Antibacterial and Antifouling Hybrid IonicâCovalent Hydrogels with

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Biological and Medical Applications of Materials and Interfaces

Antibacterial and Antifouling Hybrid Ionic-Covalent Hydrogels with Tunable Mechanical Properties Jing Zhang, Biao Shen, Lingdong Chen, Liqun Chen, Jiaying Mo, and Jie Feng ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.9b08870 • Publication Date (Web): 13 Aug 2019 Downloaded from pubs.acs.org on August 20, 2019

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ACS Applied Materials & Interfaces

Antibacterial and Antifouling Hybrid Ionic-Covalent Hydrogels with Tunable Mechanical Properties Jing Zhang*, Biao Shen, Lingdong Chen, Liqun Chen, Jiaying Mo, Jie Feng* College of Materials Science and Engineering, Zhejiang University of Technology, Hangzhou, Zhejiang, 310014, P. R. China *Corresponding authors at: College of Materials Science & Engineering, Zhejiang University of Technology, Hangzhou, Zhejiang 310014, China. E-mail addresses: [email protected] (J. Zhang), [email protected] (J. Feng).

ABSTRACT:

Due

to

their

self-recovery

ability

and

fatigue

resistance,

double-network (DN) hydrogels with hybrid ionically-covalently crosslinking have received wide attention. In this work, by a simple “one-pot” method, a novel kind of hybrid

ionic-covalent

chitosan/polysulfobetaine

(CS/PSBMA)

double-network

hydrogels was prepared. The hydrogels showed high tensile strength (2.0 MPa), strong elastic modulus (0.5 MPa), fast self-recovery ability as well as excellent fatigue resistance and high mechanical strength and toughness retention rate after soaking in water for 24 hours. Additionally, the mechanical properties of the DN gels were enhanced after stretch and relaxation because of the rearrangement of chitosan network. More excitingly, due to the antifouling feature of PSBMA and the inherent anti-bacterial property of chitosan, the hybrid DN hydrogels demonstrated a “repel and kill” effect on microorganisms. The CS/PSBMA DN hydrogels may find potential applications in biomedical fields, such as artificial connective tissues, implantable devices and wound dressing. KEYWORDS:

double

network

hydrogels;

high

self-recoverable; antibacterial; antifouling

1

ACS Paragon Plus Environment

mechanical

properties;

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

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INTRODUCTION Polymer hydrogels have been widely used in wastewater treatment, 1 agriculture and food chemistry,

2-4

tissue scaffolds,

5, 6

drug delivery systems,

7, 8

biosensors

9, 10

and

other fields due to their unique porous network structure. However, most of polymer hydrogels are lack of outstanding mechanical strength, toughness and limited recover ability,

11

which greatly limits their applications. Thus, a lot of gel scientists have

dedicated themselves to solve the problems. At present, many new tough hydrogels have been developed, such as topological hydrogels,

12

double network (DN)

hydrogels, 13 and nanocomposite hydrogels. 14, 15 Among these tough hydrogels, DN hydrogels have emerged as robust candidates for artificial connective tissues, which serve a predominated biomechanical role in the body,

13

and biosensor

9, 10

as well as wound dressing.

16

Generally, the DN gels

consist of two networks with different physical properties, wherein the first network is rigid and easily broken to dissipate energy, thereby improving strength, whereas the second network is flexible to improve the tensile properties of the hydrogel, and achieve a balance between stiffness and toughness. hydrogels have been reported.

13, 18, 19

13, 17

To date, plenty of DN

But most of them are chemically crosslinked,

and the permanent rupture of the covalent network leads to sharp decline of mechanical strength, 20 which cannot meet the needs of application requirements. For example, to replace natural articular cartilage, the hydrogels must be anti-fatigue under a high loading and millions of cycles. 13 Therefore, the features of self-recovery ability of the network and fatigue resistance are highly important for the tough hydrogels. 21, 22 To address the problem, hybrid DN hydrogels with "sacrificial bonds" available to self-repair have been proposed. 11, 23 The hybrid DN gels have a recoverable energy dissipation mechanism through physical network destruction-reconstruction behavior. 11

For example, Wu et al.

24

developed a hybrid chitosan/polyacrylamide (CS/PAM)

ionic-covalent DN hydrogel with excellent mechanical properties, which exhibits high 2


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toughness, tear strength, and tensile modulus (T=14.0 KJ/m2 ，σ=5.6 MPa, E=1.3 MPa). And Zheng et al.

20

reported a DN hydrogel (Agar/PAAm) consisting of a

hydrogen-bond crosslinked agar network and a covalently crosslinked PAAm network, which also shows good mechanical properties (T=9.4 MJ/m3 ，σ=0.3 MPa). All these hybrid DN gels can reach excellent mechanical properties. Unfortunately, the biocompatibility of the hydrogels was rarely studied in those works, which is extremely important for biomedical application. When used in vivo or other biomedical field, the absorption of proteins and bacteria should be concerned. Thus, hydrogels with excellent mechanical properties, anti-biofouling properties as well as biocompatibility will be greatly favorable. Zwitterionic polymers are uniquely nonfouling due to their super-hydrophilicity. They contain a cationic group and an anionic group in one monomer, thus their moieties exhibit neutrally charged and can bind water molecules strongly through electrostatically induced hydration.

