Self-Healing Conductive Injectable Hydrogels with Antibacterial

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Self-Healing Conductive Injectable Hydrogels with Antibacterial Activity as Cell Delivery Carrier for Cardiac Cell Therapy Ruonan Dong,† Xin Zhao,† Baolin Guo,*,† and Peter X. Ma*,†,‡,§,∥,⊥

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†

Frontier Institute of Science and Technology, and State Key Laboratory for Mechanical Behavior of Materials, Xi’an Jiaotong University, Xi’an 710049, China ‡ Department of Biomedical Engineering, University of Michigan, Ann Arbor, Michigan 48109, United States § Department of Biologic and Materials Sciences, University of Michigan, Ann Arbor, Michigan 48109, United States ∥ Macromolecular Science and Engineering Center, University of Michigan, Ann Arbor, Michigan 48109, United States ⊥ Department of Materials Science and Engineering, University of Michigan, Ann Arbor, Michigan 48109, United States S Supporting Information *

ABSTRACT: Cell therapy is a promising strategy to regenerate cardiac tissue for myocardial infarction. Injectable hydrogels with conductivity and self-healing ability are highly desirable as cell delivery vehicles for cardiac regeneration. Here, we developed self-healable conductive injectable hydrogels based on chitosan-graf t-aniline tetramer (CS-AT) and dibenzaldehyde-terminated poly(ethylene glycol) (PEG-DA) as cell delivery vehicles for myocardial infarction. Self-healed electroactive hydrogels were obtained after mixing CS-AT and PEG-DA solutions at physiological conditions. Rapid selfhealing behavior was investigated by rheometer. Swelling behavior, morphology, mechanical strength, electrochemistry, conductivity, adhesiveness to host tissue and antibacterial property of the injectable hydrogels were fully studied. Conductivity of the hydrogels is ∼10−3 S·cm−1, which is quite close to native cardiac tissue. Proliferation of C2C12 myoblasts in the hydrogel showed its good biocompatibility. After injection, viability of C2C12 cells in the hydrogels showed no significant difference with that before injection. Two different cell types were successfully encapsulated in the hydrogels by self-healing effect. Cell delivery profile of C2C12 myoblasts and H9c2 cardiac cells showed a tunable release rate, and in vivo cell retention in the conductive hydrogels was also studied. Subcutaneous injection and in vivo degradation of the hydrogels demonstrated their injectability and biodegradability. Together, these self-healing conductive biodegradable injectable hydrogels are excellent candidates as cell delivery vehicle for cardiac repair. KEYWORDS: self-healing, injectable hydrogels, cell therapy, myocardial infarction, conducting polymer

1. INTRODUCTION Myocardial infarction (MI) has been the enormous cause of death in recent years worldwide.1−3 There will be a large amount of cell loss when an acute MI happens and it will lead to heart failure and death of the patients.4 Cell therapy has been developed as a promising strategy to protect and regenerate myocardial tissue during or after MI by delivery of appropriate types of cells to the MI area.5 Many types of cells have been used in the cell therapy for MI, among which stem cells6,7 and skeletal myoblasts8−10 are most widely used. The fraction size and fibrosis will be reduced by implantation of a certain amount of bone marrow stem cells.11 Moreover, the adipose derived stem cells have been demonstrated to enhance myocardial function and reduce fibrosis when injected into the MI area.12,13 The conventional method of cell implantation is transplanting certain cells through coronary arteries, coronary veins, or peripheral veins as well as the direct intromyocardial injection. However, those delivery methods showed a poor cell retention © 2016 American Chemical Society

which leads to the failure of cell therapy. Fortunately, some research showed that using the biomaterials as cell delivery vehicles could enhance cell retention and viability, as well as cell localization.14,15 Among the biomaterials used as cell delivery vehicles, injectable hydrogels have been widely investigated as an available cell delivery carrier in recent years. Smith et al.15 reported that the cell retention increased more than 7-fold compared with that in PBS by using a hyaluronan-gelatin hydrogel as a vehicle. Wang et al.16 showed that using the chitosan based hydrogel as an injectable cell delivery vehicle could not only enhance the cell retention but also promote the cardiac differentiation of the brown adipose-derived stem cells. Meanwhile, the heart as a vital organ has special electroReceived: April 25, 2016 Accepted: June 16, 2016 Published: June 16, 2016 17138

DOI: 10.1021/acsami.6b04911 ACS Appl. Mater. Interfaces 2016, 8, 17138−17150

Research Article

ACS Applied Materials & Interfaces

supernatant until the product precipitated. Distilled water was used to wash the product until the pH was neutral. Finally, the product was dried in a vacuum oven for 48 h. 2.3. Synthesis of Dibenzaldehyde-Terminated Poly(ethylene glycol) (PEG-DA). PEG-DA polymers were synthesized using a reported method.36 2.0000 g of PEG2000, 0.3000 g of 4-formylbenzoic acid and 0.6108 g of DMAP were first dissolved in 10 mL of dry THF. Then, 1.0318 g of DCC was added into the solution under nitrogen atmosphere. The system was then put at room temperature with stirring. After 72 h of reaction, the solution was centrifuged and the supernatant was precipitated in diethyl ether. The precipitate was redissolved in THF and centrifuged again. The dissolution and precipitation step was repeated for 3 times in order to obtain the product. The final product was dried in a vacuum oven at room temperature. 2.4. Preparation of CS-AT Hydrogel. CS-AT polymer was dissolved in 1% (v/v) acetic acid aqueous solution to form a 3 wt % solution. The PEG-DA solution was prepared with different concentrations to obtain different molar ratios (R = 0.75, 1, 1.5 and 2) of −NH2 and −CHO in the hydrogel. For example, 0.35 g of CSAT solution was then mixed with 0.15 g of PEG-DA solution, and the mixture was placed at 37 °C. The mixture turned into hydrogel less than 1 min. Hydrogels formed by different CS-AT polymers were named as CS-ATX hydrogel. For example, CS-AT10 hydrogel meant that the hydrogel was made from CS-AT10 polymer with a 1:1 molar ratio of −NH2 to −CHO. 2.5. Characterization. A Bruker Avance 400 MHz NMR instrument was used to record the 1H nuclear magnetic resonance (1HNMR) spectra of CS-AT20 and PEG-DA with DMSO-d6 and CDCl3 as solvents respectively at room temperature. The FT-IR spectra of PEG, PEG-DA, chitosan, CS-AT20 polymer and CS-AT20 hydrogel were obtained using a Nicolet 6700 FT-IR spectrometer (Thermo Scientific Instrument) with a resolution of 4 cm−1. 0.5 wt % of CS-AT20 polymer solution (3.5 mL) and 2 wt % of PEG-DA (1.5 mL) were mixed together, and the solution was recorded by UV−vis spectrometer after doping with 1 M HCl. 0.35 g of 3 wt % of CS-AT20 polymer solution and 0.15 g of PEGDA solution were mixed together, daubed onto an ITO glass, and dried at room temperature. Then the dried gel was used as the working electrode accompanied by a platinum wire as the auxiliary electrode, Ag/AgCl as the reference electrode to conduct the cyclic voltammetry (CV) on an electrochemical workstation (CHI 660D). The swelling ratio of these hydrogels was also measured. Briefly, different groups of hydrogels with similar size and shape were immersed into the Dulbecco phosphate-buffered saline (DPBS) at 37 °C and taken out at a regulate time interval. After the excess water was removed with a filter paper, the hydrogels were weighed. The swelling ratio of the hydrogels was calculated using eq 1, where wt represented the hydrogel’s weight after swelling and wo meant the hydrogels weight before swelling.

