copolymer inclusion for tissue engineering repair.

Full Paper

In Situ Electroactive and Antioxidant Supramolecular Hydrogel Based on Cyclodextrin/Copolymer Inclusion for Tissue Engineering Repaira Haitao Cui, Liguo Cui, Peibiao Zhang, Yubin Huang, Yen Wei,* Xuesi Chen*

The injectable electroactive and antioxidant hydrogels are prepared from mixing the tetraaniline functional copolymers and a-cyclodextrin (a-CD) aqueous solution. UV–vis and CV of the copolymer solution showed good electroactive properties. The antioxidant ability of the copolymer is also proved. The gelation mechanism and properties of the system are studied by WAXD, DSC, and rheometer. The encapsulated cells are highly viable in the hydrogels, suggesting that the hydrogels have excellent cytocompatibility. After subcutaneous injection, H&E staining study suggests acceptable biocompatibility of the materials in vivo. Moreover, data shows the injectable electroactive material can effectively accelerate the proliferation of encapsulated cells with electrical stimuli, and the mechanism is also elaborated. Such an injectable electroactive hydrogel would more closely mimic the native extracellular matrix, thereby combining a biomimetic environment of long-term cell survival and electrical signal to support the generation of functional tissue. 1. Introduction H. Cui, L. Cui, Prof. P. Zhang, Prof. Y. Huang, Prof. X. Chen Key Laboratory of Polymer Ecomaterials, Changchun Institute of Applied Chemistry, Chinese Academy of Sciences, Changchun 130022, China E-mail: [email protected] H. Cui University of Chinese Academy of Sciences, Beijing 100049, China Prof. Y. Wei Department of Chemistry, Tsinghua University, Beijing 100084, China E-mail: [email protected] Supporting Information is available at Wiley Online Library or from the author.

a

440

Macromol. Biosci. 2014, 14, 440–450 ß 2013 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

Numerous strategies currently have been used to fabricate proper tissue engineering scaffolds which serve as synthetic extracellular matrix (ECM) to organize cells into a threedimensional architecture and to present stimuli, which direct the growth and formation of a desired tissue.[1–3] Geometrical, chemical, mechanical properties, and electrical signals from the extracellular environment play a key role in regulating cellular activities such as proliferation, migration, and differentiation.[4–6] Recently, conducting polymers have received a great deal of attention, because direct electrical stimuli or electroactivity of substances can influence cellular behaviors.[6,7] According to the literature, the conducting polymers such

wileyonlinelibrary.com

DOI: 10.1002/mabi.201300366

In Situ Electroactive and Antioxidant Supramolecular Hydrogel . . . www.mbs-journal.de

as polypyrrole (PPy) and polyaniline (PANi) enhanced the proliferation and differentiation of neurons, myoblasts, and fibroblasts.[8–11] Moreover, the conducting polymers were widely studied to scavenge free radicals for their potential in biomedical applications.[12,13] Intracellular reactive oxygen species (ROS) have been extensively shown to play an important role in various cancers, cardiovascular diseases and stem cell differentiation,[14–16] and it is also an important inflammatory mediator under pathological conditions and its overproduction is known to be able to subsequently disrupt the cellular oxidant/antioxidant balance leading to the diabetic organ damage, infection, slow tissue regeneration, and wound healing.[17–19] Therefore, the tissue engineering scaffolds containing conducting polymers could reduce the excessive levels of free radicals to protect against the oxidative damage in the process of tissue regeneration. Nowadays, the considerable efforts have been devoted to fabricate degradable conductive scaffold using electroactive oligomers and these scaffolds have showed excellent the biocompatibility, electrical properties, and controllable chemical modification.[20–24] Oligoanilines such as tetraaniline and aniline pentamer have been the most extensively studied for biomedical applications. Through a doping process, the neutral chain becomes positively charged with polarons and bipolarons as the charge carriers for electrical conduction.[25,26] It is more important that the use of oligomers solves drawbacks of the conducting polymers, that is, poor solubility, poor processability, and hard to clearance in vivo.[25,27,28] Despite many strategies for incorporating electroactive segments within the systems such as blending with and grafting into biocompatible matrix, coelectrospinning process, and covalently bonding into the hydrogels have been applied,[22,29–34] the simple fabricated strategy, which has biocompatibility, biodegradability, good formation, and sufficient cell loading still remains a challenge for tissue engineering application. Injectable hydrogels possess both several physical characteristics of the extracellular environment such as high water content, 3D network structure, facilitated mass transfer, and in situ formability, which allows effective encapsulation of various therapeutic molecules or cells and surgical operation in a minimally invasive way.[35–37] In our previous work, a concept of injectable electroactive hydrogels (IEHs) combining the advantages of degradable electrically conducting polymers with the unique properties of injectable hydrogels was proposed and we prepared multi-interactive hydrogels based on tetraaniline functional enantiomeric polylactide-poly(ethylene glycol)-polylactide (PLA-PEGPLA) copolymers.[38] These hydrogels had excellent biocompatibility and could accelerate the proliferation of encapsulated fibroblasts, cardiomyocytes, and osteoblasts with pulsed electrical stimuli. However, in this hydrogel system,

www.MaterialsViews.com

a higher gelation concentration was needed and caused larger viscosity, which led to a difficult operation in experiment. The supramolecular hydrogels resulting from the inclusion complextion of a-cyclodextrin (a-CD) molecules with poly(ethylene glycol) (PEG) and its copolymers have also been extensively studied due to their potential applications in the field of biomedical engineering, as delivery matrixes for drugs or cells.[39–42] These hydrogels are thixotropic, reversible, and injectable through needles. It is more important that the formation of these hydrogels need lower polymer concentration to facilitate operation. It is reported that such supramolecular system, which formed from other types of amphiphilic copolymers such as PEG-poly(e-caprolactone) (PEG-PCL) and PCL-PEG-PCL could facilitate the hydrogels formation and improve their stability due to the strong hydrophobic interaction between hydrophobic blocks.[43,44] The widely used supermolecular hydrogels usually lack biological activity, and the incorporation of dynamic biophysical cues and functions like electrical signal and antioxidant into these injectable hydrogels to further regulate cell behaviors has rarely been reported. Biodegradable and biocompatible poly(D,L-lactic acidco-glycolic acid) (PLGA) and PEG are FDA approved polymers for use in drug delivery carriers and tissue engineering scaffolds. Herein, we synthesized watersoluble carboxyl tetraaniline-poly(D,L-lactic acid-co-glycolic acid)-poly(ethylene glycol)-poly(D,L-lactic acid-co-glycolic acid)-carboxyl tetraaniline (CTA-PLGA-PEG-PLGA-CTA) copolymers with good biodegradability, and the copolymers were used as a building block for constructing supramolecular hydrogels together with a-CD. This strategy resulted in rapidly forming stable supramolecular hydrogels in aqueous solutions and needed a lower gelation concentration to facilitate operation in the experiment. Such an injectable electroactive hydrogel would more closely mimic the configuration of the native ECM, thereby combining a biomimetic environment of long-term cell survival and electrical signal to support the generation of functional tissue.