25

In order to endue DN gels with antifouling

property, we intent to synthesize a type of zwitterionic polymer involved DN gels to repel proteins and microorganisms. The typical zwitterionic moieties are mainly composed

of

three

kinds:

carboxybetaine

(CB),

sulfobetaine

(SB)

and

phosphorylcholine (PC). 26 As reported, SB polymers show strong intermolecular and intramolecular interactions among zwitterionic moieties.

25

Therefore, we chose

sulfobetaine methacrylate (SBMA) as a component of the DN gels hoping that the associations between PSBMA chains will increase the strength of the gels besides their anti-fouling feature. Herein, we develop a novel hybrid ionic-covalent chitosan/polysulfobetaine double-network hydrogels. Among them, chitosan is physically crosslinked by multivalent anions and behaves as the sacrificial network while SBMA is chemically crosslinked by N, N’-methylene-bis-acrylamide (MBA). We expect the destruction of chitosan network will dissipate energy, bear stress, and recover network structure upon rearrangement to promote the mechanical property of the DN gels. As a result, the hybrid DN gels have excellent mechanical properties, including high tensile 3



strength (2.0 MPa), strong elastic modulus (0.5 MPa), fast self-recovery ability, and excellent fatigue resistance. More importantly, chitosan and PSBMA are both compatible, and besides nonfouling properties of SBMA, chitosan has bactericidal effect. Thus, CS/PSBMA DN gels are highly biocompatible and show a special “repel and kill” property (shown in Scheme 1B). We believe that this hybrid DN hydrogels may find great potential in applications, such as artificial connective tissues and implantable devices. 5, 6

Scheme 1. (A) Preparation of hybrid ionic-covalent CZ-ions DN gels using the “one-pot” method. (B) Representation of “repel and kill” effect of CZ-ions hydrogel.

EXPERIMENTAL SECTION Materials [2-(Methacryloyloxy)ethyl]dimethyl-(3-sulfopropyl) (SBMA), α-Ketoglutaric acid and N, N’-methylene-bis-acrylamide (MBA) were obtained from aladdin. CS (degree of deacetylation > 90 %, viscosity 45 mPa·s, molecular weight < 10 KDa) was supplied by Jinhu Company, China. Sodium citrate was purchased from Sinopharm Chemical Reagent Co., Ltd. Sodium sulfate was obtained from Shanghai lingfeng Chemical Reagent Co., Ltd. Preparation of Gels 4


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The CS/PSBMA DN hydrogels were manufactured by a “one-pot” method. Firstly, 0.02 mol of SBMA, α-Ketoglutaric acid (1 mol% of SBMA), MBA (0.2 mol% of SBMA), 1.0 g of CS and 10 mL of H2O were mixed and stirred to completely dissolve. The CS/PSBMA composite hydrogels were obtained after irradiation with 365 nm ultraviolet light (6 W) for 8 hours under nitrogen atmosphere at room temperature (RT). Then, the CS/PSBMA composite hydrogels were immersed in salt solutions for 30 minutes to prepare CZ-SO42- and CZ-Cit3- hydrogels. Mechanical Tests The mechanical properties of the hydrogel were tested by using a universal test machine (Instron 5966, USA). The samples in the tensile test were dumbbell-shaped, with a length, width and thickness of 16, 4, and 1 mm, respectively, and the rate was 100 mm min-1. The samples in the compression test were cylindrical, with a diameter and height of 8 mm and the rate was 5 mm min-1. The compression-relaxation cycles and tensile loading-unloading cycles was tested at the rate of 5 and 32 mm min-1. The slope of linear range in the stress-strain curves was defined as the elastic modulus (E) of the gels. For self-recovery ability test, the hydrogel samples were rested for 0, 15, 30, 45, and 60 minutes after the first stretch, and then subjected to the re-stretch test. The area between the loading-unloading curves was calculated to determine the dissipated energy (hysteresis, Uhys) of the gels. Swelling Tests The hydrogel samples with a cylinder shape at diameter of 8 mm and height of 3 mm were immersed in 100 mL of deionized water after the first weighing (W0) and weighed (Ws) at regular intervals. The swelling ratio (SR) is calculated by the following equation: (1) Protein Adsorption Measurement 5