physiological behavior and the cardiac cells were in an electroactive condition.17,18 It has been demonstrated that conductive polymers may promote the proliferation and differentiation of the electrical stimuli responsive cells, such as stem cells,19,20 myoblast cells,21−23 nerve cells24,25 and cardiac cells.26−29 Therefore, injectable hydrogels with conductivity as cardiac cell delivery vehicle are highly anticipated for cardiac repair. Although conventional injectable conductive hydrogels showed a great deal of advantages such as the promotion of cell proliferation,30 they could not be longterm used with a stable function due to their nonself-healing property. Besides, the cell delivery vehicles would tolerate a considerable mechanical load after injection due to the contraction behavior of the heart,31,32 making the vehicle be damaged easily. When the damage of the vehicle happened during the service period, the self-healing capacity of the hydrogels could prolong the lifetime of the vehicle.33 Therefore, the development of the injectable conductive hydrogels with self-healing ability would be highly beneficial for cardiac cell therapy, and it is an ongoing challenge. Herein, we developed a series of self-healing conductive injectable hydrogels based on chitosan-graf t-aniline tetramer (CS-AT), and demonstrated their potential as cell delivery vehicle for cardiac cell therapy. The hydrogels were synthesized by mixing the solution of CS-AT and benzaldehyde groups capped poly(ethylene glycol) under physiological condition. The conductivity of those hydrogels (about 2 × 10−3 S cm−1) was quite similar to that of native cardiac tissue. These hydrogels showed fast self-healing capacity, good antibacterial activity, adhesiveness and in vivo degradability. They also showed a good biocompatibility that could support the proliferation of the C2C12 cells and prevent cells from injury when injected into the body. C2C12 cells and H9c2 cells were encapsulated in the hydrogel, and they showed a controlled linear-like release profile. All these results suggest that these injectable self-healing conductive hydrogels are ideal candidates as cell delivery carriers for cardiac repair applications.

2. MATERIALS AND METHODS 2.1. Materials. N-Phenyl-1,4-phenylenediamine, butanedioic anhydride, ammonium persulfate, chitosan (Mw = 100 000−300 000), N(3-(dimethylamino)propyl)-N′-ethylcarbodiimide hydrochloride (EDC·HCl), N-hydroxysuccinimide (NHS), poly(ethylene glycol) with Mn of 2000 (PEG2000), 4-formylbenzoic acid, N,N′-dicyclohexylcarbodiimide (DCC) and 4-(dimethylamino) pyridine (DMAP) were all purchased from Sigma-Aldrich and used as received. Tetrahydrofuran (THF) and dimethylformamide (DMF) were stored in a nitrogen atmosphere and dried before use. The carboxyl-caped aniline tetramer was synthesized according to reference.34 All other organic solvents were analytical grade. 2.2. Synthesis of Chitosan-graf t-Aniline Tetramer. Chitosangraft-aniline tetramer (CS-AT) was synthesized according to our previous report.35 We synthesized CS-AT with different AT contents, feed ratio ranging from 8% to 20%. Take the CS-AT8 (the feed ratio of AT in the system is 8%) as a typical example: 0.5 g of chitosan was first dissolved into 50 mL of 0.05 M HCl to obtain a 1 wt % chitosan solution. Second, a mixture of EDC·HCl (0.0916g), NHS (0.0550g) and aniline tetramer (0.044g) was dissolved in 5 mL of DMF, and stirred at room temperature to activate the carboxyl group of aniline tetramer. After 24 h, the mixture was added dropwise into the 1 wt % chitosan solution, and the reaction was continued for another 24 h. Then, the product precipitated by dropping 3 M NaOH into the solution and was then filtered. The precipitate was redissolved in a 1% (v/v) acetic acid aqueous solution and then centrifuged to remove the insoluble parts. After the centrifugation, 3 M NaOH was added to the

swelling ratio =

(wt − w0) × 100% w0

(1)

The morphology of those hydrogels before and after swelling was observed using a field emission scanning electron microscope (Quanta 250 FEG, FEI). The hydrogels were divided into two groups. One group was freeze-dried directly as prepared and another group was freeze-dried after swelling equilibrium state was reached. All the samples were sputter-coated with gold before observation. 2.6. Conductivity of the Hydrogel. Hydrogels with different AT contents were used to study the conductivity of the hydrogels. Seven g of CS-AT solution and 3 g of PEG-DA solution were vortex oscillated to form a homogeneous mixture and the mixture was put at 37 °C for 2 h to obtain completely cross-linked hydrogel. After gelation, the hydrogel was transferred into a cylinder and the conductivity of the hydrogel was detected using a pocket conductivity meter (HANNA8733). 2.7. Rheological Properties of the Hydrogels. The mechanical properties of hydrogels with different concentrations of PEG-DA were 17139