2. Experimental Section 2.1. Materials N-Phenyl-1,4-phenylenediamine, 1-ethyl-3-(3-dimethyllaminopropyl) carbodiimide hydrochloride (EDC HCl), 4-dimethylaminopyridine (DMAP), stannous octoate (Sn(Oct)2, 95%), PEG with molar masses of 4000 g mol1, a-CD, ammonium persulfate (APS), 2,2-diphenyl-1-picrylhydrazyl (DPPH) and succinic anhydride were purchased from Sigma–Aldrich. Glycolide (GA) and D,Llactide (LA) were obtained from Changchun SinoBiomaterials Co., Ltd., and recrystallized from ethyl acetate for three times before


441

H. Cui, L. Cui, P. Zhang, Y. Huang, Y. Wei, X. Chen www.mbs-journal.de

polymerization. N,N-Dimethylformamide (DMF), toluene, and ethyl acetate were dried and distilled prior to use. Dichloromethane (CH2Cl2), chloroform (CHCl3), ethyl ether, ethanol, methanol, hydrochloric acid (HCl), and ammonium hydroxide (NH3 H2O) were used as received. All chemicals were of analytical grade or higher.

2.2. Synthesis and Characterization of CTA-PLGAPEG-PLGA-CTA Copolymers As shown in Scheme 1, CTA-PLGA-PEG-PLGA-CTA copolymers were prepared by a two-step synthetic procedure. Briefly, PEG was azeotropically dried with toluene for 8 h at first. Then predetermined amounts of GA, LA, PEG, and Sn(Oct)2 (1 wt%) were added into an dried reactor and stirred in toluene at 120 8C for 24 h. The feed amounts of PEG, LA, and GA were varied to provide tunable chain lengths and block ratios of the block copolymers (Table 1). After the reaction was complete, the solution was precipitated in a mixture of cold ethyl ether/ethanol, and then dissolved in CHCl3

and precipitated twice in cold ethyl ether. The product was dried under vacuum for 48 h. Yield of the purified PLGA-PEG-PLGA copolymer was 83%. Tetraaniline was synthesized according to a similar procedure reported in the literature.[20] The carboxylcapped tetraaniline (CTA) was synthesized from the carboxylation reaction of tetraaniline and succinic anhydride in CH2Cl2, and the crude product was washed with distilled water, followed by washing in a Soxhlet extractor with CH2Cl2 till the filtrate became colorless. The product was dried under vacuum for 48 h. Yield of the purified product was 79%. 1H NMR (400 MHz, DMSO-d6, ppm): 12.11 (s, 1H), 9.71 (s, 1H), 7.77(s, 1H), 7.70–7.60 (d, 2H), 7.44–7.33 (d, 2H), 7.30–7.21 (s, 1H), 7.20–7.08 (m, 2H), 7.08–6.78 (m, 8H), 6.67 (m, 1H), 2.73 (m, 2H), 2.35– 2.27 (m, 2H). The molecular weight of CTA was measured by mass spectrometry to be 465.2 (MH þ /e). One millimole of PLGA-PEG-PLGA, 3 mmol of CTA, 10 mmol of EDC, 3 mmol of DMAP, and 20 mL of DMF were added into a dried reactor under nitrogen. The reactor was stirred for 48 h at room temperature. After the reaction, the solution was precipitated in ethanol and was dissolved in CHCl3, filtered, and precipitated in a

Scheme 1. Synthetic route of CTA-PLGA-PEG-PLGA-CTA copolymers.

Table 1. Characterizations of PLGA-PEG-PLGA and CTA- PLGA-PEG-PLGA-CTA copolymers.

Chain structure

M n a) [g mol1]

P1

PEG

4000

P2

PLGA-PEG-PLGA

127-4000-127

2.8/1

9600

1.2

Y

P3

PLGA-PEG-PLGA

386-4 000-386

3.1/1

10 700

1.1

Y

P4

PLA-PEG-PLA

360-4000-360

–

10 500

1.2

Y

P10

CTA-PEG-CTA

–c)

–

9900

1.2

Y

P20

CTA-PLGA-PEG-PLGA-CTA

–c)

2.8/1

10 400

1.3

Y

CTA-PLGA-PEG-PLGA-CTA

c)

–

3.1/1

11 600

1.3

Y

CTA-PLA-PEG-PLA-CTA

–c)

–

11 700

1.3

Y

Sample

P3

0

P40 a)

442

LA/GA –

a)

Mn b) [g mol1]

PDIb)

Water-solubility

9100

1.1

Y

Calculated from the integration of 1H NMR bands; b)Determined by GPC; c)The amount of CTA cannot be accurately calculated by 1H NMR.




mixture of cold ethyl ether/ethanol. Such a dissolution–precipitation process was repeated three times to purify the product. The product was dried under vacuum for 48 h. Yield of the purified CTA-PLGA-PEG-PLGA-CTA copolymer was 86%. 1 H NMR spectra were recorded on a Bruker AV 400 MHz spectrometer. Fourier transform infrared (FT-IR) spectra of samples were recorded on a Bio-Rad Win-IR instrument in the range of 4000–500 cm1. Matrix-assisted laser desorption/ ionization time-of-flight (MALDI-TOF) mass spectra were performed on an AXIMA-CFR laser desorption ionization time-of-flight spectrometer (COMPACT). Gel permeation chromatography (GPC) measurements were carried out with a Waters GPC instrument and DMF was used as an eluent. The molecular weights were calibrated with polystyrene standards. The ultraviolet–visible (UV–vis) spectra of sample solutions were recorded on a UV-2401PC spectrophotometer. Cyclic voltammetry (CV) of sample was conducted on a CHI660 electrochemistry system (CHI, USA) using Ag/AgCl and Pt as the reference and counter electrodes, respectively. The indium tin oxide (ITO) electrode was used as the working electrode and the scan rate was 100 mV s1. The conductivity of the samples at room temperature was measured using a broadband dielectric spectrometer between two-terminal Cu electrodes within the frequency range of 1 Hz–1 MHz (Novocontrol).

