gel-1, CHX loaded Gel-1 and Gel-1@CLN hydrogels and a gauze as control. The operating process was referred to the previous study and the absorbance of each sample was measured using a microplate reader (Allsheng, China) at 540 nm.20 To evaluate the blood coagulation effect, the blood clotting index (BCI) is calculated by Eq 4. BCI =

A B

× 100%

(4)

In this equation, A is the absorbance value of blood using each sample and B is the absorbance value of citrated blood (0.4 mL) in 30 mL DI water at 540 nm. In Vivo Blood Coagulation Study. To evaluate the hemostatic potential of the prepared hydrogels, a hemorrhaging liver mouse model was employed (C57BL/6 mouse, 22-25 g, 6 weeks, male, SLRC Laboratory Animal, Shanghai, China).54 Briefly, a mouse was anesthetized and fixed on a surgical corkboard before test. The liver of the mouse was exposed and the resultant fluids were gently wiped away. A piece of parafilm was placed beneath the liver and bleeding from the liver was induced using a 20 G needle with the corkboard tilted at about 30o. The bleeding site was immediately covered with the samples and a filter paper (replaced every 15 s) was used under the liver to evaluate blood coagulation effect of hydrogels. In Vivo Evaluation of Wound Healing. In this study, the wound healing characteristics of 11



Gel-1, CHX loaded Gel-1 and Gel-1@CLN hydrogels were evaluated in a mouse model. The wound covered with a gauze was used as control. At least 3 male mice (C57BL/6 mouse, 22-25 g, 6 weeks, male) were used for each group in the test. Each mouse was anesthetized, the dorsal area was depilated and sterilized with 75% ethanol, then a full-thickness wound for each mouse (round shape with a diameter of 8 mm) were created using a surgical scalpel.50 The wounds of each group were soaked with S. aureus suspension (50 µL, 1.0 × 107 CFU/mL) for 1 h, and then were tightly covered with hydrogels. The hydrogels were changed every 7 days. The wound healing process and the wound size was recorded. The healing extent was calculated using the wound size reduction (C) as previous study described.50 The wound site and adjacent normal skins were collected and tested for H&E analysis as previous study described.50 Blood was also obtained from each mouse after 2 weeks. The obtained blood was treated and the upper serum was collected for enzyme-linked immunosorbent assay (ELISA, Neobioscience, Shenzhen, China) as previous study described.50 Statistical Analysis. All data were evaluated as mean ± standard deviation based on at least three tests. The results are listed as mean ± standard deviation and * indicates significant difference p < 0.05; ** indicates significant difference p < 0.01. Statistical comparison between groups was performed using the student's t-test. Animal Care and Treatment. The animal tests were done in a controlled environment with a 12:12 L/D cycle at 24 ± 2 oC and 60 ± 5% humidity. The animal care and treatment in this study were obeyed the requests of the institutional authority.

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3.

RESULTS AND DISCUSSION Formation and Morphology of Hydrogels. HA-AEMA and mPEG-MA polymers were first

synthesized as shown in Figure 1a. In order to prepare different HA-AEMA hydrogels for comparison, the degree of substitution of the HA-AEMA polymers was controlled by adjusting the amount of AEMA (Figure 1a). The lysine based hydrogel was degraded with trypsin to obtain nanogels (Figure 1b). As shown in Figure S1, the 4-Lys-4 hydrogel with interconnected structure was fabricated successfully and TEM was used to investigate the morphology of the swollen lysine based nanogels with and without loading CHX. After 6 days, lysine based nanogels could be collected, and the TEM results showed the morphology of the nanogels and CHX loaded nanogels had little difference both of which had the average size of 120 nm (Figure S1d&e).

Figure 1. (a) The preparation of hydrogel precursors. (i) The synthesis process of HA-AEMA and (ii) The synthesis process of mPEG-MA. (b) The preparation of CLNs.

