Introduction
The neural disorder is an essential health challenge which can be caused by accident or disease. Due to the complication of neural regeneration and predomination of neural disease, the evolution of the novel treatment approaches has been performed by researchers [1]. Tissue engineering or regenerative medicine is one of the multidisciplinary fields that could recompense the pain from organ loss. It comprises the synthesis of tissue from biological materials and is substituted with the malfunctioning tissue. Indeed, tissue engineering improves tissue function by combining a suitable supportive matrix, biological materials and cells [2, 3]. Polymer scaffolds have significant roles as supportive matrices with an aim to develop tissue engineering. Hydrogels as an excellent class of biomaterials demonstrate the advantages in many ways for their adjustable biodegradation rate, biocompatibility with host tissue, proper porosity for the carrying of nutrients and wastes, and tunable mechanical properties. In addition, they are networks that can absorb water, and as a result, cells can adhere, proliferate and differentiate into the hydrogels. Also, their structure is similar to the native extracellular matrix so they can be a good carrier for drug/gene delivery systems. Hydrogels in biomedical applications have two categories: (1) natural polymers like chitosan, gelatin and agarose (2) synthetic polymers such as polycaprolactone (PCL), polylacticacide (PLA), polyurethane (PU) [4].
Proliferation and differentiation of cells depend on the surface properties of scaffolds. Recent studies have shown that conductive polymers, as a new generation of materials, have demonstrated a beneficial use by sending cellular signaling. Because of the intrinsic properties of neural tissue, the conductivity is more striking. Signals have transferred in neural tissue through action potential phenomenon, and these conductive polymers can mimic this action. The soft nature of conductive polymers causes a good connection with cells and better biocompatibility in comparison with metals and inorganic materials. Moreover, they can promote the adhesion, migration, proliferation, differentiation, and shape of cells, with or without electrical stimulation. Conductive polymers including polyaniline (PANI), carbon nanotubes (CNTs), polythiophene and polypyrrol (PPY), are appropriate materials in biomedical applications such as tissue engineering, drug delivery, neural prostheses, biosensors and artificial organ; because of the feasible processability and biocompatibility. Although poor solubility, poor degradation, and chronic inflammation are their drawbacks. In order to overcome these disadvantages, aniline oligomers have been utilized which have both biodegradability and solubility with proper electroactivity. Oligo anilines can be consumed by macrophage and filtered with kidney [5]. Because of the light toxicity and brittleness of oligo anilines, they have been grafted with natural biodegradable macromolecules such as chitosan, gelatin, agarose, and alginate; and create bio conductive polymers [6]. Chitosan is a linear bio-based cationic polysaccharide which is driven from deacetylation of its parent polymer chitin; and has received a great attention because of its biocompatibility, biodegradability, low toxicity and the ability to promote cell adhesion. All these properties of aniline oligomers and chitosan make them a great bio conductive scaffold for demanding tissue engineering and drug delivery [7].
Recently, injectable thermosensitive hydrogels that are soluble at room temperature, and become gel after injection into the body, have received a great attention in tissue engineering and biomedical applications. Injectable hydrogels have advantages such as they can be implanted into the body without surgery, which is a noninvasive approach; they can encapsulate cells or drugs in situ; they can take the shape of damaged tissue; and the ease of administration. Thermosensitive gelation is generally used to produce injectable conductive hydrogels, through chemically grafting conductive segments onto the end of thermosensitive materials. Pluronic® F127 is a triblock copolymer poly (ethylene oxide)-b-poly (propylene oxide)-b-poly (ethylene oxide) (PEO-PPO-PEO), which has sol-gel transition above its LCST point, because of hydrophobic interaction in PPO [8].