To observe the morphology of lyophilized hydrogels in vitro, the environmental scanning electron microscopy (ESEM) was performed on an XL 30 scanning electron microscope (Micrion FEI PHILIPS). For the investigation of the gelation kinetics for CTAPLGA-PEG-PLGA-CTA/a-CD systems, time sweep rheological analyses were performed on a MCR 301 Rheometer (Anton Paar) in oscillatory mode with parallel plate geometry (25 mm diameter, 0.5 mm gap) at 37 8C. The data were collected under a controlled strain g of 1% and a frequency of 1 rad s1. G0 is an elastic component of the complex modulus for measure of the gel-like behavior of a system, whereas G00 is a viscous component of the complex modulus and is a measure of the sol-like behavior of the system. The crossover point between G0 and G00 implied that there was a sol–gel transition. The corresponding time of the crossover from a viscous behavior to an elastic response could be regarded as the GT. To investigate the mechanical property of the resultant hydrogels, the dynamic frequency sweep tests were also conducted. The frequency applied to the hydrogel sample increased from 1 to 100 rad s1. In addition, the steady rate sweep tests were carried out to investigate the shear thinning properties of the resultant hydrogels. In this case, the hydrogel samples were allowed to consolidate for 5 h before the measurements. To prevent the evaporation of water, the outer edge of the sample was sealed by a thin layer of silicon oil.

2.3. Antioxidant (Radical Scavenging) Ability Assay

2.5. In Vitro Cell Encapsulation and Viability

The antioxidant assays (radical scavenging activity) of the material were measured according to a similar procedure reported in the literature.[13] The copolymer solution was applied to 3.0 mL of 100 mM DPPH methanol solution. The reaction mixture was vortexed for 30 s and incubated at dark condition. The absorbance of the reaction mixture was measured at 516 nm using UV–vis spectrophotometer with in the time limit of t ¼ 0 min to 3 h. The DPPH degradation was calculated using the formula: percentage of DPPH scavenging (%) ¼ (AB – AS)/AB 100, where AB is the absorption of the blank (DPPH/methanol) and AS is the absorption of the sample (DPPH/sample/methanol).

Before encapsulation, CTA-PLGA-PEG-PLGA-CTA solution and a-CD solution were sterilized under UV light for 30 min. Rat cardiomyocytes H9c2 cells suspension with a density of 5 104 per well in 24-well plates was added into 200 mL (v/v ¼ 1:1) mixture solution. After gelation, 500 mL of culture medium (Dulbecco’s modified Eagle medium (DMEM) (Gibco) supplemented with 10% fetal calf serum (Gibco) and 100 U mL1 penicillin–streptomycin (Sigma)) were added and changed every other day. The plates were incubated at 37 8C in humidified 5% CO2 atmosphere. After 1 and 3 d of culture, qualitative viability assay were performed using live-dead assay kit. At predesigned day post-encapsulation, the cell/hydrogel complexes were treated with calcein AM (2 mM) and propidium iodide (4 mM) for 30 min. The cells were observed under an inverted fluorescent microscope (TE2000U, Nikon) and analyzed using ‘‘NIH ImageJ’’ software.

2.4. In Vitro Hydrogel Formation and Its Characterization The vial inverting approach was employed to determine the gelation time (GT) at room temperature. Briefly, the CTA-PLGA-PEGPLGA-CTA copolymer solutions and a-CD solution with different concentrations were prepared by using distilled water as the solvent, respectively. Subsequently, the equal amounts of polymer solution and a-CD solution were mixed in the vial, and the sample was defined as a ‘‘gel’’ in the case of no visual flow within 30 s by inverting the vial. For the investigation of supramolecular gelation mechanism, the wide-angle X-ray diffraction (WAXD) measurement was carried out on a Bruker D8 Advance X-ray diffractometer, using Cu Ka radiation and the scattering angle ranged from 2u ¼ 58–308 at room temperature. Differential scanning calorimetry (DSC) was carried out on a TA Instrument Q100 under nitrogen atmosphere, at the heating/cooling speed of 10 8C min1 from 20 to 250 8C.


2.6. Electrical Stimuli and Proliferation Assays According to our previous work, cardiomyocytes was more sensitive to electrical signal. H9c2 cells were used to investigate the influence of electroactive hydrogel on the cell proliferation stimulated by pulsed electrical signals. The methods of electrical stimuli such as device, parameters and time are same to our previous work.[38] Before encapsulation, CTA-PLGA-PEG-PLGA-CTA (doped with HCl) solution and a-CD solution were sterilized under UV light for 30 min. The cells were seeded into various gels at a density of 5 104 cells per well in 24-well plates. After the predetermined period, the incubation medium was removed and WST-8 solution (10% v/v in medium) (Cell Counting Kit-8) was added to each well. After 4 h of incubation, the absorbance value at


443


450 nm was measured on multifunction microplate scanner (Tecan Infinite M200). The intracellular free calcium concentration ([Ca2þ (i)]) of H9c2 cells was assessed at 5-d by fluo-4/AM substrate (Fanbo biochemical). After gel fragmentation, cells were collected from each well by centrifugation. Fluo-4 substrate was added and cells were incubated for 30 min. Fluorescence intensity was measured on multifunction microplate scanner using an excitation wavelength of 494 nm and an emission wavelength of 516 nm.

2.7. Animal Procedure and In Vivo Evaluation The animal experiments were carried out according to the NIH Guide for the Care and Use of Laboratory Animals, provided by Jilin University, Changchun, China. Sprague-Dawley (SD) rats were anesthetized, and then the hydrogels composed of the CTA-PLGAPEG-PLGA-CTA and a-CD aqueous solution (0.3 mL) were subcutaneously injected into rats by a syringe with a 23-gauge needle. At the designated time intervals, the animals were sacrificed, and the injection site was carefully cut open. Then the photographs of in situ gel formation were taken. The surrounding tissues of the hydrogels were surgically removed and histologically processed using hematoxylin–eosin (H&E) stains for the examination of inflammatory responses of the hydrogels in rats.