As shown in Figure 2a, the products were verified by 1H NMR. In this study, the DS of 13



HA-AEMA was 42.8% and 36.5% according to the integrated peak areas at a and b in Figure 2a. The reference peaks of mPEG-MA at 3.3 and 3.5-3.7 ppm contributed to -O-CH3, -CH2-CH2- of mPEG moiety, and peaks of the methacrylate unit were clearly visible in the spectra (Figure 2a).55 Besides, based on the 1H NMR results, the integration ratio of CH3 in methoxy and methacryloyl end-groups for mPEG-MA-1, mPEG-MA-2 and mPEG-MA-3 was 1.00 to 0.90, 1.00 to 0.91, and 1.00 to 0.97, respectively (Figure S2), indicating the successful reaction of mPEG with MA. As shown in Figure 2b, the peak at the range of 3200 to 3600 cm-1 and the peak at 1760 cm-1 were assigned to -OH bond stretching vibration and C=O stretching from the HA, respectively.49 Besides, the peaks at around 1650 and 1540 cm-1 assigned to typical amide I and II could be observed and the peaks were stronger in HA-AEMA hydrogel.49 Similarly, the IR band of mPEG-MA was clearly seen in spectral regions around 1707 cm-1 due to C=O stretching mode vibration of ester bond (Figure 2c).55 Hydrogels were then formed via photo-crosslinking method using HA-AEMA and mPEG-MA precusors (Figure 1); in particular, CLNs were preloaded into HA-AEMA and mPEG-MA hybrid precusors to form Gel@CLN hydrogels. Optical images of pure HA-AEMA, mPEG-MA, Gel and Gel-1@CLN hydrogels showed that all hydrogels were transparent so that it would be beneficial for observing the wounds as dressings. SEM images of freeze-dried hydrogels are shown in Figure 2d-i. All hydrogels exhibited well-defined 3D pore structures with inherent interconnectivity which met the design requirement of mimicking native ECM.56,57 The swollen structures of HA-AEMA hydrogels were investigated as shown in Figure 2d and Figure S3a. For HA-AEMA hydrogels, the pore size was measured at about 14.38 ± 4.16 µm and 18.52 ± 4.47 µm for DS of 42.8% and 36.5%, 14


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respectively. The HA-AEMA-1 hydrogel with a higher DS had a smaller pore size compared with the hydrogel with a lower DS caused by a higher crosslinking extent.49 Similarly, the pore size was increased with the increase of Mw of mPEG, which was 17.76 ± 2.20, 26.78 ± 2.44 and 35.68 ± 4.03 µm for mPEG-MA-1, mPEG-MA-2 and mPEG-MA-3, respectively (Figure 2e, Figure S3b&c). That is because the mPEG-MA hydrogel with higher average Mw had a lower level of crosslinking in the case of same maromonomer feed ratio. The pore size of Gel-1, Gel-2 and Gel-3 was measured at about 14.94 ± 4.21, 20.82 ± 5.28 and 32.41 ± 9.08 µm, respectively (Figure 2f, Figure S3d&e). The results demonstrated the hybrid hydrogel combined HA-AEMA-1 with mPEG-MA with higher average Mw would lead to a looser structure compared to that of mPEG-MA with lower average Mw. The microstructure in the Gel-1@CLN hydrogel was also investigated as shown in Figure 2g-i. CLNs were uniformly distributed on the pore wall of the Gel-1 without changing the interconnected 3D structure. Figure 2i showed the cross section of Gel-1@CLN and the image further proved the addition of CLNs uniformly without aggregation. While increasing the feed ratio of CLNs, the nanogels became aggregated on the pore wall of the hydrogel as shown in Figure S3f&i. In the further experiment, compared the two CLN loaded hydrogels, we chose Gel-1@CLN with a low CLN loading amount for investigation due to its uniform structure.

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Figure 2. (a) 1H NMR spectra of HA-AEMA and mPEG-MA. (b) FTIR spectra of HA and HA-AEMA. (c) FTIR spectra of mPEG and mPEG-MA. SEM images of freeze-dried hydrogels and optical images of prepared hydrogels: (d) HA-AEMA-1, (e) mPEG-MA-1, (f) Gel-1 and (g-i) Gel-1@CLN. Swelling Ratios of Hydrogels. Water uptake capacity of hydrogels is one of the most essential properties for wound dressing.58 On one hand, sufficient liquid absorption ability is necessary for bleeding control and exudate adsorption at an injured site.17,20 On the other hand, the adequate swelling property is also crucial for keeping the wound site moist which are helpful for the epithelialization.17,20,59 The swelling data of the hydrogels were investigated as shown in 16