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were tracked with the CellTracker Green (5 μM). These two types of cells were encapsulated into hydrogels individually and incubated for 12 h. After incubation, hydrogels encapsulating with C2C12 cells and hydrogels with ADMSCs were mixed together by injection. Furthermore, two pieces of the hydrogels with different cells were put together to make it integrate into a construct. After 24 h, the samples were observed under the confocal laser microscope (FV1200, Olympus). 2.12. In Vitro Cell Delivery Property of the Hydrogels and in Vivo Cell Retention Test. To test the cell delivery ability of this hydrogel, a cell delivery test was conducted. C2C12 cells were marked and encapsulated into CS-AT10 hydrogel in a 48-well plate with a density of 1 × 106 cells mL−1, and 250 μL of DMEM with 10% FBS were added into each well. After cultured for 24 h, the DMEM in each well was transferred into a new well respectively, and 250 μL of DMEM with 10% FBS was added into the hydrogel’s well. The cell number in the transferred solution was calculated using the NIH ImageJ software after observing under inverted fluorescence microscopy (IX53, Olympus). The released cell number was analyzed from 1 to 7 days. The in vivo cell retention was measured using the in vivo optical bioluminescence imaging. ADMSCs were incubated with CellTracker Deep Red (5 μM) before encapsulated into the hydrogel or suspended with DPBS. About 1× 106 of cells were encapsulated into 100 μL of hydrogel (or 100 μL of PBS as control), and injected into the mouse subcutaneously immediately. The fluorescence intensity was correlated with the cell number.16,37 The relative fluorescence intensity was calculated by eq 2, in which FIt and FI0 mean the fluorescence intensity at t day (t = 3, 7) and 0 day. The bioluminescence image of the mouse with the cells was taken using the IVIS Spectrum instrument (Xenogen, USA) at 1, 3 and 7 days after injection.

measured using a TA rheometer (DHR-2) instrument with a parallelplate configuration (plate diameter of 20 mm, gap of 1000 μm). A dynamic frequency sweep test with frequency ranging from 1 to 100 rad·s−1 at a shear amplitude of 0.1% was employed to perform the measurement. The hydrogels with different molar ratios −NH2 to −CHO (0.75, 1, 1.5 and 2) and AT contents were made as disks (diameter of 20 mm and height of about 1 mm). 2.8. Self-Healing Behavior of the Hydrogels. The self-healing behavior of the hydrogel was tested with both microscopy and quantitive methods using the CS-AT10 hydrogel as an example. In the microscopy method, the hydrogel was first injected through a syringe into a heart shape mold and piled up at 37 °C in a humidified atmosphere. And then the hydrogel was cut into two pieces. Subsequently, the two pieces of hydrogels were put close to each other at 37 °C in a humidified atmosphere. The results of the selfhealing test were photographed in a certain time. In the quantitive way, the hydrogels were prepared as a disk with a 20 mm in diameter and a 1 mm in height. Using the strain amplitude sweep method (γ changed from 1% to 500%), the value of the critical strain region was detected. And then a new hydrogel with the same size was used to test the self-healing behavior by using the alternate strain sweep test at a fixed angular frequency (1 rad·s−1). Amplitude oscillatory strains were switched from small strain (γ = 1.0%, 100 s for each interval) to large strain (γ = 250%, 60 s for each interval), and 6 cycles were carried out in the test. The self-healing behavior under a cyclic high and low stress was taken out to measure the self-healing ability of the hydrogel under a native cardiac contractile environment. The high stress was set as 400 and 800 Pa (60 s for each interval) respectively and the low stress was 1 Pa (100 s for each interval). 2.9. Cell Encapsulation and Proliferation in the Hydrogel. CS-AT10 polymers were both dissolved in 1% (v/v) acetic acid solution (100 μL of acetic acid in 10 mL of DMEM solution) to obtain a 3 wt % CS-AT10 polymer solution. Meanwhile, the PEG-DA polymer was dissolved in Dulbecco’s modified Eagle’s medium (DMEM) to form a 23.4 wt % solution. Then the CS-AT10 and PEG-DA solution were sterilized. Before encapsulated in the hydrogel, the C2C12 cells were treated with CellTracker Green (5 μM) for 30 min. To encapsulate the cells, first, 150 μL of PEG-DA solution weas added into the 24-well plate. Then, 5 μL of cell suspension was added into the mixture and finally 350 μL of CS-AT10 polymer solution. After gelation for 1 h at 37 °C, 500 μL of DMEM with 10% FBS was added into each well and the plate was incubated at 37 °C in a humidified atmosphere incubator containing 5% CO2. After 24, 48 and 72 h cultivation, the hydrogels were taken out of the wells carefully. Then they were observed under a confocal laser microscope (FV1200, Olympus) and analyzed using NIH ImageJ software. The thickness of scanning was 60 μm for each sample. Adipose-derived mesenchymal stem cells (ADMSCs) proliferation in the CS hydrogel and CS-AT10 hydrogel was also measured. ADMSCs were incubated with CellTracker Green (5 μM) for 30 min before encapsulated into the hydrogels. The cells were encapsulated in the hydrogels as described above. After culturing for 24, 48 and 72 h, the hydrogels were taken out of the well and observed under a confocal laser microscope (FV1200, Olympus). The results were analyzed using NIH ImageJ software and the thickness of scanning was 50 μm for each sample. 2.10. Cell Viability in Hydrogel after Injection. After cells were encapsulated in the hydrogels, the hydrogels were divided into two groups. One of them was incubated directly in the well for another 24 h, and another group was injected using a 22-gauge needle. After injection, the hydrogels were incubated at 37 °C for 2 h to make it selfheal and 500 μL of DMEM were then added into each well. After 24 h of incubation, the hydrogels encapsulating cells and the injected hydrogels were observed under the confocal laser microscope (FV1200, Olympus). The picture of each sample was merged of 60 slides and analyzed using NIH ImageJ software. 2.11. Two Types of Cells Coculture in the Hydrogels. ADMSCs were incubated with CellTracker Red (5 μM) for 30 min before encapsulated into the hydrogels. At the same time, C2C12 cells

relative fluorescence =

FI t FI 0

(2)

2.13. Tissue Adhesive Test of the Hydrogels. To investigate the adhesive ability of the hydrogel to the host tissue, a lap-shear test with fresh porcine skin as adherend was conducted. Briefly, the fresh porcine skin was cut into rectangle pieces with 2 × 1 cm size and then immersed into DPBS before use. 150 μL of mixture containing 105 μL of CS-AT (3 wt %) and 45 μL of PEG-DA (23.4 wt %) was added onto the porcine skin and placed at 37 °C for 1 h to make the gelation complete. Subsequently, adhesion properties were tested using the lapshear test on an Instron Materials Test system (MTS Criterion 43, MTS Criterion) equipped with a 50 N load cell at a rate of 5 mm/min. All the tests employed more than 3 samples. 2.14. Antibacterial Properties of the CS-AT Polymers and Hydrogels. Minimum inhibitory concentrations (MICs) of the CSAT polymers for both Escherichia coli (ATCC 8739) and Staphylococcus aureus (ATCC 29213) were investigated using a reported method.38 Briefly, CS-AT copolymers were dissolved in 1% (v/v) acetic acid solution to obtain 3 wt % copolymer solutions. Subsequently, these 3 wt % copolymer solutions were diluted with MH broth using the 2-fold dilution method in a 96-well microplate and a series of copolymer solutions were obtained (100 μL in each well). 100 μL (104−105 CFU mL−1) of bacterial suspension was added into each well and mixed repeatedly to acquire a homogeneous solution. The gradiently diluted copolymer solutions with no inoculum were chosen as the negative control and MH broth with inoculum served as the positive control. The 96-well plate was incubated at 37 °C for 18 h, and the absorbance of the solutions at 600 nm was read using a microplate reader (Molecular Devices) at the end of time. The minimum concentration of the copolymer that could inhibit the bacterial growth was intended to be the MICs of the copolymers. To evaluate further the antibacterial activity of the hydrogels, surface antibacterial assay of the hydrogels was conducted. Hydrogels were prepared as previously described: 140 μL of CS-AT solution (3 wt %) and 60 μL of PEG-DA solution (23.4 wt %) were mixed homogeneously, and then 200 μL of mixture was added into each well in a 48-well plate. The plate was placed at 37 °C for complete gelation, and 5 μL of bacterial suspension (in PBS, 105 CFU mL−1) was spread 17140