3. Results and Discussion The CTA-PLGA-PEG-PLGA-CTA copolymers were synthesized by the two-step reactions, according to the synthetic procedure in the Scheme 1, in which the first step was the ring-opening polymerization of the monomer GA and LA using PEG as the initiator and the second step was condensation coupling reaction of the PLGA-PEG-PLGA triblock and carboxyl-capped tetraaniline (CTA). Table 1 shows the feed ratios and the results of structure characterization of the copolymers. Supporting Information, Figure S1 shows 1H NMR spectra of typical a-CD, CTAPLGA-PEG-PLGA-CTA copolymer (P20 ) and lyophilized P20 /aCD hydrogel formed from 10 wt% of P20 and 20 wt% of a-CD (or lyophilized 10 wt% P20 /20 wt% a-CD hydrogel) in DMSO. The peaks at 5.20 and 1.45 ppm were assigned to a methine

proton and methyl protons of the LA units, respectively. The peaks at 4.80 and 3.50 ppm were attributed to methylene protons of the GA units and PEG, respectively. The proton of tetraaniline in the CTA blocks appeared at 7.80–6.80 ppm. Based on the comparison of the molecular weight of P copolymers, these P0 samples all had molecular weights greater than P copolymers, and the calculated CTA/triblock ratios were higher than 1.5. In the experiment, the excessive CTA were added in the system, the reaction was conducted as full as possible. Due to the calculated deviation from GPC data, it is possible that the coupling degree is correspondingly lower than theoretical value. The combined data of the 1H NMR, GPC, and FT-IR spectra confirmed the successful synthesis of the copolymers. The FT-IR spectra of a-CD, P20 , and lyophilized 10 wt% P20 /20 wt% a-CD hydrogel are shown in Supporting Information, Figure S2. In general, the weak peak at 1750 cm1 was attributed to the C5 5O asymmetric stretching mode of ester in the copolymer and inclusion complexes (ICs). Furthermore, the peaks at 2854 and 2927 cm1 corresponded to the methylene groups of PEG, which was slightly weakened in ICs. In addition, a sharp peak at 1058 cm1 corresponded to the ether group of —CH2—O—CH2— which disappeared in ICs. In the spectra of P20 and P20 /a-CD, the typical absorption peak of tetraaniline observed at 1510 cm1 (s, —N—B—N—) and 1605 cm1 (s, —N5 5Q5 5N—) was attributed to the benzenoid unit and the quinoid unit of the CTA blocks in the chain, respectively. The data showed the successful synthesis of the copolymers and indicated that the a-CDs were successfully threaded onto the linear PEG chain. The conductivity of the sample films was also evaluated and the results were shown in Supporting Information, Figure S3. The electrochemistry of the CTA-PLGA-PEG-PLGA-CTA copolymer was investigated by CV and UV–vis spectra. Figure 1A shows the CV of the P20 solution obtained by directly dissolving P20 into 1.0 mol L1 HCl aqueous solution. The P20 solution showed one pair of obvious reversible redox peaks, and the mean peak potential E1/2 was 450 mV, which was due to the transition from the leucoemeraldine

Figure 1. A) CV of P20 in 1 M HCl aqueous solution. B) UV–vis spectra of P20 aqueous solution oxidized by APS. C) UV–vis spectra of P20 aqueous solution in LM state, EM state and EMS state (doped with 1 M HCl aqueous solution).

444




(LM) state to the emeraldine (EM) state. Figure 1B shows the UV–vis spectra of P20 oxidized by APS in DMF. It exhibited a stepwise oxidation process from the LM to the EM state. Compared with the LM copolymer that showed only one peak at 310 nm, the EM copolymer showed two peaks at around 305 and 575 nm, which were attributed to the p–p transition of the benzene ring and the benzenoid to quinoid (pB–pQ) excitonic transition, Figure 2. A) The time dependent of antioxidant activity test of P2’. B) The antioxidant respectively. These observations were activity of different samples. also confirmed by FT-IR spectra. The predominance of the absorption band of DPPH with time increasing. The mechanism of at 1510 cm1 to over that at 1605 cm1 for the LM copolymers. When the LM copolymer was oxidized, the the reaction between aniline and DPPH has been 1 intensities of the absorption band at 1605 cm for the investigated by many researchers. After reacting with DPPH, the EMS form got converted into higher oxidized EM copolymer is approximately equal to the one at 1 form of oligoaniline until it could not possess any hydrogen 1510 cm , which are consistent with intrinsically atom for reducing DPPH. Figure 2B shows a comparative oxidized structure.[25] The characteristic absorbance graph of the antioxidant activity of the same weight spectra of the EM P20 copolymer in aqueous solution at (1.0 mg) of different samples. It has been observed that different oxidation states were recorded as shown in the antioxidant property of the samples increases with Figure 1C. The LM sample showed one peak at 310 nm, the increase in the proportion of TA block. This can be and the EM sample showed the blue-shift peak at attributed to the fact that TA owing to its redox active 305 nm and the appearance of a new peak at 575 nm. nature is efficient in scavenging DPPH and as such with When the EM P20 copolymer was doped with the 1 M HCl the increasing proportion of TA there is a corresponding aqueous solution, the peak at 575 nm decreased and almost disappeared. At the same time, the p–p transition increase in the antioxidant activity of the material. For such tetraaniline functional block copolymers, it peak at about 302 nm persisted and decreased in its could be dissolved in water to form homogeneous solution. intensity, and the polaron peaks at 410 nm and the In particular, their aqueous solutions could be transformed localized polaron peak at 800 nm confirmed the into the invertable hydrogels in some cases when a-CD generation of emeraldine salts (EMS) and the ability of solution was introduced. Although solubility of a-CD was conducting electrons of copolymer. After the copolymer lower at room temperature, it could increase with the was doped, the extended chain conformation increases the temperature rising. Figure 3A shows the solution states of conjugation length, which leads to the disappearance of 10 wt% aqueous P20 solution before and after mixing pB–pQ excitonic transition and the delocalization of [20,25] polarons. with 20 wt% aqueous a-CD solution. Depending on CTAThe results of CV and UV–vis identified PLGA-PEG-PLGA-CTA and a-CD amounts, this gelation the good electroactivity of the copolymer. could occur at a certain time under mild condition. When The ROS produced by the injured tissue itself or by aqueous a-CD solution with the concentration of 20 or infiltrating inflammatory cells lead to new cellular damage. 25 wt% was mixed with aqueous copolymer solution with Since oligoanilines could react with radical species, we the concentration of 10 or 20 wt% at room temperature, a hypothesized that the copolymer could reduce free radicalrapid physical gelation occurred instantaneously as shown mediated oxidative damage through removing O2 in in Supporting Information, Table S1. This phenomenon the process of ROS extracellular transmission. In order to may be attributed to the formation of ICs between PEG evaluate the effectiveness of scavenger activity of the blocks and a-CD in their aqueous mixed system and ICs copolymer, the ability of the compound to scavenge may be thought to aggregate into microcrystals, which act the stable DPPH free radical as a model was tested. In as physical cross-links and induce the formation of its radical form, DPPH showed an absorbance peak at the supramolecular polymer network. In the comparison 516 nm, which progressively disappeared as a function experiment without hydrophobic blocks by investigating of DPPH reduction. The time-dependent antioxidant the gelation process of P1/a-CD, the much shorter GT were activity of the copolymer was tested as shown in required, however, the hydrogel was easy to dissociate Figure 2A. The results of the reaction of DPPH with P20 after 1 or 2 d. Therefore, the hydrophobic interaction of revealed a progressive decrease in the absorption band