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Figure 3a-c. The results showed HA-AEMA hydrogels displayed a sudden increase in the initial period, in detail, during the first 10 min HA-AEMA-1 and HA-AEMA-2 hydrogel absorbed nearly 855.44% and 1004.04% of their weight of PBS, respectively (Figure 3a). Moreover, the pure HA-AEMA hydrogels of the two DS had almost reached their equilibrium swelling in 1 hour. In the equilibrium, HA-AEMA-1 and HA-AEMA-2 absorbed nearly 979.38% and 1235.27% of their weight of PBS, respectively. As shown in Figure 3b, the PBS uptake data of mPEG-MA-1, mPEG-MA-2 and mPEG-MA-3 hydrogel in 10 min were 351.67%, 1041.88% and 2105.54%, respectively. The swelling ratio of all mPEG-MA hydrogels almost reached their equilibrium in 24 hours, which was 556.63%, 1840.29% and 2657.24%, respectively, which was corresponded with the requirement for water uptake. The swelling ratio of Gel-1, Gel-2 and Gel-3 and Gel-1@CLN was characterized as Figure 3c. The PBS absorption of Gel-1, Gel-2 and Gel-3 was 439.61%, 1040.79% and 1505.57%, respectively, in the first 10 min and finally was 617.65%, 1341.34% and 1844.89%, respectively, in 24 hours. With the incorporating of CLNs, the water uptake capacity of Gel-1@CLN had little change compared with that of Gel-1 (Figure 3c). The results indicated the abundant pores in all hydrogels could contain large amounts of water. The main reason for the difference of swelling ratio between the hybrid hydrogels is caused by the different pore size (Figure 2). Benefited from the porous structure and hydrophilic property of HA and mPEG based components, the prepared hybrid hydrogels have an excellent absorption capacity which could contribute to the clotting by concentrating the plates and red blood cells as a homeostatic dressing.

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Figure 3. The swelling ratio of (a) HA-AEMA hydrogel, (b) mPEG-MA hydrogel, (c) Gel and Gel-1@CLN. (d) Compression modulus of the HA-AEMA，mPEG-MA, Gel and Gel-1@CLN hydrogels. Mechanical Properties of Hydrogels. An ideal wound dressing should present good mechanical properties so it chould maintain their integrity during use.60 The compression moduli of the hydrogels are shown in Figure 3d. The results showed the HA-AEMA-1 had a higher modulus than that of HA-AEMA-2 which suggested that a higher DS would contribute to a higher compression modulus. The mechanical property was related to the structure of hydrogels, i.e., the pore size of HA-AEMA-1 hydrogel with a higher crosslinking density was smaller than 18


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that of HA-AEMA-2 hydrogel, which led a more compact structure. Compared with our previous study,49 the molecular weight of the HA in this study is significantly lower than that of the HA before, leading to the pendent vinyl group denser which cause higher compression moduli of HA-AEMA hydrogels. For the mPEG-MA hydrogels, compared with the mPEG-MA-2 and mPEG-MA-3 hydrogels, mPEG-MA-1 had the highest compressive modulus due to the densest structure. It could be noted that with the incorporating of CLNs there was little change in compression modulus. According to Peak’s study, polymer hydrogels encapsulated nanometer size particles could reinforce the hydrogels,61 however, in this study, the amounts of CLNs might be too small to influence and increase the compression moduli of the hydrogel. Drug Release Kinetics. CHX was preloaded during the formation of hydrogels, the DEE and DLE results were calculated as shown in Table S1. The CHX release rate profile at 37 oC in PBS solutions from CHX loaded hydrogels and Gel-1@CLN was shown in Figure 4a. CHX loaded HA-AEMA, mPEG-MA and Gel all displayed a sudden drug release ratio in the early period, especially in the initial 24 hours, and then exhibited a steady growth stage ones during the next 96 hours. In detail, at the time of 5 hours, the cumulative CHX release percent from the HA-AEMA-1 and HA-AEMA-2 hydrogel was about 14.38% and 19.97%, respectively, and that was 15.16%, 25.11% and 35.64% for the mPEG-MA-1, mPEG-MA-2 and mPEG-MA-3 hydrogel, respectively. For the same category, the drug release rate was mainly depended on the hydrogel structure; the release ratio increased with the increase of pore size and the decrease of crosslinking density, i.e., for HA-AEMA hydrogel, the hydrogels with a higher DS in AEMA showed a lower CHX release rate, and for mPEG-MA, the hydrogel formed from mPEG with 19



