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Synthesis and characterization of semi-conductive nanocomposite based on hydrolyzed collagen and in vitro electrically controlled drug release study
Accepted Manuscript Synthesis and characterization of semi-conductive nanocomposite based on hydrolyzed collagen and in vitro electrically controlled drug release study Ali Pourjavadi, Mohadeseh Doroudian PII:
S0032-3861(15)30069-0
DOI:
10.1016/j.polymer.2015.06.050
Reference:
JPOL 17948
To appear in:
Polymer
Received Date: 7 April 2015 Revised Date:
21 June 2015
Accepted Date: 25 June 2015
Please cite this article as: Pourjavadi A, Doroudian M, Synthesis and characterization of semiconductive nanocomposite based on hydrolyzed collagen and in vitro electrically controlled drug release study, Polymer (2015), doi: 10.1016/j.polymer.2015.06.050. This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
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Synthesis and characterization of semi-conductive nanocomposite based on
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hydrolyzed collagen and in vitro electrically controlled drug release study
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Ali Pourjavadi∗, a, Mohadeseh Doroudian a
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Polymer Research Laboratory, Department Of Chemistry, Sharif University of Technology, Azadi Avenue,P.O.Box11365-9516, Tehran, Iran
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Abstract
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Corresponding author E- mail address: [email protected]
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ACCEPTED MANUSCRIPT In this study, a semi-conductive nanocomposite for electrically controlled drug delivery is
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introduced. Hydrolyzed collagen known as a naturally abundant polypeptide was modified
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with polycaprolactone. This modification changed the mechanical properties of the
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hydrolyzed-collagen. A hydrogel compound was synthesized through radical co-
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polymerization of acrylic acid on the backbone of this biocompatible polymer in the presence
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of a crosslinker. The reaction parameters affecting the water absorbency of the hydrogel were
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optimized using Taguchi method. In situ polymerization of aniline, incorporated conductive
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nanofiber pathways throughout the hydrogel matrix. 1H NMR, TGA, AFM, SEM, FTIR, UV-
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Vis, cyclic voltammetry and conductivity measurements were used for the characterization of
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this system. Moreover, in vitro conductive-stimuli drug release of hydrocortisone as a model
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drug was investigated. MTT assay showed no cytotoxicity for the conductive and non-
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conductive hydrogels. Results suggest that this nanocomposite acts as an appropriate
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externally controlled drug delivery system that can be tailored to match physiological
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processes.
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Keywords: semi-conductive hydrogel; drug delivery; polyaniline.
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1. Introduction
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Recently, conductive scaffolds systems have been developed for both in vitro and in vivo
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biomedical applications. Theses conductive scaffolds such as current-stimuli polymeric
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structures have successfully introduced extra controlling drug delivery systems which provide
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switchable release profile. Nowadays, drug release from conductive polymeric systems by
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applying desired voltage at precise duration time is accessible.
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Inherently conductive polymers have wide applications including in microelectronic industry,
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electrode production, optoelectronic or photovoltaic devices and recently in biotechnological
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systems such as nerve regeneration [1] and drug delivery systems. The major problem of the
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ACCEPTED MANUSCRIPT conductive polymers is their insolubility in common solvents that makes difficulties in film
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forming or production of three dimensional structures. On the other hand the biodegradability
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of the conductive polymers is still undetermined.
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Three dimensional hydrogels demonstrate a distinct resemblance to the physiological
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conditions because of their unique properties such as swelling-deswelling behavior at
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different environmental situations, good mechanical properties and high hydration level.
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Therefore, there are numerous studies that remark application of hydrogels in transdermal,
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ocular and subcutaneous delivery systems. However, because of low electrical conductivity
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of the hydrogels their response to the electrical signal is very slow.
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Considering properties of both conductive polymers and hydrogels, their composite could
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have positively synergic effects. Elastomeric nature and compatibility with physiological
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conditions of hydrogels and high sensitivity of conductive polymers to the electrical signals;
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provide appropriate three dimensional conductive systems for biomedical quests.