445


Figure 3. A) Photographs of 10 wt% P20 aqueous solution and 10 wt% P20 /20 wt% a-CD hydrogel. B) Schematic illustration of the gelation mechanism for the formation of the supramolecular hydrogel.

the PLGA and CTA block copolymer improved hydrogel stability. It could be explained that the hydrophobic chains hindered a-CD to thread onto and off PEG blocks in the formation process of the supramolecular hydrogel. Compared with P30 that had similar molecular weight, P40 composed of PLA blocks showed poor solubility and longer GT. Thus GA units could improve solubility of the hydrophobic blocks. In water, due to the hydrophobic aggregation between PLGA and CTA blocks, the amphiphilic block copolymer tended to aggregate together to form core–corona structured particles. As a result, a-CD preferred to thread onto PEG block rather than hydrophobic block and the hydrophobic blocks left unthreaded.[34] The hydrophobic interaction between uncovered hydrophobic blocks facilitated the hydrogel formation and improved the hydrogel stability.[44] The driving force for gelation was a combination of the inclusion complexation and the hydrophobic interaction.

Above all, the schematic structure of the supramolecular hydrogels formed is illustrated in Figure 3B. To confirm this mechanism, the WAXD patterns and DSC curves of a-CD, P20 , and lyophilized 10 wt% P20 /20 wt% a-CD hydrogel were measured. As shown in Figure 4A, the a-CD was characteristic of multiple diffraction peaks corresponding to the crystalline form, while P2 and P20 were characteristic of two main strong peaks at 19.308 and 23.508. For the supramolecular hydrogels investigated, however, both patterns exhibited ˚) two characteristic diffraction peaks at 19.908 (d ¼ 4.44 A ˚ and 22.808 (d ¼ 3.96 A), which were different from that of either a-CD or copolymers. According to a similar result reported in the literature, the patterns represented the channel-type structure of a crystalline-necklace-like complex of a-CD and PEG.[45] While a-CD did not show any melting behavior over our experimental temperature range, a distinct endothermic peak was observed at 50.5 8C for P2 (or at 49.3 8C for P20 ), corresponding to the melting point of PEG chain crystalline (Figure 4B). After the gelation, the melting point was shifted to a lower temperature and the melting enthalpy was also decreased, implying that PEG chain was included in the channel of a-CD so as to impede the crystalline phase formation. These results indicated that the PEG chains participated in the formation of the channel-typed crystalline structure together with the a-CD molecules in the hydrogel, which might act as physical cross-linker and then resulted in the supramolecular hydrogel networks. From the viewpoint of practical application, the supramolecular hydrogel samples were studied by dynamic mechanical rheology at 37 8C. To investigate the effects of hydrophobic chain length and the concentration of a-CD or copolymer on the supramolecular gelation, the time sweep measurements for the viscoelastic properties of each system were carried out. As shown in Supporting Information, Figure S4, with the time increasing, the G0 values of the electroactive hydrogel

Figure 4. A) The WAXD patterns and B) the DSC curves obtained from the second heating (a, a-CD; b, P2; c, lyophilized 10 wt% P2/20 wt% a-CD hydrogel; d, P20 ; e, lyophilized 10 wt% P20 /20 wt% a-CD hydrogel).

446




(15 wt% P20 /20 wt% a-CD) and its HCl doped product were higher than that of the pure hydrogel (P2/a-CD). This date suggested that the introduction of PLGA and CTA blocks positively affected the strength and viscosity of the hydrogels. Because each of such hydrophobic domains connected many PEG chains, we called it a supra-crosslink. Because CTA was coupled to the triblock copolymer, the hydrophobic and rigid CTA provided additional physical crosslinks via their hydrophobic aggregation, hydrogen banding, and p–p stacking.[38] In other words, it was concluded that the CTA introduction greatly improved the hydrogel strength. The time dependences of the storage modulus (G0 ) and loss modulus (G00 ) for various copolymer/ a-CD systems were also investigated. As shown in Figure 5A, with the increase of Mw of PLGA block in the copolymers, the GT of the hydrogel formed from 15 wt% of

copolymer and 20 wt% of a-CD increased whereas the strength of hydrogel was improved that the G0 value increased from 10.0 to 120.0 kPa. When the concentration of P30 increased from 5 to 10 wt% after mixing with 25 wt% a-CD, the G0 value increased from 10.0 to 65.0 kPa. When the concentration of a-CD increased from 15 to 25 wt% after mixing with 10.0 wt% P30 , the G0 value increased from 15.0 to 65.0 kPa (Figure 5B). These results indicated that a higher copolymer or a-CD amount would be favorable for the supramolecular gelation, which might be attributed to the increase of interaction sites. Figure 6A presents the storage modulus (G0 ) evolutions of the hydrogel formed from 15 wt% of P20 and 20 wt% of a-CD as a function of frequency. The G0 value of hydrogel exhibited a substantial elastic response and was greater than the loss modulus G00 over the entire range

Figure 5. Time sweep measurements of the storage modulus (G0 ) and loss modulus (G00 ) for various copolymer/a-CD systems. A) Effects of Mw of PLGA block on the supramolecular gelation kinetics (&, G0 and &, G00 of 15 wt% P10 /20 wt% a-CD hydrogel; &, G0 and &, G00 of 15 wt% P20 /20 wt% aa-CD hydrogel; ~, G0 and &, G00 of 15 wt% P30 /20 wt% a-CD hydrogel). B) Effects of copolymer and a-CD amounts on the supramolecular gelation kinetics (&, G0 of 5 wt% P30 /15 wt% a-CD hydrogel; ~, G0 of 5 wt% P30 /20 wt% a-CD hydrogel; 3, G0 of 5 wt% P30 /25 wt% a-CD hydrogel; *, G0 of 10 wt% P30 /15 wt% a-CD hydrogel; !, G0 of 10 wt% P30 /20 wt% a-CD hydrogel; ", G0 of 10 wt% P30 /25 wt% a-CD hydrogel).