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Figure 1. Scheme of synthesis procedure of the CS-AT polymer (a), PEG-DA polymer (b) and CS-AT hydrogel (c).

Figure 2. FT-IR spectra of CS-AT20 hydrogel, CS-AT20 polymer and PEG-DA polymer (a); UV−vis spectra of the doped CS-AT20 hydrogel and CS-AT20 polymer (b); CV curve of the CS-AT20 hydrogel (c); Conductivity of the hydrogels (d). onto the surface of each hydrogel subsequently. Afterward, the

kill% =

hydrogels were incubated at 37 °C in a humidified atmosphere for 4 h.

× 100%

Any bacterial survivors were resuspended by adding 1 mL of PBS into

2.15. Subcutaneous Injection, In Vivo Degradation of CS-AT Hydrogels and Histopathological Analysis. Female Harlan Sprague−Dawley rats (200−220 g) were anesthetized with 10% chloral hydrate and shaved. The CS-AT solution and the PEG-DA solution were sterilized. The hydrogels with different AT contents (CS-AT8, CS-AT10 and CS-AT20) were prepared and then injected into the shaved rats subcutaneously with the same volume (200 μL)

each well then. One mL of PBS solution with 5 μL bacterial 5

cell count of control − survivor count on experimental hydrogels cell count of control

−1

suspension (in PBS, 10 CFU mL ) served as the control. After incubated at 37 °C for 12−24 h, the colony-forming units (CFU) on the Petri dish were counted. Each test was repeated for 3 times and the kill% was used to measure the antibacterial activity of the hydrogels. 17141


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Figure 3. Swelling ratio of the CS-AT hydrogels with different AT contents in PBS (a). Storage modulus of CS-AT hydrogels with different AT contents and molar ratio between −NH2 and −CHO (b). SEM images of CS-AT hydrogels before and after swelling (c), scale bar: 300 μm. on the dorsum using a 22-gauge needle in a sterile condition. 15 min later, the rats were sacrificed by cervical dislocation and the hydrogels were taken out to photograph. For in vivo degradation test, the hydrogels were injected under the dorsum of the rats as described above and were taken out from the dorsum at 5, 10, 15, 20, and 45 days and photographed. The degradation properties of the hydrogels in vivo were evaluated by measuring the size of them under the dorsum. CS-AT10 hydrogels were injected into the SD rats subcutaneously on the dorsum after sterilized. After 7 days, the tissue near the hydrogel was excised and frozen in O.C.T medium. The frozen section was cut into 20 μm in thickness. The sections were stained with hematoxylin−eosin (H&E) stains to evaluate the inflammatory response of the hydrogels. All experiments were performed in compliance with the relevant laws and institutional guidelines, and had been approved by the Animal Ethics Committee of the University. 2.16. Statistic Analysis. All the experiments were conducted using more than 3 samples, and the data were expressed as mean ± standard. The statistical significance was analyzed using the student t test, and it was considered to be a statistically significant difference when p < 0.05.

synthetic scheme of CS-AT, PEG-DA polymer and the CS-AT hydrogels. The CS-AT hydrogel was prepared by Schiff base reaction between the amino groups in CS-AT and the aldehyde groups in PEG-DA. Because of the high reactivity of amine and aldehyde to form aromatic Schiff base,39,40 these CS-AT hydrogels were obtained at a mild condition, making it suitable for the in vivo application. Besides, with the increase of AT content in the CS-AT polymer, the gelation time increased from ∼20 to ∼60 s. Chemical structures of the resultant hydrogels were verified using FT-IR spectroscopy. A peak at 1720 cm−1 appeared in PEG-DA, which was attributed to the stretching vibration of the CO bond in the aldehyde group (Figure 2a). Peaks at 1580 and 1644 cm−1 of CS-AT20 copolymer owned to the characteristic stretching vibration of CC bond in the aromatic ring and that of the amide groups (−NHCO−) in AT, indicated that AT was successfully grafted onto the chitosan main chain. The peak of aldehyde group (1720 cm−1) was almost disappeared in the CS-AT20 hydrogel. The ratio of absorption intensity at 1644 and 1580 cm−1 was higher in CSAT20 hydrogel than that in the CS-AT20 polymer (Figure 2a), due to the characteristic absorption of the newly formed Schiff base41 (at 1644 cm−1) from amine group of CS and aldehyde of PEG-DA. These results indicated that the CS-AT20 hydrogels were successfully synthesized. It has been widely demonstrated that the electroactive materials could enhance cell proliferation and differentiation.42,43 Electroactivity of the hydrogel was indicated using the UV−vis and cyclic voltammetry (CV). After doping, the UV spectra of the CS-AT20 hydrogel and CS-AT20 solution both showed characteristic peaks at 436 and 307 nm (Figure 2b), which was caused by the formation of polarons in the AT segment. The peak at 800 nm attributed to the localization of

3. RESULTS AND DISCUSSION 3.1. Synthesis and Characterization of Self-Healing Injectable Conducting Hydrogels. To develop a suitable vehicle for cardiac cell delivery, an injectable conductive hydrogel with self-healing property was needed. CS-AT polymer was chosen as a component of our hydrogel because of its good electroactivity and biocompatibility.35 Furthermore, to obtain a self-healing hydrogel, PEG-DA was used as the cross-linking agent, owing to the dynamic covalent Schiff-base linkage between amine group from chitosan and benzaldehyde groups from PEG-DA chain.39 The structure of CS-AT polymer and PEG-DA was characterized by NMR and FT-IR as shown in Figure S1 in the Supporting Information. Figure 1 shows the 17142