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Conductive hydrogels have been used frequently in biotechnological systems including nerve
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regeneration [2-3], actuators [4] and transdermal drug delivery systems. Swelling or
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shrinking of the conductive hydrogels due to electrical signals provides an on-off switching
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system for release of appropriate amount of drug in accurate time and place. For example, the
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release of growth hormones, insulin and estrogen were investigated using conductive
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hydrogels [5]. Therefore, synthesis of conductive hydrogels have been studied by several
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research groups. J. O. You et al. prepared a conductive hydrogel containing Au nanoparticles.
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The conductivity of this hydrogel changes at narrow range of pH which is very important for
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physiological applications [6]. C. Lau et al. used carbon nanotube filler for the fabrication of
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conductive composite with chitosan [7]. C. Hou et al. synthesized a hydrogel with graphene
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and investigated its volume changes in terms of external electrical field [8]. P. M. George et
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ACCEPTED MANUSCRIPT al. used polypyrrole doped biotin for current-controlled drug delivery [9]. S. Niamlang et al.
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examined salicylic acid release via electrical pulse from conductive poly (p-phenylene
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vinylene)/polyacrylamide hydrogel [10]. N. Paradee et al. prepared a conductive drug
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delivery system based on alginate and investigated the effects of crosslinker type and
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electrical field on drug release profile [11]. S. H. Takahashi synthesized a conductive
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hydrogel from polypyrrole and polyacrylic acid. This composite showed zero-order release
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profile of safranin through electrical signals [12]. S. S. Indermun et al. examined release of
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indomethacin from conductive PVA-vinyl imidazole hydrogel [13]. However, it remains a
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challenge to consider all structural factors of the hydrogels [14]. Therefore, target application
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of hydrogels must be cleared in order to fast introduction of them into clinical formulation.
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However, there are some studies about in vivo drug release from conductive hydrogels. For
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example, methyl methacrylate-CNT hydrogel was used for in vivo release of
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[15].
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In this study, we synthesized an electro-responsive hydrogel containing biodegradable and
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available
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physiological systems of body, researchers have used collagen at tissue engineering
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frequently [16-17]. We found that the mechanical properties and film forming of hydrolyzed
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collagen improve via polycaprolactone modification. It has been well documented that
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polyaniline possesses good conductivity at emeraldine salt state as well as environment
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stability [18]. Studies have shown that polyaniline has no destructive effects on tissue cells
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but there are concerns about chronic inflammation due to its non-biodegradability. K.
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Sengothi et al. showed that polyaniline have no bad or carcinogen effect on Sprague-Dawley
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rat cells after two years of implantation [19]. There are various methods for incorporation of
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conductive polymers into an insulator material. One of the most studied methods is in situ
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polymerization of monomer on the pores of the hydrogels [20-21]. Through polymerization
collagen
polypeptide.
Considering
good
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with
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ACCEPTED MANUSCRIPT of aniline at emeraldine oxidation form, we performed hydrolyzed collagen-based conductive
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hydrogel. Here, the optimized hydrogel synthesis with Taguchi method was carried out and
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its potential as an extra controlled drug release systems examined with hydrocortisone
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hydrochloride as a model drug.
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2. Experimental
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2.1. Materials
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Hydrolyzed collagen type II (from Parvar Novin-E Tehran, Iran), acrylic acid (distilled under
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reduced pressure), N,N´-methylene bisacrylamide (MBA from Loba chemie, India),
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ammonium persulfate (APS from Fulka), aniline (distilled under reduced pressure), stannous
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octoate (Sn(Oct)2), ε-caprolactone (from Merck) and Hydrocortisone (from Jaber Ebne
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Hayyan pharmaceutical Co., Iran) were used without further purification.