Adsorption of HRP-conjugated anti-IgG was used and quantified to evaluate nonspecific protein adsorption of the hydrogels. The hydrogel samples were cuboid with a length, width and thickness of 5, 5, and 1 mm, respectively, and tissue culture polystyrene (TCPS) was used as an experimental control group. The hydrogel samples were placed in 24-well plate containing 1 mL of 1 μg mL-1 HRP-conjugated anti-IgG and incubated at 37 oC for 1 hour and a half. Then the samples were taken out and divided into two groups. The two groups of the samples were immersed in PBS buffer solution, for 0.5 and 3 hours, respectively, and were taken out and washed for 5 times with PBS buffer solution. After that, the samples were placed in a new 24-well plate and 1 mL of 0.1 M citrate-phosphate buffer (pH=5) containing 1 μg mL-1 of o-phenylenediamine (OPD) and 0.03 % hydrogen peroxide was added to the 24-well plate. The reaction was terminated by addition of 2 mL of 2 mol L-1 H2SO4 after 15 minutes. Relative protein adsorption was measured at 492 nm using a microplate reader (DG5033A). The relative adsorption value of the hydrogel samples was calculated by taking the absorbance of the TCPS sample to be 100 %. In Vitro Antibacterial Tests Rectangular sample (10 mm ×10 mm ×1 mm) was sterilized via soaking with 75 % (v/v) ethanol for 30 minutes and then it was immersed in PBS buffer solution for 30 minutes. The sample was placed in a 12-well plate, 1 mL of E. coli/S. aureus solution was added with the optical density (OD) of 0.1, and then it was incubated in shaking table at 120 rpm for 24 hours at 37 °C. Afterward, it was incubated with dye for 15 minutes in the darkness. Fluorescence photographs were taken by AXIO Observer A1 fluorescent inverted microscope. Cytotoxicity 400 mg hydrogels were placed in a 12-well plate, and 2 mL of Roswell Park Memorial Institute (RPMI) 1640 medium containing 10 % fetal bovine serum (FBS) was added, and the hydrogels were cultured at 37 °C for 24 hours. At the same time, L929 cells were seeded in a 96-well plate with a density of 5×104 per well and 6


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cultured at 37 °C for 24 hours. Then the cell culture medium was replaced with 200 μL of hydrogel percolate with different concentrations, and continued to culture the cells for another 24 hours. Subsequently, the medium was removed and 200 μL of MTT solution (0.5 mg mL-1) was added. After 4 hours of reaction, the medium was replaced by 150 μL of DMSO. The dyeing step was carried out under dark conditions. At last, the plate was put in a microplate reader (DG5033A), gently shaken and the absorbance of the solutions in the wells was measured at 492 nm. Results were defined as the mean percentage of cell viability relative to untreated cells and calculated as follows: (2) Where OD treated is obtained in the presence of the hydrogel percolate, OD control is obtained in the absence of the hydrogel percolate, and OD

black

is obtained in the

absence of the hydrogel percolate and cells. Each OD value was measured for four times on independent parallel samples, and the results were expressed as mean ± standard deviation, referring to our previous work. 27 Cell Attachment Experiments Prior to the experiment, the hydrogel samples were sterilized under UV light for 60 minutes. Hydrogel (5×5×1 mm) and TCPS samples were put in a 24-well plate. 1 mL of L929 cell suspension at concentration of 5×104 per mL in RPMI 1640 medium with 10 % FBS was added. After 24 hours, the medium was discarded and all the samples were washed for 5 times with PBS buffer solution. Then all the samples were placed in a new 12-well plate and incubated with FDA solution at concentration of 5 mg mL-1 for 5 minutes. At last, all the samples were observed by inverted fluorescence microscope (Nikon Ti) after washed with PBS for 5 times. In Vivo Biocompatibility Test The biocompatibility test of the CZ-Cit3- hydrogel was evaluated in vivo in a mice subcutaneous model. And CS/PHEAA DN hydrogels were used as a control group. 7



Six independent C57 mice weighing 20-30 g were used for each kind of hydrogels. All hydrogel samples were sterilized using 75 % (v/v) alcohol first. And then two CZ-Cit3- or CS/PHEAA DN hydrogel samples with a disk shape (diameter 10 mm) were implanted into the subcutaneous pocket of the back of each mouse, respectively. After implantation, the mice were housed, fed individually and observed daily. For each group, three mice were euthanized after one week while the other three were euthanized after four weeks. The samples were excised together with each pocket. The specimens were immersed in 10 % formalin for more than 24 hours and embedded in paraffin. Tissue sections were placed on slide, stained with Hematoxylin and Eosin, and observed under microscope. RESULTS AND DISCUSSION Synthesis and Mechanical Properties of CS/PSBMA DN Gels Scheme 1A shows the general “one-pot” method procedure to prepare hybrid ionic-covalent CS/PSBMA DN gels. First, CS, SBMA, α-Ketoglutaric acid and MBA were added to deionized (DI) water, followed by stirring to dissolve completely. Then, through photo induced polymerization process, CS/PSBMA composite hydrogels were synthesized. Afterward, the CS/PSBMA composite hydrogels were immersed into saturated solutions with sulfate or citrate to form CS ionic networks by bidentate or tridentate coordination between the N-glucosamine units of CS chains and SO42- or Cit3- anions and obtain hybrid ionic-covalent CS/PSBMA DN hydrogels (CZ-SO42--x and CZ-Cit3--x, wherein C refers to chitosan, Z refers to zwitterionic polymer and x refers to immersion minutes). 24 Firstly, we conducted a series of tensile tests on CZ-Cit3--30 to investigate the effect of concentration of SBMA monomer, molar ratio of MBA to SBMA, concentration of CS on mechanical performance of the DN gels, respectively (Figure S1-S3). According to Figure S1-S3, in general, the gels achieve the optimal tensile properties (E of 0.5 MPa, σ of 2.0 MPa) when the concentration of SBMA is 2.0 M, the molar ratio of MBA to SBMA is 0.20 mol% and the concentration of CS is 0.10 8


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g/mL. Unless otherwise stated, the experiments were carried out with CS/PSBMA DN gels at the optimal prepared concentration below.