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Figure 4. Photographs of self-healing process of CS-AT10 hydrogels. The hydrogel was first injected into a heart shape mold (a) and healed together after 3 h (b−c); this healed hydrogel was cut into two pieces (d), put together and healed into one block (e−f). (g) G′ and G″ on strain sweep and (h) the rheological properties of the hydrogel when alternate step strain switched from 1% to 300%. The scale bar is 1 cm.

ratio between 25% and 50% at the equilibrium state. Owing to the inverse correlation between the cross-linking density and the AT content, the swelling ratio of the hydrogels reduced when the AT content increased. The relationship between the cross-linking density and AT content could be further demonstrated by the morphology of the hydrogels observed under FE-SEM (Figure 3c). The hydrogels showed porous structures and the mesh sizes of the CS-AT8, CS-AT10 and CS-AT20 hydrogels before swelling are about 102.4 ± 9.5 μm, 156.0 ± 21.6 μm and 216.4 ± 26.4 μm, and the mesh sizes of these hydrogels after swelling are about 174.5 ± 31.6 μm, 289.8 ± 32.2 μm and 334.9 ± 41.6 μm, respectively. With the increase of the AT content, the pore size increased. Besides, the pore size of the swollen hydrogels was almost twice than that in the original dried hydrogels. Those large pores in the hydrogels provide a sufficient space for the migration, proliferation and communication of the cells. 3.3. Rheological Behavior of CS-AT Hydrogels. The storage modulus (G′) and loss modulus (G″) of the hydrogels versus time were detected to depict the process of the gelation. The storage modulus (G′) surpassed the loss modulus (G″) in a few seconds and highest modulus value was reached at about 15 min, due to the fast reaction between the −NH2 and −CHO groups to form the Schiff base. Furthermore, to analyze the stability of the hydrogels, the storage modulus G′ and loss modulus G″ versus angular frequency were conducted. When the angular frequency changed from 1 to 100 rad·s−1, the hydrogels’ storage modulus G′ and loss modulus G″ showed no significant changes, indicating the good stability of our hydrogels.

the radical polaron confirmed the generation of the emeraldine salts. The electroactivity of the hydrogel was further investigated using the cyclic voltammetry (CV). The CSAT20 hydrogel was prepared on the ITO glass and doped in a mixture of DMSO and 1 M HCl. Two pairs of well-defined reduction/oxidation peaks appeared in the CV spectra of the hydrogel (Figure 2c), owing to the transition from “leucoemeraldine” to the “emeraldine” state (0.32 V) and that from the “emeraldine” to the “pernigraniline” state (0.54 V). Both the CV and UV−vis results confirmed the good electroactivity of the CS-AT hydrogels. The conductivity of the hydrogels in swollen state was measured using a pocket conductivity meter. As shown in Figure 2d, with the increase of AT content in the CS-AT hydrogels, the hydrogels showed different conductivity from 2.29 × 10−3 to 2.42 × 10−3 S cm−1, which is much higher than that of chitosan (1.6 × 10−4), indicating that AT made a great contribution to the conductivity of the hydrogel. Because the contraction of myocardium was regulated by the conductance of electrical signals, it is necessary for the materials to have a conductivity that matches the native myocardium well. Interestingly, our hydrogel showed a conductivity that was very similar to that of the myocardium (about 1−4 × 10−3 S cm−1),44 making it an suitable candidate for the application in cardiac repair. 3.2. Swelling Ratio and Morphology of the Hydrogels. Swelling behavior of the hydrogels was carried out by immersing the hydrogels in DPBS at 37 °C. As shown in Figure 3a, the equilibrium state of swelling was reached at 120 min and hydrogels with different AT contents showed swelling 17143


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Figure 5. Confocal microscope images of the cells in hydrogels at 1 (a), 2 (b) and 3 (c) days; cell proliferation of the C2C12 myoblasts encapsulated in the hydrogels (d). Scheme of the injection of the hydrogel encapsulated with cells (e). Confocal microscope images of the cells in CS-AT10 hydrogels before (f) and after (g) injection. Live cell numbers in the hydrogel before and after injection (h). 2D confocal microscope image (i) and 3D confocal microscope image (j) of two pieces of hydrogels encapsulating with different type of cells injected together after 3 h. 2D confocal microscope image (k) and 3D confocal microscope image (l) of two pieces of hydrogels encapsulating with different type of cells put together after 3 h. Scale bar was 50 μm.

To analyze the influence of different ratios (R) of −NH2 to −CHO on the mechanical properties of the hydrogels, rheological properties of hydrogels with different molar ratios (R = 0.75, 1, 1.5 and 2) of −NH2 to −CHO were measured. As shown in Figure 3b, the storage modulus G′ showed an enhancement when the R value increased, owing to the higher possibility of the formation of Schiff base when more −CHO groups existed in the system. Besides, compared with the groups with low AT content, the groups with high AT content showed a lower G′ when R value was the same. This is probably because higher content of AT in CS-AT polymer consumed more amine group from chitosan, leading to the decrease of chemical cross-linking point. Compared with the hydrophobic and π−π interaction between AT segment, the chemical crosslinking affects the storage modulus more strongly. Thus, the higher AT content in the CS-AT hydrogels led to the lower storage modulus. 3.4. Self-Healing Behavior of CS-AT Hydrogels. Because of the contraction of the heart, materials used for cardiac tissue repair would bear a high mechanical load, making them easy be damaged. Self-healing materials, with the ability of self-repair, have the ability to prolong the lifetime of the materials.45,46 The self-healing properties of the hydrogels were detected with both visible and qualitative methods for the CSAT10 hydrogel without adding any additive. The hydrogel was injected into small pieces in a heart shape mold and an intact block of hydrogel was obtained in 2 h (Figure 4a−c). Furthermore, when cut into two pieces and put together, the CS-AT10 hydrogels immediately healed together (see Figure 4d−f). The movement of the polymer chains and the dynamic nature of the Schiff base in the hydrogel especially on the surface made them easily re-cross-link together, which led to the fast self-healing ability of the hydrogel.