lower Mw had a slower release rate. At the time of 240 hours the CHX cumulative release from all CHX loaded hydrogels reached nearly 100%. In comparison, the release of CHX from the Gel-1@CLN hydrogels was slower and more stable during the whole periods in this study. At the time of 5 and 24 hours, the CHX cumulative release percent was 4.04% and 56.70%, respectively. The release mechanism for CHX loaded hydrogels and CLNs loaded hydrogels are both mainly a combination of diffusion and erosion.62 However, the reason for the initial release difference is that pores of micrometer-scale in diameter in the hydrogel scaffolds are large enough to allow drug to release immediately and result in the burst drug release. In other words, the burst drug release from the CHX loaded hydrogels was caused by CHX on the hydrogel surface which could be released as soon as being immersed in the medium. In this study, compared with HA-AEMA, mPEG-MA-1 hydrogel which had similar pore size with HA-AEMA showed a slower release rate. Based on previous studies,63-65 the PEG based chains without cleavable segments had no degradation nearly without enzyme in PBS at 37 oC, while the HA-AEMA hydrogel had a little of degradation,49 which probably caused the small difference of release property. While the release of CHX from the Gel-1@CLN hydrogels was attributed to the CHX diffusion from CLN and the erosion of polymers in the PBS solution subsequently. When the hydrogels are applied as wound dressings, the rapid drug release is usually harmful to the wound and shortens the drug efficacy, the feature of CLNs loaded hydrogel will tremendously retain the release in wound healing, effectively lessen the systemic side effect.4,25

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Figure 4. (a) Controlled release of CHX from the hydrogels. (b) The cell viability of hydrogels for the Fibroblast cells after 24 hours’ culture. (c) and (d) Micrographs of cell culture after 24 and 48 hours (the scale bar is 100 µm). Toxicity and Cell Culture. The cell viability was evaluated by quantitative MTS assay as shown in Figure 4b. In this study, the HA-AEMA, mPEG-MA and Gel all exhibited favorable cytocompatibility and Gel-1@CLN had low cytotoxicity after 24 hours’ culture. Besides, the cell viability of Gel-1@CLN after 48 hours and CHX loaded Gel-1 after 24 and 48 hours’ culture was tested for comparison (Figure S4). There was a significant drop in the cell viability of the Gel-1@CLN below 70% while the cell viability of the CHX loaded Gel-1 was below 40% after 21



48 hours’ culture. The results showed the toxicity of CHX loaded Gel was much higher than that of Gel-1@CLN because the release of CHX from the Gel was faster than that of Gel-1@CLN hydrogels. The results further confirmed the slow release kinetics of CLNs as wound dressing were beneficial for minimizing the risk of cells which was consistent with the concept of wound management.30 Besides, in vivo toxicity of hydrogel solution was further examined after 1, 3 and 7 days of injection, and the saline-treated mouse was used as the control group. The results showed the key organs including liver, spleen, kidney, heart and lung in each group had integrated tissue structure without edema, inflammation and abnormal defects (Figure S5), and the weight of mice had little change (Figure S6), indicating that all hydrogels had no significant in vivo toxicity during 7 days of treatment.52,53 Figure 4c&d showed the images of the L929 Fibroblast cells attached and proliferated on the hydrogel surfaces after 24 and 48 hour. It was found that more cells were attached on the surface of mPEG-MA-1, Gel-1 and Gel-1@CLN hydrogels compared with the HA-AEMA-1 hydrogel after 24 hours’ culture (Figure 4c). Furthermore, the image showed the Fibroblast cells had not spread on the surface of HA-AEMA-1 hydrogel completely. After 48 hours’ culture almost no significant cellular morphology and amount changes were observed on the all hydrogel surfaces (Figure 4d). Antibacterial Activity Evaluation. As the whole wound healing process are usually accompanied with high infection risk which will delay the healing process, modern dressings are designed to be endowed with the anti-infection function. The antibacterial activities of HA-AEMA-1, mPEG-MA-1, Gel-1 and Gel-1@CLN hydrogels were first investigated qualitatively and the CHX loaded Gel-1 hydrogel was used as control by the agar diffusion assay 22