Figure 1. SEM images of: (A) CS/PSBMA composite hydrogel, (B, E) CZ-SO42-, (C, F) CZ-Cit3- hydrogels and (D) PSBMA hydrogel. (G) Digital images of the CZ-SO42- and CZ-Cit3- hydrogels before and after swelling in water for 24 hours. (H) Tensile curves of the CZ-SO42- and CZ-Cit3- hydrogels before and after 24 hours of swelling. (I) Retention coefficient of σ (tensile strength), E (elastic modulus), εb (fracture strain), and T (toughness) of the DN hydrogels after 24 hours of swelling relative to the as-prepared ones.

Then, we tried to study the influence of salt solution concentration and further ionic crosslinking time on the mechanical properties of CS/PSBMA DN gels. As the salt solution concentration increases, the tensile strength of the hydrogel shows an increasing trend (Figure S4), because the increase of the salt solution concentration elevates the crosslinking degree of chitosan network, thus improving the rigidity of the CS/PSBMA DN gels. Therefore, the saturated salt solutions were used for further study. As shown in Figure S5, both CZ-SO42- and CZ-Cit3- hydrogels exhibit the best tensile properties after immersed in saturated sulfate or citrate solutions for 30 minutes. Therefore, the ionic crosslinking time was set as 30 minutes for the 9



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following experiments and CZ-SO42- and CZ-Cit3- are short for CZ-SO42--30 and CZ-Cit3--30 below. As reported, the physical properties of hydrogels can be greatly influenced by their microstructures.

28

So the formation of CS ionic crosslink network in the DN

gels was investigated by a lot of loading-unloading tensile tests. It is well documented that the hysteresis loops indicates the energy dissipation at the time of deformation. 29 In Figure S6, under the same concentrations and conditions, the hysteresis loops of PSBMA single-network hydrogels and the CS/PSBMA composite hydrogels are both ignorable. Conversely, after further ionic crosslinking, the CZ-SO42- and CZ-Cit3hydrogels show obvious hysteresis loops, manifesting the existence of an effective energy dissipation mechanism which is attributing to double-network structure.

11, 30

Thus, we assume that the CS ionic network is formed in the DN hydrogels. Moreover, we used CS/PSBMA composite hydrogels with an interpenetrating network to study the swelling ratio when there is no CS separate network exists. We can see from Figure S7A, if the chitosan and anions exist in free form and no CS ionic network is formed, the swelling ratio of the hydrogels achieves about 30. But according to Figure S7A, after further ionic crosslinking, the CZ-SO42- and CZ-Cit3hydrogels show negligible swelling, which can further prove the formation of a CS ionic network. Through SEM images (Figure 1A-F), we can see that the microstructures of the CZ-SO42- and CZ-Cit3- hydrogels become much denser compared with that of the composite hydrogels, that the pore sizes decrease dramatically to nanoscale after ionic crosslinking compared with the pore size of micro scale (more than 30 µm) for the CS/PSBMA composite hydrogel. The interconnecting pores existed in the crosslinked network also indicate that CS ionic crosslink networks are formed. Additionally, as shown in Figure S8, after immersed in salt solution for 30 minutes, the water contents of CZ-SO42- and CZ-Cit3- decrease to 56 % and 44 %, respectively, from 62 % for CS/PSBMA composite hydrogels, because the interconnecting pores extrude part of water molecules from the network. 10


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Figure 2. Tensile (A) and compressive (B) curves of the CS/PSBMA composite hydrogel, CZ-SO42- and CZ-Cit3- hydrogels. (C-F) Outstanding performance of the CZ-Cit3- hydrogel (C: stretch, D: twisted stretching, E: knotted stretching, F: compressing).

Tensile and compressive curves are displayed in Figure 2A and B. As expected, the smaller pore structures are more effective to disperse stress for toughing hydrogels, and higher crosslink density significantly enhances the stiffness of hydrogels. As a result, mechanical properties of CS/PSBMA DN hydrogels are much improved by further ionic crosslinking. Specifically, the tensile strength of the CZ-SO42- and CZ-Cit3- hydrogels achieves 1.0 and 2.0 MPa, respectively, which is 25 and 50 times of the CS/PSBMA composite hydrogel (0.04 MPa). At the same time, the CZ-SO42and CZ-Cit3- hydrogels achieve high compressive stresses of 108 and 120 MPa at 99.9 % strain, respectively, compared with 1.0 MPa of CS/PSBMA composite hydrogels. And the tensile elastic modulus increases obviously from 18 KPa for the CS/PSBMA composite hydrogel to 150 KPa and 0.5 MPa for the CZ-SO42- and CZ-Cit3- hydrogels, respectively. Additionally, the superior stiffness and toughness of CZ-Cit3- hydrogel than CZ-SO42- hydrogel indicate that N-glucosamine-Cit3- interaction is stronger than N-glucosamine-SO42-. According to the previous literatures, 31-33 owing to swelling equilibrium, most of hydrogels will absorb plenty of water and expand isotropically, resulting in a sharp 11



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drop in mechanical properties. Here, we first studied the swelling ratio of the CS/PSBMA DN gels. Figure 1G and Figure S7B show that CZ-SO42- and CZ-Cit3hydrogels perform small volume expansion and low equilibrium swelling ratio (1.2 and 1.1). As reported, each pair of cationic group and anionic group in the SBMA monomer can bind about 25 water molecules around them.