Figure 6. A) The storage modulus (!, G0 ) and the loss modulus (&, G2) evolutions of the 15 wt% P20 /20 wt% a-CD hydrogel as a function of frequency. B) Change of steady shear viscosity as a function of shear rate for the supramolecular hydrogels (*, h of 15 wt% P20 /20 wt% a-CD hydrogel; &, h of 15 wt% P2/20 wt% a-CD hydrogel).



447


of frequency. The fact confirmed that the hydrogels were well cross-linked with insignificant sol fraction. Figure 6B shows the viscosity change of hydrogel formed from 15 wt% of P20 and 20 wt% of a-CD as a function of shear rate. The viscosity of the P20 /a-CD hydrogel displayed progressive decrease in viscosity from 5000 to 0.0008 Pa s with increasing shear rate from 1 to 50 s1. It was because that the self-assembly of ICs were continuously break under shearing, which led to a substantial decrease in the degree of crosslinking. Moreover, it was observed that a disrupted sol phase could be turned reversibly into a gel after shearing. The shear-thinning and thixotropic properties attributed to the physical nature of the supramolecular crosslinks are important for injectable matrix through needles. The morphology of lyophilized hydrogels was observed by SEM as shown in Supporting Information, Figure S5. The hydrogels were highly porous networks due to the high water content. The pores of 10 wt% P1/20 wt% a-CD hydrogel were homogeneous and the average pore size was around 3 mm. The morphology of 10 wt% P2/20 wt% a-CD hydrogel was very different, which showed loose holes and the average pore size was larger. The 10 wt% P10 /20 wt% a-CD hydrogel and 10 wt% P20 /20 wt% a-CD hydrogel had pore sizes of 10 mm, indicating that CTA blocks had a significant influence on the pore size of the lyophilized hydrogels. Moreover, the image of hydrogels before and after coupled with CTA blocks showed different interconnected network structures. The interactions and the rigidity of CTA blocks induced separation of interconnected network to form macropore structure. Because of the appropriate flowability and storage modulus values, 15 wt% P20 /20 wt% a-CD hydrogel was applied to the in vitro and in vivo experiments and 15 wt% P2/20 wt% a-CD hydrogel was used as a control experiment. Because of mild encapsulation, the supramolecular hydrogel was expected to maintain cells viability. To confirm this, H9c2 cells were incorporated in the hydrogels and the cell-encapsulated hydrogels were incubated in DMEM up to 3 d. Cell viability was determined by live– dead staining. The viable cells (stained green) and dead cells (stained red) were observed as shown in Figure 7A. Staining of cultures 1 and 3 d after encapsulation indicated that the cell viability was preserved and not significantly influenced in the electroactive hydrogel scaffold. These results showed that the hydrogels had a good cytocompatibility. Due to the wide variety of bioelectrical presence and its significant influence on in vivo tissue, researchers have been exploring the potential of using this stimulus in vitro. Electrical stimuli that provide better control over the in vitro and in vivo proliferation and differentiation of cells such as neurons, myoblasts, fibroblasts was proved. In our previous paper, it has been demonstrated that the IEHs

448

Figure 7. A) Live (green)/dead (red) staining images of H9c2 cells entraped in 15 wt% P20 /20 wt% a-CD hydrogels for different times. B) The cell proliferation activity and C) the [Ca2þ (i)] of hydrogels for different times (pure hydrogel, 15 wt% P2/20 wt% a-CD hydrogel; electroactive hydrogel, 15 wt% P20 /20 wt% a-CD hydrogel; [], without electrical stimuli; [þ], with electrical stimuli).

can promote the proliferation of fibroblasts (L929 cells), cardiomyocytes (H9c2 cells), and osteoblasts (Mc3T3-E1) upon stimulation by electrical signals, especially for cardiomyocytes. Therefore, the proliferation activity was quantitatively determined to measure the total population of H9c2 cells growing into the hydrogels with or without electrical stimuli for 5 d as shown in Figure 7B. A general increase in the cell proliferation rate was observed in the electroactive hydrogel groups, when compared to control groups (pure hydrogel). The cell proliferation rate at 5 d had more significant difference compared with that at 3 d when the cells were seeded in electroactive hydrogels. Cell proliferation rate of the 15 wt% P20 /20 wt% a-CD hydrogel group with electrical stimuli in various cells was highest as expected, compared with that of the 15 wt% P2/20 wt% a-CD hydrogel group without electrical stimuli at 5 d stimulation period. Many studies elaborated electrical stimuli affect cells behaviors through the calcium/calmodulin pathway.[46–48] A key component in this process was the intracellular calcium level. The elevated calcium level in both cases activated the cytoskeletal calmodulin, resulting in enhanced proliferation, increased vascular endothelial growth factor (VEGF), and transforming growth factor (TGF)-b expression.[48–50] As shown in Figure 7C, cells were seeded for 5 d, compared with cells in the hydrogel groups without electrical stimuli, a higher [Ca2þ (i)] was observed for those exposed to electrical field. Moreover, highest [Ca2þ (i)] was observed when cells were seeded in electroactive hydrogels. The increase of [Ca2þ (i)] led to an activation of calmodulin, which is responsible for many




calcium-mediated processes including proliferation and differentiation of cells.[48,49] Thus, modulation of calcium transport thorough voltage grated channels, as a result of the electrical stimulation, could be mediating cellular behaviors. We assumed the electroactivity of the hydrogel accelerated the chemical and energy exchange between cells and its surroundings with electrical stimuli, which could be desirable and useful for tissue engineering scaffold. In vivo gel formation and degradation were observed in SD rats. The gels were injected into the subcutaneous layer of rats. At 1 d post-injection, the integrated hydrogels were observed in situ in the subcutaneous layer. The hydrogels was kept up to half one month with the size decreasing gradually. After three weeks, the gels almost disappeared at the injection sites. Typical photographs are shown in Figure 8A. The inflammatory response of the implanted hydrogels was studied by H&E staining of the surrounding tissue at different intervals (Figure 8B). Acute inflammatory reactions by lymphocytes and macrophages were observed in the initial days after injection, neutrophils infiltrated into the implanted gel, and the thin fibroblastic capsules formation was observed

Figure 8. A) In vivo hydrogels formation in the subcutaneous tissue at different intervals. B) Images of H&E stained of surrounding tissues at indicated days for examination of the inflammation reaction (F, fibroblastic capsules; I, lymphocytes; M, macrophages, N, neutrophils).


around the gel. After two weeks, the neutrophils disappeared and the density of fibroblastic capsules around the implant diminishes. Notably, at three weeks postinjection, the histology of the tissue sample surrounding the injection site was almost restored to the normal tissue (control). Although the inflammatory reaction was found surrounding the injection site, it was reduced significantly and eventually eliminated accompanying the degradation of the hydrogels. Hence, we concluded that the electroactive hydrogel exhibited acceptable biocompatibility in vivo, which might be suitable for in vivo applications.