In our previous work, a novel drug loaded colloidal hydrogel by reacting carboxyl capped oligo-aniline and gelatin was synthesized with tunable properties for neural tissue engineering. Colloidal hydrogels have been used for regulating the neural interface modulus, but they did not have enough conductivity, therefore, these hydrogels can be apt for neural tissue engineering [9]. Also, alginate-aniline tetramer-agarose was synthesized and exhibited applicable interaction with PC12 cells. Agarose, an inert and non-immunological material, was used to prevent the toxicity of crosslinking agent. In spite of its good mechanical features, agarose decreased the ionic conductivity because of the large size molecule which acts as a barrier between conductive units. In addition, Aniline tetramer in alginate structure, augments the conductivity. Because of the ion mobility improvement, ionic conductivity is increased and activation energy is reduced with temperature being incremented [10]. In addition, Colloidal hydrogels have been applied for regulating the neural interface modulus, however, their inadequate conductivity hindered their applications to some extent, so our group has reported a bio-conductive colloidal hydrogel based on agarose-aniline pentamer, which improved the cellular activity and amended the nerve regeneration. For the reason that conductivity, on-demand drug release, mechanical features and self-gelling properties; this hydrogel can be a good option for neural electrodes [5].
Recent cases reported by Gue et al. also support the hypothesis that in comparison with traditional scaffolds, the injectable hydrogels with electrical stimuli not only are feasible to synthesize, but also there is no need to use invasive surgical operations. They developed injectable conductive hydrogels for drug delivery with electro and pH responsibility based on chitosan-poly aniline, and oxidized dextran as a cross-linker. The in vitro and in vivo studies demonstrated inherent antibacterial activity and suitable cytocompatibility [11]. Also, Dong et al. developed a series of injectable conductive self-healing hydrogels based on aniline tetramer-graft-chitosan and dibenzaldehyde-terminated poly ethylene glycol, and exhibited their possibility as cell delivery vehicle for myocardial infarction for cardiac cell therapy [12]. Moreover, they have developed an injectable dopamine based hydrogel by NaIO4 as the oxidizing agent for oxidizing a mixture of chitosan, gelatin and dopamine. This system could deliver dopamine and inflammatory drugs, which they could help treatment of Parkinson’s disease [13]. In a similar case, a series of in situ forming electroactive biodegradable hydrogels based on gelatin-graft-poly aniline and genipin, as a cross-linker, were synthesized. They demonstrated a linear release profile of diclofenac sodium as a drug. The in situ forming convenience administration of these materials in a non-invasive way [14].
Material and methods
Materials
Dimethyl sulfoxide (DMSO), dimethyl formamide (DMF), N-hydroxysuccinimide (NHS) and N, N-di cyclohexyl carbodiimide (DCC), ammonium peroxidisulfate (APS), camphor sulfonic acid and succinic anhydride were purchased from Merck. Medium molecular weight Chitosan, N-Phenyl-p-phenylenediamine, p-Phenylenediamine, Pluronic F127 and β-glycerophosphate was received from Sigma-Aldrich.
Experimental methods
NHS-capped Aniline Pentamer Synthesis
NHS-capped Aniline Pentamer was synthesized as the same way in the references. Carboxyl capped aniline pentamer was synthesized by reacting of N-Phenyl-p-phenylenediamine and succinic anhydride to produce carboxyl capped aniline dimer, and then p-phenylenediamine was added to the mixture to make carboxyl capped aniline pentamer. APS was added dropwise to the mixture. Finally, N-hydroxysuccinimide (NHS) and N,N-di cyclohexyl carbodiimide (DCC) was added to achieve NHS-capped Aniline Pentamer [15].
Synthesis of Chitosan -graft-NHS capped aniline pentamer copolymer
The presence of amine groups in the structure of Chitosan (CS) facilitates the reaction with carboxyl group of aniline pentamer (AP) using carboiimides. 0.5 gr of CS were dissolved in 20 ml camphor sulfonic acid. Also, 0.1 gr NHS-capped-AP were dissolved in 5 ml DMSO. Subsequently, the two solutions were mixed and stirred for 24 h at 50 °C temperature under a nitrogen atmosphere [16].
Synthesis of CS-graft- NHS capped AP / Pluronic F127
The product was mixed with 0.5, 0.25, 0.12 gr pluronic with was dissolved in 5 ml DMSO and stirred at room temperature under nitrogen atmosphere.
Preparation of CS-g-NHS capped AP/Pluronic F127/ β-glycerophosphate hydrogel
0.5 gr β-glycerophosphate was dissolved in 5 ml of deionized water and then added to the aforementioned mixture to reach intended hydrogel.