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2.2. Characterization
The vibrational spectra were measured using FTIR spectrometer (MBB Bommem, MB-100)
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in the range of 400 to 4000 cm-1. Samples were compressed into a KBr disk for FTIR
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measurements. Thermal characteristics of the samples were studied using thermal gravimetric
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analysis (TGA). TGA was carried out under a nitrogen atmosphere with a heating rate of 10
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ºC/min and the scan range from 30 to 600 ºC. Veeco’s Autoprobe (CP-research) atomic force
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microscope in non-contact mode and field-emission scanning electron microscope (FE-SEM,
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S-4160) were used for morphological studies. Voltammetry measurement was carried out
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using an electrochemical cell made up of gold wire working electrode, platinum wire counter
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electrode and Ag/AgCl reference electrode immersed in phosphate buffer solution with
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pH=7. UV-Vis spectra were measured from 200 to 800 nm in DMSO using PerkinElmer
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(Lambda 25) UV-Vis spectrometer. A Bruker (DRX-500 Avanes) NMR was used to record
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ACCEPTED MANUSCRIPT the 1H NMR spectra in DMSO. A SANTAM universal testing machine (STM-50 series) was
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used for mechanical measurements of the hydrogels under tensile loading force 1 kN.
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2.3. Modification of the hydrolyzed collagen with polycaprolactone (C-PCL)
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3 g of hydrolyzed collagen was dissolved in 7 mL of distilled water and separated from
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insoluble phosphate salts. 3 mL of ε-caprolactone (0.027 mol) was added into the
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collagen solution and the mixture stirred in 80 ºC for one hour. Then, 200 µL of Sn(Oct)2
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was added to start ring opening polymerization of ε-caprolactone. The reaction was
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allowed to continue at 110 ºC for 12 hours and then, the unreacted monomers and
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homopolymers of ε-caprolactone were separated by liquid extraction with toluene at 60
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ºC for 24 hours. The modified hydrolyzed collagen (C-PCL) in aqueous phase was dried
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at 50 ºC for three days.
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2.4. Experimental design
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Repeatability is an important characteristic that must be considered in order to achieve
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appropriate target drug delivery systems. It is understood that swelling of the hydrogels
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influences both uptake and release of the drugs and thus, all of structural parameters
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which are known to be important must take into account. Taguchi model as a powerful
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statistic method is able to provide an approach to optimize these structural parameters
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[22].
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Accordingly, an orthogonal array for three parameters (initiator, crosslinker and
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monomer) in three levels was designed as presented in Table 1 and Table 2.
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Table 1
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Table 2
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and contribution of each factor.
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2.5. Synthesis of conductive hydrogel (CH)
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In general, 1.00 g of modified collagen was dissolved in 40.0 mL of distilled water and
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stirred at 80 ºC. After 2 minutes certain amount of APS (0.03-0.1) was added for macro
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radical initiator formation. After 2 minutes acrylic acid (3-5 mL) and several minutes later
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MBA (0.05-0.2) were added respectively and finally, hydrogels formed during 15-60
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minutes. The hydrogels were placed in distilled water for 2 days to remove unreacted
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monomers and then dried at 50 ºC for 3 days. The conductive polyaniline insertion was
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done through in situ polymerization of aniline. In a typical procedure the hydrogels were
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swollen in HCl solution (1 M) containing aniline (1 mL per 5 g of dry hydrogel). After
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one day the hydrogels were removed and washed with water then, placed in 1.1 M APS
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solution. In situ polymerization of soaked aniline and color change from yellow to green
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was performed gradually in the hydrogel matrix during one hour.
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Swelling measurements were done by the gravimetric method. The Equilibrium Swelling
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(ES) capacity was calculated at room temperature using the following formula (Eq. 1):
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ES (g/g) = (W2 – W1)/W1
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where W1 and W2 are the weights of dry and swollen gel, respectively.
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2.6. Drug loading (DL) and release profiles
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100.0 mg of hydrocortisone was dissolved in saline phosphate buffer (PBS) at pH=7.