34

These water molecules

are kind of polarized and bonded with zwitterionic moieties by ionic interaction rather than hydrogen-bond interaction. As known, ionic bonding is much stronger than hydrogen bonding. This is the reason why PSBMA can form a strong hydration layer and exhibit favorable non-fouling property. After polymerization, the intramolecular and intermolecular space is reduced and part of the bonded water molecules is extruded due to stronger ionic interactions between zwitterionic moieties themselves. On the one hand, the strong interaction between PSBMA chains in PSBMA network limits the entry of free water molecules. On the other, the polarized water molecules already bonded in the network of PSBMA cannot form hydrogen bonds with free water molecules thus cannot help PSBMA hydrogel to swell. Therefore, PSBMA single-network hydrogel always shows a dehydration effect and negligible swelling in water. Besides, chitosan network is highly crosslinked by anions, making it difficult to swell. As shown in Figure S7C, the equilibrium swelling ratios of PSBMA hydrogels with different crosslinking degrees do not exceed 1.3 while that of chitosan hydrogels does not exceed 0.2. Based on the properties of these two networks, the resulted CS/PSBMA DN gels show negligible swelling. Then we tested the mechanical properties of the CZ-SO42- and CZ-Cit3- hydrogels at swelling equilibrium. From Figure S7B, it can be found out that the CZ-SO42- and CZ-Cit3- hydrogels can reach swelling equilibrium after soaking in water for 10 hours. So we tested the mechanical properties of CZ-SO42- and CZ-Cit3- hydrogels immersed for 24 hours in order to ensure that the DN gels have completely reach swelling equilibrium. In Figure 1H and I, the tensile strength of CZ-SO42- and CZ-Cit3hydrogels maintains at 0.8 MPa and 1.6 MPa compared with original 1.0 MPa and 2.0 MPa. After swollen for 24 hours, the retention coefficients of tensile properties are over 70 %, indicating that the N-glucosamine-Cit3- tridentate interaction and 12


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N-glucosamine-SO42- bidentate interaction are stable for a long time in water. Thus, the CZ-SO42- and CZ-Cit3- hydrogels could be competitive candidates for applications in watery environment. In addition, as shown in Figure 1G, after the CZ-SO42- and CZ-Cit3- hydrogels equilibrate in water for 24 hours, the transparencies of the DN gels are significantly reduced. That is because the PSBMA polymer chains are stretched in salt solution while coiled in water, causing the refractive index of gels changed. 35 Since crosslink density of CZ-SO42- hydrogels is lower than that of CZ-Cit3- hydrogels, the polymer chains in network of CZ-SO42- hydrogels can move more freely and coil more than those in CZ-Cit3- hydrogels. Thus the transparency of CZ-SO42- hydrogels decreases more significantly than that of CZ-Cit3- hydrogels. Self-Recovery and Fatigue Resistance of the DN Gels As mentioned above, to be used as artificial connective tissues, the hydrogels need to be self-recoverable and anti-fatigue. Thus, we tried to study the destruction and reconstruction mechanism of the networks in the DN gels. Figure 3A shows that the hysteresis energy per unit volume of the CZ-SO42- and CZ-Cit3- hydrogels increases from 0.003 MJ/m3 and 0.025 MJ/m3 at 50 % strain to 0.534 MJ/m3 and 1.360 MJ/m3 at 250 % strain, respectively. Meanwhile, the dissipation coefficient of the CZ-SO42and CZ-Cit3- hydrogels increases to 81.0 % and 87.0 % of the total work at 250 % strain, respectively. The remarkable differences manifest that the N-glucosamine-ion dentate interaction serves as a "sacrificial bond" during the stretching process and destruction of the CS ionic network are extremely effective in dissipating energy. As the strain increases, the area of the hysteresis loops increases significantly indicating more energy is dissipated by more destruction of CS ionic networks (Figure 3B and C). The mechanism is represented in Figure 3G. Moreover, as demonstrated in Figure 2C-F, CZ-Cit3- hydrogels are flexible to endure knotted stretching and compressing with high-level deformation. After removing the deformation force, the gels can quickly recover to their initial shapes, indicating that the gels exhibit excellent shape-recovery property. 13



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Figure 3. (A) The dissipated energy and coefficient of the CZ-SO42- and CZ-Cit3- hydrogel. Loading-unloading curves of CZ-SO42- (B) and CZ-Cit3- (C) hydrogel under different strains (50 %, 100 %, 150 %, 200 % and 250 %). (D) Time-dependent recovery of dissipated energy and elastic modulus of the CZ-SO42- and CZ-Cit3- hydrogel. Self-recovery behavior of the stretched CZ-SO42- (E) and CZ-Cit3- (F) hydrogel at different time (0 minutes, 15 minutes, 30 minutes, 45 minutes, 60 minutes). (G) Dissociation of the N-glucosamine-ions dentate coordination upon deformation and recovery behavior of CS ionic network during relaxing process.