Characterization and experiments
Fourier transformed infrared spectroscopy (FT-IR)
FT-IR spectra were obtained to study the chemical interactions between the various functional groups within the components present in the CS, NHS-capped AP, CS/NHS-capped AP/Pluronic and final hydrogel, using a Bruker instrument with the KBr disc, made by Germany, working in the range of 4000-600 cm-1 at a resolution of 4 cm – 1 at ambient temperature.
Electro activity and conductivity
UV-visible (UV-Vis) was used to evaluate the AP concentration and transition states of the copolymer with a spectrophotometer (Shimadzu, Kyoto, Japan). To do so, AP has dissolved in DMSO/deionized water (1:1) solvent mixture and the spectra of the undoped AP using camphor sulfonic acid, as a dopant, were assessed at 340 nm. The AP content (C) in the sample was evaluated by dividing of the slope of samples concentration (P) on pure AP slope (P0)(C% = P/P0 * 100) [17].
Micro auto lab type ш apparatus was used for cyclic voltammetry (CV) and electrochemical impedance spectroscopy (EIS). Carbon pasted electrode (CPE) modified with 20% polymer was the working electrode, platinum wire was the counter electrode, Ag/AgCl was the reference electrode and indium tin oxide (ITO) electrode was the working electrode. All of which were put in the 1 M HCl solution and the 50 mVs−1 was applied. Also, EIS was performed at 0.01 Hz. EIS by the neural model which consist of one capacitance and two resistors. With EIS records, Nyquist plot and conductivity properties were obtained [5].
The conductivity of the hydrogel was determined by measuring the resistances using the four-probe method [18]. Samples were prepared in pellet then volt and ampere were applied, the value of conductivity was obtained by the Eq. (1):
σ = 1/R d/S (1)
Where σ is the conductivity, d is the thickness, R is the resistance, and S is the area of the sample, respectively. Ionic conductivity was obtained at the ultimate swollen state in various temperatures using Eq. (2):
σ =σ_0/T expâ¡ã€–(- E_a/RT)〗 (2)
In which, σ0 is the initial conductivity value before the sample being swelled, T is the absolute temperature at which test is performed, Ea is activation energy and R is the molar gas constant.
Conductivity is related to mobility and hole concentration according to (3):
σ=q× μ_p×p (3)
Where q, μp and P are mathe gnitude of electronic charge (1.6 × 10-19 C), mobility and hole concentration [10, 19].
Morphology
The surface morphology, porous structure and cell adhesion of cells to scaffolds the freeze-dried hydrogel were studied using scanning electron microscopy (SEM). Before the examination, the surface of the samples was coated with a gold layer.
Thermal properties
Thermal characteristics of the samples were conducted with differentiation scanning calorimetry (DSC) under a nitrogen atmosphere with a heating rate of 10 °C/min which is swept from 0 to 250 °C and 250 to 0 °C for heating and cooling, respectively.
Furthermore, Thermogravimetric analysis (TGA) of hydrogels were persuaded using a Perkin–Elmer Pyris-1 TGA apparatus, with a heating rate of 10 °C/min from 35 to 900 °C under air and nitrogen atmosphere [20].
Rheological characterization
The gelation time and storage modulus of these hydrogels were measured with MCR 300 Anton Paar Rheometer with cylinder geometry. Temperature sweep test was done in the range of 15°C to 70°C, a frequency 0.1 Hz and strain of 1%. Time sweep test was conducted with a frequency 1 Hz and a strain of 1% in 30 °C. Also, 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.
Contact angle
Owing to the fact that the surface characteristics are connected with the hydrophilicity of materials, it is a vital factor for cell attachment that has to be measured. Contact angles of hydrogels in room temperature were determined using a data physics contact angle system OCAH 200 device, functioning in static mode. A drop of distilled water around 0.1–0.2 μl was introduced onto the surface of the dry hydrogel. Images of the water droplet were documented within 5 s and the contact angles were measured on both sides of the hydrogels and averaged. Digital pictures were analyzed by computer for angle determination.
Porosity
The porosity of the hydrogels was determined by liquid displacement method. The hydrogels were immersed in Specified volume (V1) of hexane in a graduated cylinder for 24h. Hexane was selected because it permeates inside scaffolds without causing a high shrinkage and swelling in comparison with other liquids unlike ethanol, etc.