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method. Before test, nanogels without loading CHX and nanogel (without CHX) loaded Gel-1 were proved to be not antibacterial against E. coli and S. aureus, respectively (Figure S7a&b). Figure 5a&b showed images of the samples against E. coli and S. aureus, respectively. It has been noticed that no inhibition zone was observed for HA-AEMA-1, mPEG-MA-1 and Gel-1 group, whereas clear inhibition zones were formed for CHX loaded Gel-1 and Gel-1@CLN group towards E. coli and S. aureus for several days. As shown in Figure 5c&d, it was found the inhibition zone of CHX loaded Gel-1 showed a sudden and significant decrease at day 5 against E. coli and S. aureus caused by a robust and fast drug release which was consistent with the drug cumulative release profile in Figure 4a. In comparison, Gel-1@CLN had a more sustained and stable antibacterial property against E.coli and S. aureus for more than 10 days (Figure S7c&d). The actual wt% of CHX was calculated as 0.676‰ and 0.096‰ for Gel-1 and Gel-1@CLN hydrogels, respectively. Although the CHX content in Gel-1@CLN hydrogel was much lower than that in Gel-1 hydrogel, Gel-1@CLN hydrogel had more durable performance of antibacterial property. In order to make a more detailed comparison of the two drug loading systems, the bacteria suspensions were cultured with CHX loaded Gel-1 and Gel-1@CLN hydrogels and the bacteria survival percent at day 1, 2, 4, 7 and 10 was tested. As shown in Figure 5e, nearly 100 percent of bacteria (S. aureus and E. coli) were almost killed at day 1 and 2 for both CHX loaded Gel-1 and Gel-1@CLN hydrogels, and at day 4 the two above hydrogels had a similar and good antibacterial effect. After a week, the ability of both hydrogels in killing S. aureus and E. coli diminished, especially for the CHX loaded Gel-1 hydrogel. After ten days the CHX loaded Gel-1 hydrogel had lost the antibacterial property while the antibacterial ratio of 23



Gel-1@CLN hydrogel still remained 65.71% and 73.56% against E. coli and S. aureus, respectively, further confirmed the durable release profile of nanogels loaded hydrogels. Figure 5f&g displayed the bacterial colonies cultured with Gel-1@CLN and CHX loaded Gel-1 hydrogels, respectively at day 1 and 7. At day 1 there was few visible bacterial colonies without dilution for both groups against E. coli and S. aureus, respectively. At day 7, the amount of formed bacterial colonies (as pointed by red arrows) of the Gel-1@CLN group was significantly less than that of CHX loaded Gel-1 group. Figure 5h showed the SEM images of bacteria morphology before and after being treated with Gel-1@CLN hydrogel and the results showed E. coli and S. aureus displayed an integrated and regular morphology whereas the cell walls of E. coli and S. aureus appeared to be damaged and disorganized after one day of culture.9,66,67 The bacteria morphologies before and after being treated with Gel-1 and CHX loaded Gel-1 hydrogels were also investigated as control (Figure S7e&f). The surface of bacteria treated with CHX loaded Gel-1 hydrogel was partly damaged while the E. coli and S. aureus bacteria treated with Gel-1 hydrogel had little difference at 24 hours’ treatment. The results suggested the Gel-1@CLN hydrogel has the best and most durable antibacterial effect which was helpful for wound infection and had a low frequency of change as wound dressings due to the long-term antibacterial property.

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Figure 5. Optical images of antibacterial activity of hydrogels against E. coli (a) and S. aureus (b). The inhibition zone of hydrogels against E. coli (c) and S. aureus (d). The plates were incubated overnight at 37 oC. (e) Antibacterial efficiency of the hydrogels against E. coli and S. aureus. Bacterial colonies formed using an agar diffusion assay as pointed by red arrows showed treated with Gel-1@CLN hydrogel (f) and CHX loaded Gel-1 hydrogel (g). (h) SEM images of E. coli and S. aureus before and after being treated with Gel-1@CLN hydrogel, red arrows represent damaged bacteria (Scale bars = 1 µm). Hemostatic evaluation. The in vitro blood coagulation effect of HA-AEMA-1, mPEG-MA-1, Gel-1, CHX loaded Gel-1 and Gel-1@CLN hydrogels was evaluated by whole blood clotting 25