A qualitative test of the hydrogels’ self-healing ability was conducted using the rheological measurement. The strain amplitude sweep results showed that when the strain was above 150%, the G′ value decreased rapidly, which means that the hydrogel network suffered from a collapse at that point (Figure 4g). Continuously, a strain amplitude sweep method was conducted to analyze the self-healing behavior. As shown in Figure 4h, the storage modulus decreased from ∼2000 Pa to ∼15 Pa when the strain changed from 1% to 300% and it returned to the original level immediately when unloaded the force, suggesting the fast self-healing ability of the hydrogel with the oscillatory shear strain. The end-diastolic pressure of right ventricle is 3−6 mmHg (∼400−800 Pa).47 The self-healing test with the high stress at 400 and 800 Pa was carried out and the hydrogels still showed a good self-healing property (Figure S2 in Supporting Information). Moreover, the native global strain of the heart was about 17% in average48 and the conductive hydrogels remain intact with a stain around ∼250% (Figure 4g). Therefore, our injectable conductive hydrogels could meet the mechanical demand as a carrier for cardiac cell therapy. 3.5. Cell Proliferation in CS-AT Hydrogel. Polyaniline and aniline oligomers have been demonstrated to enhance the proliferation ability of the electrical stimuli sensitive cells, such as mesenchymal stem cells,49,50 C2C12 myoblast,50−52 neuronal cells53,54 and H9c2 myocardial cells.55,56 The CS-AT hydrogels exhibited a good electroactivity and conductivity, which will have a good effect on the proliferation of the cells. Thus, proliferation ability of C2C12 cells was used to measure the biocompatibility of the hydrogels. As shown in Figure 5a−c, an increase of cell amount was observed during the 3 days’ culture from the confocal microscope images. The cell amount calculated in different days showed that the cell number of C2C12 myoblasts increased dramatically for the second day compared to day 1 (Figure 5d), and it continuously increased 17144


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Figure 6. Cumulative delivery profiles for 7 days with different types of cells (a). Images of inversed fluorescent microscope of delivered cells in each day (b).

fusion of the hydrogels and the migration of the cells in the hydrogels. Those results also indicated that this hydrogel still had a good self-healing ability when encapsulated with cells. In most cases, to repair a damaged cardiac tissue, more than one kind of cells is needed, and the carriers that can encapsulate different types of cells are enormously demanded in the cardiac cell therapy. 3.8. Cell Delivery Properties of the Hydrogels. Because of the expectation of these self-healing conductive hydrogels to be used in the cell delivery for cardiac repair, the cell delivery ability of this hydrogel was further investigated. H9c2 cells had been widely used in the regeneration of cardiac tissues.29,60 Skeletal myoblasts had been demonstrated to have the ability to differentiate into myotubes or myocardium which was promising in improving cardiac function postinfarction.8,9,61 Thus, C2C12 myoblasts and H9c2 cardiac myoblasts were chosen to measure the cell delivery ability of the hydrogels. It was tested for 7 days and the delivered cell numbers were monitored every day. Different types of cells showed different releasing properties (Figure 6). Encapsulated with the same cell number, C2C12 cells showed a higher released rate than H9c2 cells, which was probably caused by the faster proliferation of the C2C12 cells. When the same type of cell (C2C12 myoblast) was used, the released cell numbers were impacted by the initial cell number that was encapsulated in the hydrogel, that is, hydrogels with the higher cell concentration encapsulated in the matrix released more cells. These results suggested that the delivered cell number could be effectively tuned by changing the initial cell number. Interestingly, the release curve of C2C12 cells for the first 4 days and of H9c2 cells for 7 days all exhibited a linear-like cumulative delivery behavior, which means that these hydrogels could provide stable cell supply for the cardiac tissue repair. 3.9. In Vivo Cell Retention of the CS-AT Hydrogels. In the traditional way of cell therapy, the cells were injected into the patients’ body as suspension, leading to the low survival ratio and unsatisfactory engraftment efficiency, which greatly limited the therapy efficiency.16 Using biomaterials as cell delivery vehicles could prevent the anoikis in the cell transplantation and enhance the cell engraftment efficiency.1 In this study, the in vivo cell retention ratio was measured by the fluorescence intensity of the cells. At the first day (the day the cells were injected into the hydrogel), the intensity of the cells in both PBS and CS-AT10 hydrogel was the same (Figure S4 in the Supporting Information). The fluorescence intensity of the cells in PBS group showed a sharp decrease by about

for the third day. When using ADMSCs to measure the proliferation in the hydrogels, the cell number after 3 days has an increase by about 50% in the CS-AT10 hydrogel group, while there was no obvious increase of cell number in the CS hydrogel (Figure S3 in the Supporting Information), indicating that the CS-AT hydrogel greatly promoted the proliferation of the ADMSCs compared with the CS hydrogel. Used as the cell carrier in cardiac cell therapy, the ability of keeping a certain amount of cells in the hydrogel is of great significance. This hydrogel not only provided the cells an extracellular matrix to avoid the anoikis of the cells but also promote the proliferation of the cells which could continuously afford cells for the repair of the damaged cardiac tissues. 3.6. Cell Viability in CS-AT Hydrogels after Injection. Injectable hydrogels have been widely used in cell therapy57−59 and the injectable conductive hydrogels have a great potential in cardiac tissue engineering. Considering the goal of using these conductive hydrogels as an injectable cell delivery vehicle, we further examined the cell viability in the hydrogels after injection. Hydrogels encapsulating with C2C12 cells (1 × 106 cells mL−1) were injected through 22 gauge needles (Figure 5e) and the cell number before and after injection was calculated. After 3 h, the morphology of the cells showed no change after injection (Figure 5f,g) and there’s no significant decrease of the cell number in the hydrogel (Figure 5h), indicating that cells still remained good viability after injection. Thus, this conductive hydrogel showed a great possibility to be used as a cell delivery vehicle through injection into the body. 3.7. Two Different Types of Cells Coculture in the Hydrogel. Cardiac tissue consists of different types of cells. In order to repair the damaged cardiac tissue, different types of cells were needed. Stem cells and myoblasts were two types of cells which are widely used in cardiac tissue repair, thus ADMSCs and C2C12 myoblasts were used to coculture in this hydrogel. ADMSCs and C2C12 cells were encapsulated in two pieces of hydrogels separately, and then those two hydrogels were injected together (Figure 5i,j) or put together (Figure 5k,l). After 3 h, for the hydrogels that injected together, the cells in different hydrogels remained a good viability and the cells mixed homogeneously, indicating that this hydrogel could be used to encapsulate different types of cells and mix them. For the hydrogels that were put together, both C2C12 cells and ADMSCs were seen at the healing interfaces of the two pieces of hydrogels (Figure 5k and j), and there’s no obvious boundary of these two cells. Some of ADMSCs (red) were observed in the area of C2C12 cells (green), indicating the 17145


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Figure 7. Antibacterial activities of the CS-AT copolymers and their hydrogels. MICs of the CS-AT solutions for E. coli and S. aureus (a). Surface antibacterial activity of the CS-AT hydrogels (b−d).