Afterward, in order to investigate self-recovery ability of the DN gels, we tested the tensile properties of the CZ-SO42- and CZ-Cit3- hydrogels with different relaxing time. The data are demonstrated in Figure 3E and F. Without relaxing time, the tensile property of the CZ-SO42- and CZ-Cit3- hydrogels after one loading-unloading cycle decreases dramatically compared to the original sample because the dissociated ionic bonds between N-glucosamine and SO42- or Cit3- cannot be recovered immediately. With relaxing time increasing, the hysteresis loops area is getting closer to the original one. According to Figure 3D, after the DN hydrogels relaxed for 60 minutes, the dissipated energy of the CZ-SO42- and CZ-Cit3- hydrogels recovered to 87.8 % and 80.6 % of the original values, while and elastic modulus recovered to

81.0 % and

87.0 %. These results strongly suggest that the CS ionic network are gradually restored, suggesting an excellent self-recovery capability of the CZ-SO42- and 14


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CZ-Cit3- hydrogels. Due to their recoverable dissipative ability, the ionic-covalent crosslinked DN hydrogels are expected to be anti-fatigue. We studied the fatigue resistance of the CZ-SO42- and CZ-Cit3- hydrogels by continuous loading-unloading tensile tests. As known, hysteresis energy is dissipated by breakage of the chains in gel network. Figure 4A and D shows that the as-prepared and recovered (after relaxing for 24 hours) CZ-SO42- and CZ-Cit3- hydrogels dissipate a large amount of energy in the first cycle, but sharp drops in hysteresis energy and elastic modulus occur in the second cycle, indicating that the CS ionic network of the DN gels are destroyed and cannot restore in a short time. Whereas according to Figure 4C and F, the other eight cycle curves are overlapped and show similar hysteresis energy, implying that the weak hydrogen bonds in the CZ-SO42- and CZ-Cit3- hydrogels are capable of momentarily healing. Moreover, it is worth noting that the recovered CZ-Cit3- hydrogels possess better mechanical properties than the as-prepared ones. In Figure 4, we can see that after relaxing for 24 hours, the stress of the DN gels increases from 0.9 MPa to 2.0 MPa, and the dissipated energy increases from 1.0 MJ/m3 to 1.5 MJ/m3. Furthermore, the curves of other nine cycles in Figure 4A and D show the same phenomena. Since the interaction between N-glucosamine and ions is dynamic, a better network is reconstructed after the rearrangement of CS chains and ions,

21

thus enhances the

mechanical properties of the DN gels. This result can also be proved by Figure 3E and F. As shown, with relaxing time increasing, due to reconstruction of the chitosan ionic network, the maximum stress of the CZ-SO42- and CZ-Cit3- hydrogels increased from 0.43 MPa and 0.95 MPa to 0.55 MPa and 1.35 MPa at 200 % strain, respectively. Therefore, it can be concluded that recoverable CS ionic network imparts excellent fatigue resistance to the CZ-SO42- and CZ-Cit3- hydrogels.

15



Figure 4. Ten successive loading-unloading cycles of: (A, D) as-prepared and (B, E) recovered (resting for 24 hours at RT) CZ-SO42- (A, B) and CZ-Cit3- (D, E) hydrogel. The corresponding dissipated energy in every cycle of CZ-SO42- (C) and CZ-Cit3- (F) hydrogel.

In order to imitate the repetitive force-bearing situation of the natural connective tissues, the CZ-SO42- and CZ-Cit3- hydrogels were conducted to undergo 50 successive compression-relaxation cycles at compressive strain of 50 %. As shown in Figure 5A and D, the maximum force decreases slightly in the force-time curve. Except for the first cycle, the stress-strain curves for the 2-50th cycles are almost overlapped in Figure 5B and E. And it can be found out in Figure 5C and F that after 50 successive cycles, the maximum stress of the CZ-SO42- and CZ-Cit3- hydrogels remains 0.22 MPa and 0.68 MPa, which is 92.0 % and 80.0 % of first cyclic, 0.24 MPa and 0.85 MPa at 50 % strain, and dissipated energy maintains over 3 kJ/m3 and 10 kJ/m3. All these results prove that the network structure of the CZ-SO42- and CZ-Cit3- hydrogels have not been seriously affected after successive force loading and unloading. This supports the results obtained from Figure 5 and suggests that the CZ-SO42- and CZ-Cit3- hydrogels have excellent elasticity and shape-recovery ability. More importantly, it is worth noting that compressive strength of the CZ-Cit3hydrogels is similar to the artificial cartilage (0.24-0.85 MPa at strain of 40-60 %) in a continuous compression-relaxation cycle. 36 So we believe that the CZ-Cit3- hydrogels have a great potential in application that requires highly mechanical properties, such as artificial cartilage, tendon, muscle, and blood vessel. 16


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Figure 5. (A, D) The load-time curves and (B, E) stress-strain curves of fifty successive compression-relaxation cycles of CZ-SO42- (A, B) and CZ-Cit3- (D, E) hydrogel. The maximal stress and dissipated energy of CZ-SO42- (C) and CZ-Cit3- (F) hydrogel every three cycles in 50 continuous compression-relaxation cycles.