measurement compared to a gauze group (Figure 6a&b). As shown in Figure 6a, the BCI of a gauze group were 90.64% and 79.72% in 5 and 10 min, respectively. In comparison, the BCI of prepared hydrogels was much lower than that of the gauze group. Among hydrogels, the mPEG-MA group had the lowest BCI value in both 5 and 10 minutes. The BCI values of Gel-1, CHX loaded Gel-1 and Gel-1@CLN were slightly lower than that of HA-AEMA hydrogel. As the larger BCI value indicated the lower clotting efficiency,20 in this study the results showed prepared hydrogels significantly increased blood clotting rates. The images of samples after test were exhibited as Figure 6b and all hydrogels had absorbed a large amount of bloods in 5 and 10 min. The clots formed in 5 min were less than that in 10 min for the same group and the clots formed on the surface of mPEG-MA-1, Gel-1 and Gel-1@CLN hydrogels are obviously larger than that on HA-AEMA hydrogel. As the controlling of bleeding is a rapid process thus in this study there was almost no difference in hemostasis performance between Gel-1 and Gel-1@CLN hydrogels. The hemostatic effect of hydrogels on bleeding was further investigated in vivo using the mouse liver model. HA-AEMA-1, mPEG-MA-1, Gel-1@CLN hydrogels were applied in the hemostatic experiment. As shown in Figure 6c, the mouse liver was exposed and induced bleeding, and the hydrogel was placed on the bleeding site. It was found that when hydrogels were placed on the site the bleeding was significantly arrested and the complete hemostasis was visually observed compared to the gauze group as control (Figure 6d) and the images of hydrogels using in vivo test were shown in Figure 6e. As shown in Figure 6d, there was no more bleeding coming out from the liver with 100 s for mPEG-MA-1 and Gel-1@CLN groups, and 120 s for the HA-AEMA-1 group. In comparison, the liver could not stop bleeding even after the 26


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gauze was applied for 150 s. Both in vitro and in vivo studies had indicated the mPEG based hydrogels had a better hemostatic effect than pure HA based hydrogel. The possible reasons for the hemostasis mechanism are as follows: when the hydrogels are applied onto the bleeding site, they first act as a physical barrier to blood loss and display the function of physical barrier immediately. Then the water will be absorbed from the serum benefited from the excellent water uptake capacity of hydrogels. This absorption behavior would be helpful for concentrating the blood coagulation factor, red blood cells and plates hence leading to the quick hemostasis.6,20,68 On the other hand, compared with the pure HA-AEMA hydrogel, mPEG based hydrogels are more adhesive to cells (Figure 4c&d) which might be the reason for the better hemostasis effect of mPEG based hydrogels as when a surface is exposed to blood, plasma proteins, platelet adhesion and activation, and other blood cell responses post the main factors to the coagulation.69-72

27



Figure 6. (a) The blood clotting index (BCI) of hydrogels and control sample within 10 min. (b) Photographs of samples used in vitro hemostasis test after 5 and 10 min. (c) The procedures for a liver model in a mouse model. (d) The evaluation of hemostatic ability of the gauze (control) and hydrogels (e) Photographs of samples after in vivo hemostasis test. Evaluation of Wound Healing. The wound healing efficiency using the prepared hydrogels was investigated in vivo and the animal test was divided into control group, Gel-1, CHX loaded Gel-1 and Gel-1@CLN. The wounds were photographed on day 3, 7 and 14. As shown in Figure 7a, the control group showed infection with a tough and severe bacterial biofilm after 3 days’ treatment and had the least wound size reduction. In comparison, the wounds treated with Gel-1, CHX loaded Gel-1 and Gel-1@CLN hydrogels had no visible bacterial biofilm which indicated less bacterial infections for these groups and had an obvious size reduction after 3 days’ 28