4. Conclusion Supermolecular hydrogels formed from a-CD and CTAPLGA-PEG-PLGA-CTA copolymers were successfully prepared. Based on the UV–vis spectra and CV results in solution, these copolymers possessed electroactivity and antioxidant ability. The WAXD and DSC characterization results showed that the main driving force for in situ gelation was the host–guest interaction between a-CDs and PEG segments. The crystal inclusion complexation, hydrophobic aggregation, hydrogen bonding, and p–p stacking acted as physical crosslink interactions. The morphology of the electroactive hydrogel showed loose holes and the average pore size was larger than the pure hydrogel. The rheological behavior of the hydrogels confirmed that the introduction of CTA and PLGA segments increased the storage modulus and improved hydrogel stability. The storage modulus value of 15 wt% P20 /20 wt% a-CD hydrogel increased up to 50-fold compared with that of pure hydrogel. For the purpose to use the hydrogel as tissue engineering scaffold, the in vitro cytotoxicity and in vivo biocompatibility of the hydrogels were evaluated and the positive results were obtained. At three weeks postinjection, the histology of the tissue sample surrounding the injection site was almost restored to the normal tissue. The cell culture results showed that the electroactive hydrogels could accelerate the proliferation of cells with electrical stimuli compared with pure hydrogels. However, the application of these supramolecular hydrogels as tissue engineering scaffolds still faces some problems such as relatively fast gel erosion and poor mechanical property. Therefore, further improvements in material design to obtain enhanced mechanical property and stability in vivo are required for widespread utility of shear-thinning hydrogels for biomedical applications.

Acknowledgements: This work was supported by grants from National Natural Science Foundation of China (Project 50973109, 51103149, 51233004, and 51021003) and the Ministry of Science


449


and Technology of China (International Cooperation and Communication Program 2011DFR51090).

Received: August 7, 2013; Revised: September 27, 2013; Published online: November 8, 2013; DOI: 10.1002/mabi.201300366 Keywords: antioxidants; electroactivity; injectable hydrogels; tetraaniline; tissue engineering

[1] [2] [3] [4] [5] [6] [7] [8]

[9] [10] [11] [12]

[13] [14]

[15] [16] [17] [18]

[19] [20] [21]

450

P. X. Ma, Adv. Drug Delivery Rev. 2008, 60, 184. L. J. Lee, Ann. Biomed. Eng. 2006, 34, 75. V. L. Tsang, S. N. Bhatia, Adv. Drug Delivery Rev. 2004, 56, 1635. F. Brandl, F. Sommer, A. Goepferich, Biomaterials 2007, 28, 134. H. L. Lim, J. C. Chuang, T. Tran, A. Aung, G. Arya, S. Varghese, Adv. Funct. Mater. 2011, 21, 55. N. K. Guimard, N. Gomez, C. E. Schmidt, Prog. Polym. Sci. 2007, 32, 876. R. Ravichandran, S. Sundarrajan, J. R. Venugopal, S. Mukherjee, S. Ramakrishna, J. R. Soc. Interface 2010, 7, S559. M. Y. Li, P. Bidez, E. Guterman-Tretter, Y. Guo, A. G. MacDiarmid, P. I. Lelkes, X. B. Yuan, X. Y. Yuan, J. Sheng, H. Li, C. X. Song, Y. Wei, Chin, J. Polym. Sci. 2007, 25, 331. M. Gizdavic-Nikolaidis, S. Ray, J. R. Bennett, A. J. Easteal, R. P. Cooney, Macromol. Biosci. 2010, 10, 1424. Y. Wan, A. Yu, H. Wu, Z. Wang, D. J. Wen, Mater. Sci.: Mater. Med. 2005, 16, 1017. N. Gomez, C. E. Schmidt, J. Biomed. Mater. Res. A 2007, 81, 135. M. Gizdavic-Nikolaidis, J. Travas-Sejdic, G. A. Bowmaker, R. P. Cooney, C. Thompson, P. A. Kilmartin, Curr. Appl. Phys. 2004, 4, 347. J. P. Saikia, S. Banerjee, B. K. Konwar, A. Kumar, Colloids Surf. B 2010, 81, 158. X. Wang, Q. Tu, B. Zhao, Y. An, J. C. Wang, W. Liu, M. S. Yuan, S. M. Ahmed, J. Xu, R. Liu, Y. Zhang, J. Wang, Biomaterials 2013, 34, 1155. K. Sugamura, J. F. Keaney, Free Radic. Biol. Med. 2011, 51, 978. L. Zhang, Y. Zhou, J. Zhu, Q. Xu, Vascul. Pharmacol. 2012, 56, 280. P. Pacher, J. S. Beckman, L. Liaudet, Physiol. Rev. 2007, 87, 315. W. Hoetzenecker, B. Echtenacher, E. Guenova, K. Hoetzenecker, F. Woelbing, J. Bruck, A. Teske, N. Valtcheva, K. Fuchs, M. Kneilling, J. H. Park, K. H. Kim, K. W. Kim, P. Hoffmann, C. Krenn, T. Hai, K. Ghoreschi, T. Biederman, M. Rocken, Nat. Med. 2012, 18, 128. Y. F. Feng, L. Wang, Y. Zhang, X. Li, Z. S. Ma, J. W. Zou, W. Lei, Z. Y. Zhang, Biomaterials 2013, 34, 2234. L. H. Huang, X. L. Zhuang, J. Hu, L. Lang, P. B. Zhang, Y. Wang, X. S. Chen, Y. Wei, X. B. Jing, Biomacromolecules 2008, 9, 850. J. Hu, L. H. Huang, X. L. Zhuang, P. B. Zhang, L. Lang, X. S. Chen, Y. Wei, X. B. Jing, Biomacromolecules 2008, 9, 2637.