In Vitro Anti-fouling Properties and Antibacterial Test of the DN Gels Anti-fouling properties of the DN gels were investigated by adhesion of unspecific proteins and cells on gels. Firstly, the hydrogel samples were immersed in 1 μg/mL HRP-conjugated anti-IgG contained solution for 90 minutes and then soaked in PBS for a certain time. After washing with PBS for five times, the adsorption of the protein onto the hydrogels was evaluated by ELISA. Compared with tissue culture polystyrene (TCPS) control group, all the prepared hydrogels have excellent resistance to adsorption of non-specific proteins. In detail, as shown in Figure 6, by soaking in PBS for 30 minutes, the PSBMA and CZ-Cit3hydrogels show relative lower protein adsorption (4.75 % and 4.82 %), compared with the CZ-SO42- hydrogels (10.75 %). However, all the three prepared hydrogels exhibit low protein adsorption after 3 hours soaking in fresh PBS, which are 1.62 %, 7.39 %, and 4.16 % for the PSBMA, CZ-SO42- and CZ-Cit3- hydrogels, respectively. As reported, adsorption of protein is mainly due to the entrapping of protein molecules by hydrogels during incubation.

37

Thus, protein molecules will take more

time to diffuse out from highly crosslinked hydrogels containing much more pores with smaller size. Figure 1D-I have demonstrated that the pore size of PSBMA 17



hydrogels is much bigger than the CZ-SO42- and CZ-Cit3- hydrogels. So PSBMA hydrogels have an obvious reduction of protein adsorption after incubated for 3 hours in fresh PBS than incubated for 0.5 hours while the CZ-SO42- and CZ-Cit3- hydrogels only have a slight reduction of protein adsorption during the same process. Meanwhile, as displayed in Figure 6, the CZ-Cit3- hydrogels have a smaller static water contact angle (25.4o) than the CZ-SO42- hydrogels (30.8o), which means that the CZ-Cit3- hydrogels could combine water more easily to form a dense hydration layer than the CZ-SO42- hydrogels. This can support the result that the CZ-Cit3- hydrogels have superior anti-proteins behavior than the CZ-SO42- hydrogels.

Figure 6. The relative adsorption of HRP-conjugated IgG on different hydrogels and TCPS.

Liu et al

38

developed a PSBMA fibrous membrane through electrospinning.

Results show that no cell attachment on PSBMA membranes after culturing with Bovine aortic endothelial cells (BAECs) for 48, 72, or 96 hours, indicating an excellent anti-cell adhesion property of PSBMA. Herein, the attachment of mouse fibroblast (L929 cells) on the PSBMA, CZ-SO42- and CZ-Cit3- hydrogels was tested. As shown in Figure 7A1-A3, cell attachment on the surface of TCPS samples is obvious since the 1st day. At the 7th day, a very high density of cells can be observed on the surface of TCPS. Conversely, no L929 cells adhered to the PSBMA, CZ-SO42and CZ-Cit3- hydrogels after 7 days according to Figure 7B-D. This is consistent with 18


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the above results that surfaces of the PSBMA, CZ-SO42- and CZ-Cit3- hydrogels are resistant to protein adsorption. Since no adhesive protein is supplied from FBS or secreted by the cells on the surface of the hydrogels, cells could not attach to the surfaces. Therefore, we can come to the conclusion that the CZ-SO42- and CZ-Cit3hydrogels have an excellent antifouling property, exhibiting a strong “repel” effect to proteins which provides them a promising prospect in applications, such as implantable devices and wound dressing.

Figure 7. The fluorescence microscopic images of mouse fibroblast attached on (A1-A3) TCPS (blank) and (B1-B3) the PSBMA, (C1-C3) CZ-SO42-, (D1-D3) CZ-Cit3- hydrogels at different time (1 day, 4 days and 7 days).

As known, with their unique structures, bacteria are capable of attaching to the surface of the object and producing a polysaccharide protein complex, leading to an irreversible adhesion. 39 The anti-protein adhesion properties of DN hydrogel prevent this process from happening. In addition to “repel” effect to proteins, because of the synergistic effects of antifouling PSBMA and antibacterial CS, the CS/PSBMA DN hydrogels demonstrate a “repel and kill” effect to bacteria. Consequently, bacteria can hardly adhere to it, wherein the little amount of adhered ones will be killed by chitosan. Herein, we used Staphylococcus epidermidis (Gram-positive bacterium) and Escherichia coli (Gram-negative bacteria) to evaluate antibacterial activity of the DN hydrogels. 19



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Figure 8. Representative fluorescence microscopy images of S. aureus attachment on (a1-a3) PSBMA, (b1-b3) CZ-SO42- and (c1-c3) CZ-Cit3- hydrogels for 24 hours. (green staining represents live bacteria and red staining represents dead bacteria)

As we could see in Figure 8, for S. aureus, the antibacterial activity of both DN gels is comparable to that of pure PSBMA hydrogels. Meanwhile, for E.coli, their antibacterial activities are a little inferior relative to pure PSBMA hydrogel (Figure 9). We think this is related to the surface properties of the gels. Because the two main components in the DN gels are positively charged chitosan and neutrally charged PSBMA, making the Zeta potential of the gels becomes positive, and thus the DN gels are more adhesive to Gram-negative bacteria, E.coli. Specifically, there are ignorable bacteria adhered on the CZ-Cit3- hydrogels while little amount of them adhered on the CZ-SO42- hydrogels. We can see from Figure 6, the static water contact angle of CZ-Cit3- hydrogels (25.4o) is lower than the CZ-SO42- hydrogels (30.8o), which means that the CZ-Cit3- hydrogels is more hydrophilic than the CZ-SO42- hydrogels thus could combine water more easily to form a dense hydration layer to resist adsorption of bacterial. Nevertheless, although there is a few E.coli adhered on the CZ-SO42hydrogels, most of them were killed by chitosan (as shown in Figure 9b3).