79% after 7 days. However, the fluorescence intensity was only reduced by about 16% in the CS-AT10 hydrogel group, which means the injectable conductive hydrogels significantly enhanced the cell retention at the aimed point. 3.10. Antibacterial Property of CS-AT Polymer and Hydrogels. The indiscriminate use of antibiotics could lead to the antibiotic resistance, which is a major public-health problem.62 Thus, controlling use of the antibiotics is of great importance.63 Utilization of the antibacterial biomaterials could partly reduce the dosage of the antibiotics. Research has been carried out on the antibacterial tissue engineering scaffold64−68 and other implanted biomaterials69 in recent years. The introduction of antibacterial ability into biomaterials for the cell therapy would bring benefits, such as reducing the infection risk. Chitosan has been widely used in the field of antibacterial materials due to the positive-charged protonated amino groups in the polymer chain.70 Aniline tetramer and polyaniline were also reported to have antibacterial activity owing to their ability to destroy the bacteria cell-wall.71 Especially, Zhao et al.38 reported that quaternized chitosan grafted polyaniline showed enhanced antibacterial property than quaternized chitosan, which was attributed to the synergistic effect of quaternized chitosan and polyaniline. Thus, we hypothesized that our conductive CS-AT copolymers could show enhanced antibacterial activity than the pure chitosan due to the synergistic contribution of chitosan and aniline tetramer. Indeed, with the introduction and increase of AT in chitosan chain, the MICs of the CS-AT copolymers decreased from 937 to 117 μg/mL for E. coli and decreased from 468 to 117 μg/mL for S. aureus (Figure 7a), which demonstrated that the introduction of AT into chitosan chain enhanced the antibacterial activity for both E. coli and S. aureus. To investigate further whether these materials still have good antibacterial activity in a hydrogel state, the surface antibacterial activity of the CS-AT hydrogels

was also evaluated. As shown in Figure 7b,c,d, CS-AT hydrogels all showed 100% kill efficiency for both bacteria while pure chitosan hydrogel just showed a 50% kill efficiency for E. coli and 87% kill ratio for S. aureus. These results demonstrated that CS-AT copolymers both in solution state and gel state showed more enhanced antibacterial activity than that of pure chitosan. Positive charged oligomer AT and chitosan polymer interact with the negative charged membrane of the bacteria, which are similar to other cationic antimicrobial compounds.72,73 Meanwhile, as a conducting compound, AT also affords a possibility for the transfer of the electron between bacterial cells and itself which results in the dead of bacterial cells.72−74 Thus, the synergetic antibacterial effect of chitosan and AT contributed to the excellent antibacterial activity of CS-AT copolymers and their hydrogels. These conductive CS-AT hydrogels with excellent antibacterial activity are promising candidates as cell delivery vehicles to reduce the risk of bacterial infection during cell therapy process. 3.11. Tissue Adhesive Properties of the Hydrogels. Materials used in cardiac repair would bear a considerable mechanical load due to the contraction of the heart. Thus, to ensure the delivery vehicles would not move during the contraction period, the adhesive properties of them are in great significance. The aldehyde groups in the CS-AT hydrogels, which could react with the amino groups in the native tissues, lead to a good tissue adhesive property of the delivery system. Using the lap-shear test (Figure 8a), the hydrogel’s adhesive properties were investigated by measuring the stress at failure. Figure 8b shows the stress at failure of different hydrogels, and the strength of the hydrogel varied from (3.1 ± 0.69) kPa to (5.1 ± 0.83) kPa, and there’s no statistically significant difference between those groups. The adhesive stress of our hydrogel is much higher than that of the hydrogels reported by Shin et al.14 ((1.356 ± 0.084) kPa). For the hydrogels with 17146


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Figure 8. Representative tissue adhesive image of the lap shear test (a). Adhesive stress of different hydrogels on porcine skin (b). Hydrogels after subcutaneous injection (c), the hydrogels wrapped in the rat’s skin and the hydrogels peeled from the rat’s skin (d). Degradation properties of the hydrogels in vivo (row A shows the hydrogel under the skin and row B shows the hydrogel that peeled from the skin) (e).

5.3 μm in average) around the hydrogel, indicating that this hydrogel has a slight acute inflammatory response. Degradability is an important property that could avoid the risk and pain of the resurgery.75 Grafted conducting polymers have been widely investigated,76 and degradable polymers grafted with aniline oligomers showed a good degradability.42,43 In vivo degradation of the hydrogel was measured for 45 days, and the results are shown in Figure 8e. The size of the hydrogel was reduced and the hydrogel was almost disappeared after 45 days. According to the photographs of the hydrogels, the degradation of the hydrogels exhibited a stable rate, providing a favorable property for their in vivo application.

different AT contents, the stress increased with the AT contents increasing. This may be caused by the decrease of the amino group density in the CS-AT polymer with the increase of AT content. Therefore, the possibility of the reaction between tissue and aldehyde groups increased, leading to the enhancement of adhesive stress. 3.12. Subcutaneous Injectability and in Vivo Degradation of the Hydrogels. A surgery was always needed to implant the original cell delivery vehicles into the human body, and injectable hydrogels reduced pain and risks for the patients. To test the injectable properties, the hydrogels with different AT contents were injected subcutaneously into the SD rats using a 22-gauge needle. After injection, the hydrogels remained round shapes at the injected sites (Figure 8c) and completely cross-linked hydrogels were formed when peeled from the rats’ skin after 15 min (Figure 8d). These results indicated that this hydrogel could be injected into the body keeping a certain shape and the injected hydrogels were integrated into a construct in a short time. The histopathological analysis of the tissues near the hydrogels was measured using the H&E staining. After 7 days, the implanted hydrogels and the tissue around them were excised and the H&E staining result is shown in Figure S5 in Supporting Information. There is a thin fibrous capsule (26.5 ±

4. CONCLUSIONS A series of self-healing injectable conductive hydrogels based on chitosan-g-aniline tetramer have been developed as a promising cell delivery vehicle for cardiac cell therapy. The injectable conductive hydrogels showed a fast self-healing ability due to the dynamic cross-linking network formed by the Schiff-base reaction. The conductivity of these hydrogels ranged from 2.29 to 2.42 × 10−3 cm s−1, which is very close to the native myocardium. These injectable electroactive hydrogels showed good adhesiveness to host tissue and high antibacterial activity. The gelation time, swelling ratio and storage modulus of these 17147