[22] B. L. Guo, A. Finne-Wistrand, A. C. Albertsson, Chem. Mater. 2011, 23, 1254. [23] L. Lang, X. L. Zhuang, Y. D. Liu, X. S. Chen, Y. Wei, Acta Polym. Sin. 2010, 956. [24] L. Lang, X. L. Zhuang, Y. D. Liu, P. B. Zhang, X. S. Chen, Y. Wei, Chem. J. Chin. Univ. (Chin. Ed.) 2011, 32, 411. [25] L. H. Huang, J. Hu, L. Lang, X. Wang, P. B. Zhang, X. B. Jing, X. H. Wang, X. S. Chen, P. I. Lelkes, A. G. MacDiarmid, Y. Wei, Biomaterials 2007, 28, 1741. [26] L. Chen, Y. Yu, H. Mao, X. Lu, W. Zhang, Y. Wei, Mater. Lett. 2005, 59, 2446. [27] J. Hu, L. H. Huang, L. Lang, Y. D. Liu, X. L. Zhuang, X. S. Chen, Y. Wei, X. B. Jing, J. Polym. Sci. A: Polym. Chem. 2009, 47, 1298. [28] H. T. Cui, Y. D. Liu, M. X. Deng, X. Pang, P. B. Zhang, X. H. Wang, X. S. Chen, Y. Wei, Biomacromolecules 2012, 13, 2881. [29] Y. D. Liu, J. Hu, X. L. Zhuang, P. B. Zhang, Y. Wei, X. H. Wang, X. S. Chen, Macromol. Biosci. 2012, 12, 241. [30] Y. D. Liu, H. T. Cui, X. L. Zhuang, P. B. Zhang, Y. Cui, X. H. Wang, Y. Wei, X. S. Chen, Macromol. Biosci. 2013, 13, 356. [31] Y. D. Liu, J. Hu, X. L. Zhuang, P. B. Zhang, X. S. Chen, Y. Wei, X. H. Wang, Macromol. Biosci. 2011, 11, 806. [32] B. L. Guo, A. Finne-Wistrand, A. C. Albertsson, Macromolecules 2011, 45, 652. [33] B. L. Guo, Y. Sun, A. Finne-Wistrand, K. Mustafa, A. C. Albertsson, Acta Biomater. 2012, 8, 144. [34] B. L. Guo, A. Finne-Wistrand, A. C. Albertsson, Biomacromolecules 2011, 12, 2601. [35] S. R. Van Tomme, G. Storm, W. E. Hennink, Int. J. Pharm. 2008, 355, 1. [36] M. K. Nguyen, D. S. Lee, Macromol. Biosci. 2010, 10, 563. [37] Y. Li, J. Rodrigues, H. Tomas, Chem. Soc. Rev. 2012, 41, 2193. [38] H. T. Cui, J. Shao, Y. Wang, P. B. Zhang, X. S. Chen, Y. Wei, Biomacromolecules 2013, 14, 1904. [39] M. Guvendiren, H. D. Lu, J. A. Burdick, Soft Matter 2012, 8, 260. [40] B. He, J. Zeng, Y. Nie, L. Ji, R. Wang, Y. Li, Y. Wu, L. Li, G. Wang, X. L. Luo, Zhang, Z. R. Z. W. Gu, Macromol. Biosci. 2009, 9, 1169. [41] D. Ma, L. M. Zhang, Biomacromolecules 2011, 12, 3124. [42] J. H. Yu, H. L. Fan, J. Huang, J. H. Chen, Soft Matter 2011, 7, 7386. [43] X. Li, J. Li, J. Biomed. Mater. Res. A 2008, 86, 1055. [44] S. P. Zhao, L. M. Zhang, J. Phys. Chem. B 2006, 110, 12225. [45] K. M. Huh, T. Ooya, W. K. Lee, S. Sasaki, I. C. Kwon, S. Y. Jeong, N. Yui, Macromolecules 2001, 34, 8657. [46] R. Balint, N. J. Cassidy, S. H. Cartmell, Tissue Eng. B 2013, 19, 48. [47] I. S. Kim, J. K. Song, Y. L. Zhang, T. H. Lee, T. H. Cho, Y. M. Song, D. K. Kim, S. J. Kim, S. J. Hwang, Biochim. Biophys. Acta 2006, 1763, 907. [48] I. Titushkin, M. Cho, Biophys. J. 2009, 96, 717. [49] H. Zhuang, W. Wang, R. M. Seldes, A. D. Tahernia, H. Fan, C. T. Brighton, Biochem. Biophys. Res. Commun. 1997, 237, 225. [50] J. Xu, W. Wang, C. C. Clark, C. T. Brighton, Osteoarthritis Cartilage 2009, 17, 397.



Repair and tissue engineering techniques for articular cartilage.

Tissue engineering strategies for fetal myelomeningocele repair in animal models.

Biodegradable Materials for Bone Repair and Tissue Engineering Applications.

Design of 3D scaffolds for tissue engineering testing a tough polylactide-based graft copolymer.

Tensioned synoviocyte neotissues for meniscal tissue engineering: closer to meniscal tissue repair in dogs.

Tissue Engineering to Repair Diaphragmatic Defect in a Rat Model.

Tissue engineering and regenerative repair in wound healing.

Engineering topochemical polymerizations using block copolymer templates.

Engineering the bacterial shapes for enhanced inclusion bodies accumulation.

Tissue Engineering for Pediatric Applications.

3D Printing for Tissue Engineering.

Combining tissue repair and tissue engineering; bioactivating implantable cell-free vascular scaffolds.

The vascularized periosteum flap as novel tissue engineering model for repair of cartilage defects.

New perspectives for articular cartilage repair treatment through tissue engineering: A contemporary review.

Tissue-engineering strategies to repair joint tissue in osteoarthritis: nonviral gene-transfer approaches.

Abstract 94: decellularized porcine stomach extracellular matrix for tissue engineering and repair.

Tissue engineering-based therapeutic strategies for vocal fold repair and regeneration.

or Regeneration.

Tissue engineering is a promising method for the repair of spinal cord injuries (Review).

Tissue-engineering strategies to repair chondral and osteochondral tissue in osteoarthritis: use of mesenchymal stem cells.

Tissue engineering for rotator cuff repair: an evidence-based systematic review.

Immune modulation to improve tissue engineering outcomes for cartilage repair in the osteoarthritic joint.

Tissue engineering in periodontal tissue.

Scaffolds for tissue engineering of cardiac valves.