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treatment, especially the CHX loaded Gel-1 and Gel-1@CLN group. As shown in Figure 7a, the CHX loaded Gel-1 and Gel-1@CLN groups showed better debridement and antibacterial effects than that of the other groups and the wounds showed significant healing and the size reduction was nearly reached 100% after 14 days’ treatment while the wounds of the control and Gel-1 groups do not heal well as shown in Figure 7b. Besides, H&E analysis was utilized to assess the newly formed skin at day 7 and 14, respectively, as shown in Figure 7c. The results showed there was nearly no epidermis structure and neutrophils still could be found (as the black arrow showed) after the first week’ treatment. The granulation tissue of wound was gradually thickening with the healing time increasing for all groups. Granulation tissue of the defect was replaced by collagen fibers (as circled by the blue circles) after 14 days’ treatment. An almost complete layer of epidermis was visible and hierarchical structure composed of epidermis and dermis was gradually formed in CHX loaded Gel-1 and Gel-1@CLN groups (as the blue and red arrow showed).50 However, no clear epidermal tissue structure was formed for Gel-1 and control groups, and there were still neutrophils existing in the dermis structure for control group. The impact of these wound dressings on the inflammatory response was analyzed by testing the serum inflammation indexes including tumor necrosis factor-α (TNF-α), interleukin-1 IL-1β, interleukin-6 (IL-6) and monocyte chemoattractant protein-1 (MCP-1) via ELISA method. As shown in Figure 8a-d, a significant decrease of TNF-α, IL-1β, IL-6 and MCP-1 was found compared with the control group after 14 days of treatment, indicating a reduced inflammatory response during wound healing process. This suggests a more effective and sustained anti-inflammation property is achieved by CHX loaded Gel-1 and Gel-1@CLN hydrogels. It is 29



believed that when the Gel-1, CHX loaded Gel-1 and Gel-1@CLN hydrogels are applied onto the wounds, the hydrogels all had a debridement function and could avoid the growth of bacteria rapidly due to the excellent hydrophilicity. Compared with Gel-1 hydrogel, CHX loaded Gel-1 and Gel-1@CLN hydrogels with the existence of CHX have a better therapeutic efficacy and they show little difference in the initial healing period. While considering the different released amounts and time, Gel-1@CLN hydrogel with a sustained drug release profile have a slightly shorter healing time which is consistent with Figure 4a&5.

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Figure 7. (a) Photographs of wounds treated by the control, Gel-1, CHX loaded Gel-1 and Gel-1@CLN hydrogel samples (scale bar of 5 mm). (b) Assessment of the wound size reduction. (c) H&E Histological analysis of the regenerated skin tissue after 7 and 14 days of treatment (the white scale bar of 200 µm and the black one of 100 µm).

Figure 8. Expression level of TNF-α (a), IL-1β (b), IL-6 (c) and MCP-1 (d) after 14 days’ treatment: (A) Control, (B) Gel-1, (C) CHX loaded Gel-1 and (D) Gel-1@CLN hydrogels.

4. CONCLUSIONS In this study, a HA and mPEG based hydrogel incorporated CHX preloaded nanogels with cytocompatibility, sustainable antibacterial property, hemostasis and assisted healing ability was developed. The three dimensional hybrid hydrogel acts as a carrier for nanogels as well as 31



provides mechanical support. The Gel@CLN hydrogel is proved to have a more prolonged and stabilized period of drug release than CHX loaded Gel hydrogel. For this reason, the ability of the Gel@CLN hydrogel in killing bacteria was enhanced to 10 days against E. coli and S. aureus, and the cytotoxicity of Gel@CLN hydrogel was also decreased. The Gel@CLN hydrogel was applied for a bleeding site and full-thickness wound model and the results showed this hydrogel had a rapid hemostasis effect and obvious accelerated healing process. In summary, the hydrogel incorporating drug preloaded nanogels might have potential applications for controlling bleeding and wound repairing.

ASSOCIATED CONTENT Supporting Information Synthesis of 4-Lys-4 hydrogel, morphology and size distribution of lysine based nanogels (Figure S1); 1H NMR spectra of mPEG-MA-1, mPEG-MA-2 and mPEG-MA-3 (Figure S2); morphology of HA-AEMA-2, mPEG-MA-2, mPEG-MA-3, Gel-2, Gel-3 and Gel-1@CLN (Figure S3); calculation of drug entrapment efficiency and drug loading efficiency (Table S1); the cell viability of hydrogels (Figure S4); in vivo toxicity of hydrogel solutions (Figure S5); the weight change of mice after being injected with hydrogel solutions (Figure S6); antibacterial activity of unloaded nanogels, unloaded nanogels in Gel-1, Gel-1, CHX load Gel-1 and Gel-1@CLN hydrogels (Figure S7) (PDF)

AUTHOR INFORMATION 32


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Corresponding Author *E-mail: [email protected], [email protected]. Notes The authors declare no competing financial interest.

ACKNOWLEDGEMENTS The research is supported by the National Key Research and Development Program of China (Project No. 2017YFB0309001); the Fundamental Research Funds for the Central Universities and sponsored by the Shanghai Pujiang Program 14PJ1400300.

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