The entire volume of hexane and the hexane saturated hydrogel was recorded as V2. The hexane-saturated hydrogels were then removed from the cylinder and the residual hexane volume was documented as V3 [21]. The porosity of the scaffold ε was obtained by:
ε= (V_1-V_3)/(V_2-V_3 ) ×100 (4)
Swelling/deswelling behavior of hydrogels
Swelling ratio of hydrogels with similar size and shapes was measured over a day at room temperature in distilled water (pH=7.2). First, Hydrogel samples were freeze-dried and then they were immersed in distilled water and over a specified period of time, they were taken out, and the weight was determined after sweeping excess water on the hydrogel’s surface, until a swelling equilibrium was reached. The swelling ratio was calculated using equation (5), in which SR, Ws, Wd are swelling ratio, swollen weight of sample and dry weight of the sample.
SR=(W_s 〖-W〗_d)/W_s ×100 (5)
The equilibrium water content (EWC) and the rate parameter were intended according to equation (6) and (7).
EWC=(W_e 〖-W〗_d)/W_e ×100 (6)
S_t =S_e (1- e^((-t)â„Ï„)) (7)
Equation (7) is called Voigt equation, and We, St and Se are the weight of swollen hydrogel at equilibrium state, swelling ratio at time t and equilibrium swelling ratio.
In order to assess deswelling behavior, after hydrogels were reached swelling equilibrium, the samples were weighted immediately in specific time periods. The water retention (WR) was measured from Equation (6). Where Wt is the weight of hydrogel at a certain time of deswelling
WR=(W_t 〖-W〗_d)/(W_e 〖-W〗_d )×100
In vitro degradation of hydrogel
The degradation assay was performed with phosphate buffered saline (PBS) at pH 7.4 and 5.5 in 37 °C. Hydrogel bulks were immersed in PBS, the PBS was updated every week and at the interval time point, the hydrogels were taken out and rinsed with DI water to remove excess salinity, and they were then dried in an oven at 60 °C for 48 h and weighed to determine the degradation rate.
weight remaining ratio (WRR)=W_t/W_0 ×100
Where Wt and W0 were the weight of the degraded hydrogels at different time intervals and the weight of the hydrogels which were at swelling equilibrium state, respectively.
In vitro cell compatibility and cytotoxicity of hydrogels
Cell proliferation and biocompatibility were carried out using 3-(4,5-dimethylthiazol-2-yl)-2,5 diphenyltetrazolium bromide (MTT) assay. The samples were sterilized by UV for 20 mins and then ethanol was poured into them, then they immersed in cell culture media for 1 day. After rinsing with PBS, the culture medium was applied to the samples and incubated overnight. the pheochromocytoma (PC12) cells in Dulbecco’s modified eagle medium(DMEM) with 10% fetal bovine serum (FBS) and 105/L penicillin were seeded on the samples, then incubated at 37 °C in 95% moisture and 5% CO2. The MTT test was performed in 1, 3 and 5 days to appraise the biocompatibility of hydrogels.
In order to evaluate the toxicity of hydrogels, the cytotoxicity test was conducted. Samples were incubated at the same condition for 24 hours to reach their secreted liquid. 96-well plate seeded PC12 cells were incubated with various concentrations of medium (1, 5, 10, 20, 50 mg/ml). Subsequently, the cell viability was graphed by using MTT test. For cell attachment assessment, the cells were fixed on the samples which displayed the best compatibility using the glutaraldehyde and alcohol gradients.
Drug release behavior
To evaluate the drug release behavior of the samples in stimulated conditions and passive, 50 mg of dexamethasone was dissolved in 50 ml methanol, and 0.5 g of dry samples was added to methanol solution and stirred for 24h at ambient temperature. Then, the drug loaded hydrogel was washed to remove the weak drug bonds on its surface, and the sample was submerged in PBS (pH 7.4, 37 °C) and at particular interval times the same amount of fresh PBS was replaced. Dexamethasone is a kind of corticosteroid family that demonstrate the immune-suppressant and anti-inflammatory properties. To evaluate the drug release profile, UV-Vis (λ=237 nm) was utilized. In stimulated release, electrical current was used to the hydrogel and after that, the release pattern was determined.