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Structure and Function after Myocardial Infarction. Circulation 2002, 106, I-131−I-136. (9) Pouzet, B.; Ghostine, S.; Vilquin, J.-T.; Garcin, I.; Scorsin, M.; Hagège, A. A.; Duboc, D.; Schwartz, K.; Menasché, P. Is Skeletal Myoblast Transplantation Clinically Relevant in the Era of Angiotensin-Converting Enzyme Inhibitors? Circulation 2001, 104, I223−I-228. (10) Suzuki, K.; Murtuza, B.; Fukushima, S.; Smolenski, R. T.; VarelaCarver, A.; Coppen, S. R.; Yacoub, M. H. Targeted Cell Delivery into Infarcted Rat Hearts by Retrograde Intracoronary Infusion: Distribution, Dynamics, and Influence on Cardiac Function. Circulation 2004, 110, II-225−II-230. (11) Kudo, M.; Wang, Y.; Wani, M. A.; Xu, M.; Ayub, A.; Ashraf, M. Implantation of Bone Marrow Stem Cells Reduces the Infarction and Fibrosis in Ischemic Mouse Heart. J. Mol. Cell. Cardiol. 2003, 35, 1113−1119. (12) Haenel, A.; Ghosn, M.; Schulz, D. G.; Vykoukal, J.; Shah, D.; Raizner, A. E.; Alt, E. Fresh Adipose Tissue Derived Stem Cells Significantly Enhance Ventricular Function in a Chronic Porcine Myocardial Infarction Model. J. Am. Coll. Cardiol. 2015, 65, A1911− A1911. (13) Hong, S. J.; Rogers, P. I.; Kihlken, J.; Warfel, J.; Bull, C.; DeuterReinhard, M.; Feng, D.; Xie, J.; Kyle, A.; Merfeld-Clauss, S.; et al. Intravenous Xenogeneic Transplantation of Human Adipose-Derived Stem Cells Improves Left Ventricular Function and Microvascular Integrity in Swine Myocardial Infarction Model. Catheter. Cardio. Inte. 2015, 86, E38−E48. (14) Shin, J.; Lee, J. S.; Lee, C.; Park, H. J.; Yang, K.; Jin, Y.; Ryu, J. H.; Hong, K. S.; Moon, S. H.; Chung, H. M.; et al. Tissue Adhesive Catechol-Modified Hyaluronic Acid Hydrogel for Effective, Minimally Invasive Cell Therapy. Adv. Funct. Mater. 2015, 25, 3814−3824. (15) Smith, R. R.; Marbán, E.; Marbán, L. Enhancing Retention and Efficacy of Cardiosphere-Derived Cells Administered after Myocardial Infarction Using a Hyaluronan-Gelatin Hydrogel. Biomatter 2013, 3, e24490. (16) Wang, H.; Shi, J.; Wang, Y.; Yin, Y.; Wang, L.; Liu, J.; Liu, Z.; Duan, C.; Zhu, P.; Wang, C. Promotion of Cardiac Differentiation of Brown Adipose Derived Stem Cells by Chitosan Hydrogel for Repair after Myocardial Infarction. Biomaterials 2014, 35, 3986−98. (17) Janse, M. J.; Anderson, R. H.; van Capelle, F. J.; Durrer, D. A Combined Electrophysiological and Anatomical Study of the Human Fetal Heart. Am. Heart J. 1976, 91, 556−562. (18) Anderson, R. H.; Durrer, D.; JANSE, M. J.; VAN CAPELLE, F. J.; BILLETE, J.; BECKER, A. E.; DURRER, D. A Combined Morphological and Electrophysiological Study of the Atrioventricular Node of the Rabbit Heart. Circ. Res. 1974, 35, 909−922. (19) Woo, D. G.; Shim, M.-S.; Park, J. S.; Yang, H. N.; Lee, D.-R.; Park, K.-H. The Effect of Electrical Stimulation on the Differentiation of Hescs Adhered onto Fibronectin-Coated Gold Nanoparticles. Biomaterials 2009, 30, 5631−5638. (20) Serena, E.; Figallo, E.; Tandon, N.; Cannizzaro, C.; Gerecht, S.; Elvassore, N.; Vunjak-Novakovic, G. Electrical Stimulation of Human Embryonic Stem Cells: Cardiac Differentiation and the Generation of Reactive Oxygen Species. Exp. Cell Res. 2009, 315, 3611−3619. (21) Xie, M. H.; Wang, L.; Guo, B. L.; Wang, Z.; Chen, Y. E.; Ma, P. X. Ductile Electroactive Biodegradable Hyperbranched Polylactide Copolymers Enhancing Myoblast Differentiation. Biomaterials 2015, 71, 158−167. (22) Shuai, C.; Deng, J.; Gao, C.; Feng, P.; Peng, S.; Xiao, T.; Deng, Y. Mechanisms of Tetraneedlelike Zno Whiskers Reinforced Forsterite/Bioglass Scaffolds. J. Alloys Compd. 2015, 636, 341−347. (23) Mawad, D.; Stewart, E.; Officer, D. L.; Romeo, T.; Wagner, P.; Wagner, K.; Wallace, G. G. A Single Component Conducting Polymer Hydrogel as a Scaffold for Tissue Engineering. Adv. Funct. Mater. 2012, 22, 2692−2699. (24) Tseng, T. C.; Tao, L.; Hsieh, F. Y.; Wei, Y.; Chiu, I. M.; Hsu, S. h. An Injectable, Self-Healing Hydrogel to Repair the Central Nervous System. Adv. Mater. 2015, 27, 3518−3524.

hydrogels were tuned by the AT content and cross-linking density. Different types of cells were encapsulated into the hydrogel and showed good viability and proliferation. Furthermore, the viability of the cells encapsulated in the hydrogels showed no change after the injection of those hydrogels. The delivery of C2C12 myoblasts and H9c2 cells from those hydrogels showed a linear-like profile and the release rate was controlled by the cell density and cell types. Subcutaneous injection confirmed the injectability and in vivo degradability of those hydrogels. All these results suggest that these self-healing injectable conductive hydrogels with antibacterial activity are excellent vehicles for cell therapy for cardiac tissue repair.

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ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsami.6b04911. NMR and FT-IR of the polymers, dynamic modulus of hydrogel, ADMSCs proliferation in hydrogels, fluorescence intensity of the cells in vivo, and hematoxylin− eosin staining of hydrogel (PDF).

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AUTHOR INFORMATION

Corresponding Authors

*B. Guo. Tel.: +86-29-83395363. Fax: +86-29-83395131. Email: [email protected]. *P. X. Ma. E-mail: [email protected]. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS The National Natural Science Foundation of China (grant number: 21304073) and Xi’an Jiaotong University were acknowledged for financial support of this work.

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REFERENCES

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