40

These

results show that incorporation of chitosan into PSBMA hydrogels does not affects their antifouling properties much and donates antibacterial activity to the gels.

20


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Figure 9. Representative fluorescence microscopy images of E.coli attachment on (a1-a3) PSBMA, (b1-b3) CZ-SO42- and (c1-c3) CZ-Cit3- hydrogels for 24 hours. (green staining represents live bacteria and red staining represents dead bacteria)

Biocompatibility of CS/PSBMA DN Gels For applications in biomedical field, biocompatibility of the CS/PSBMA DN gels should be considered. Therefore, we investigated the in vitro compatibility of the DN gels by evaluating of cell viabilities after culturing cells with the hydrogel percolate. Figure 10 shows that cell viabilities are all over 80.0 % in the concentration range of the hydrogel extraction, indicating an ignorable cytotoxicity of the CZ-SO42- and CZ-Cit3- hydrogels on the cells. Then, in vivo biocompatibility of the CZ-Cit3- hydrogels was investigated by implanting hydrogels subcutaneously in mice. Here we used another nonfouling polymer, ploy(N-Hydroxyethyl acrylamide) (PHEAA), to replace PSBMA to synthesize CS/PHEAA DN gels and set them as control. Tissue surrounding the hydrogels was excised and stained with hematoxylin and eosin (H&E) to assess infection. As shown in Figure 11A and C, at 7th day post-surgery, neutrophils were observed in both the control group and CZ-Cit3- hydrogel-treated group, indicating inflammation reactions occurred in the surrounding tissues of the hydrogels. Differently, inflammation reactions in tissues surrounding the CS/PHEAA DN gels 21



are much more severe than that of CZ-Cit3- DN gels. Additionally, there are much less neutrophils adsorbed on the surface of CZ-Cit3- hydrogel than CS/PHEAA DN hydrogel, which further confirmed the antifouling properties of CZ-Cit3- hydrogels in vivo. More excitingly, at 28th day after surgery, there were ignorable neutrophils in the tissue treated with CZ-Cit3- hydrogel, which means that the inflammation gradually dissipated, while inflammation reactions were still observed for the control group. These results manifest that CZ-Cit3- hydrogels own excellent biocompatibility which is greatly favorable in biomedical applications.

Figure 10. Viability of mouse fibroblast after incubated with the extraction of the PSBMA hydrogel, CZ-SO42- and CZ-Cit3- DN hydrogels at different concentrations for 24 hours.

Figure 11. Micrographs of subcutaneous implant with CS/PHEAA DN hydrogel (A, C) and CZ-Cit3- hydrogel (B, D) stained with H&E at day 7 (A, B) and day 28 (C, D). 22


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CONCLUSION In summary, we prepared a novel kind of ionically and chemically crosslinked CZ-ions (CZ-SO42- and CZ-Cit3-) hydrogels through “one-pot” method. The efficient energy dissipated mechanism of double network structure endows the CZ-SO42- and CZ-Cit3- hydrogels with excellent mechanical properties. Because of the reconstruction of reversible CS ionic networks, the CZ-SO42- and CZ-Cit3- hydrogels show fast self-recovery ability, and excellent anti-fatigue capacity. Moreover, the DN hydrogels possess excellent antifouling and antibacterial properties, which is called “repel and kill” effect. The in vitro and in vivo experiments prove that the DN gels are highly biocompatible. Thus we believe that the CZ-ions hydrogels hold great potential for applications in artificial connective tissues, implantable devices and wound dressing. SUPPORTING INFORMATION Effect of different component contents on mechanical properties of DN hydrogels. Tensile curves of the DN hydrogels versus post-crosslinking times. Tensile curves of the DN hydrogels immersed in different concentrations of salt solution. Water contents percentage of DN hydrogels versus post-crosslinking times. Swelling ratio of DN hydrogels, PSBMA hydrogels, CS SN hydrogels and CS/PSBMA composite hydrogels. The loading-unloading curves of the hydrogel at strain of 200 %. CONFLICTS OF INTEREST There are no conflicts of interest to declare. ACKNOWLEDGMENTS This work was funded by Zhejiang Provincial Natural Science Foundation of China (LY17E030005) and the National Natural Science Foundation of China (Grants 21404091 and 21404089). This work was also supported by China Postdoctoral Science Foundation (2019M650143). And we thank the People’s Hospital of Zhejiang Province for helping us to take micrographs of subcutaneous implant with hydrogel. 23



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Antibacterial and Antifouling Hybrid IonicâCovalent Hydrogels with

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