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5.00 g of dry conductive hydrogel (CH) (1.5×0.5×0.2 cm pieces) was placed in the drug
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solution for one day. The swollen hydrogels were raised and placed in the fresh buffer in
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order to remove the weak adsorbent drugs in the surface of the hydrogels.
(Eq. 1)
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ACCEPTED MANUSCRIPT For determination of actual amount of the drug, the dry hydrogel was cut into tiny pieces
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and placed in 20 mL of phosphate buffer with pH=7 for 3 days. The concentration of drug
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was determined using UV-Vis absorption peak at 245 nm. Eq. 2 was used for calculation
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of actual amount of the drug:
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DL=
(Eq. 2)
× 100
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where A is the actual quantity of hydrocortisone which determined with UV-Vis
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spectroscopy and A0 is the theoretical amount of the drug.
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For drug release investigation, PBS (NaCl: 0.8 g, KCl: 0.02 g, Na2HPO4: 0.178 g,
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KH2PO4: 0.027 g) with pH=5.5 was chosen for simulation of skin conditions. Five
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different release profiles were investigated. At the first release experiment, samples were
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placed in 30.0 mL of PBS at 37 ºC and 3.00 mL of buffer solution were withdrawn at
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selected times (5, 8, 18, 21, 31, 34, 44, 47, 57 and 60 minutes after incubation) for drug
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concentration analysis and returned to the mother solution (Release 1).
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At the second release experiment, samples were placed in 30.0 mL of PBS at 37 ºC and
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potential difference (3 V) applied to the hydrogel by using two electrodes for 3 minutes in
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the selected times (5, 18, 31, 44, 57 minutes after incubation). 1 mL of PBS was
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withdrawn with the same time intervals as Release 1 and diluted to 10 mL for
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determination of the drug concentration with UV-Vis analysis (Release 2).
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The third release experiment was performed quite similar to the Release 2 profile except
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that potential difference between two electrodes was 1.5 V (Release 3).
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The fourth and fifth experiments were done similar to the Release 2 and 3 respectively,
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expect that duration time of the applied potential was 1 minute (Release 4 and 5). UV-
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ACCEPTED MANUSCRIPT Vis analyses were done at different selected times (5, 6, 18, 19, 31, 32, 44, 45, 57 and 58
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minutes after incubation). All reported data are the average of three experiments.
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2.7. Cell viability
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To evaluate cytotoxicity of the conductive hydrogel, MTT colorimetric assay based on
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general methods was used. For extract of products, sterilized samples (8 cm3) using UV
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irradiation were placed in individual wells and added cell culture media (EMEM (EBSS)
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+ 1% NEAA + Bovine insulin-optional +10% FBS) in each well and incubated at 37 ºC
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for one week. This media were removed for MTT assay. A431 (human skin cell
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carcinoma) were seeded at a density of 1×104 cells/mL on a 96-well tissue culture
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polystyrene plate and incubated at 37 ºC for 24 hours. The extract of the products (at
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various concentration) were added to the wells and cells cultured for more 24 hours. Then
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the culture medium was removed and 100 µL of MTT solution diluted in PBS (0.5
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mg/mL) added to the wells. Plate incubated at 37 ºC for 4 h. after removal of the medium,
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isopropanol solution was added to each well in order to complete dissolving of the
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produced blue crystals. The absorbance was measured at 545 nm using spectrometric
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microplate reader. The cell viability was calculated by normalization of optical density of
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the samples to the control. . The wells with more alive cells show higher optical density
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(OD) therefor, the cell viability could be calculated using equation (3) and (4).
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Toxicity % = (1 – (mean OD of sample / mean OD of control) × 100
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Viability % = 100 – Toxicity % Eq. (4)
Eq. (3)
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3. Result and Discussion
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3.1. Synthesis of the conductive hydrogel
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In this study, a conductive hydrogel based on hydrolyzed collagen was synthesized.
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Hydrolyzed collagen was modified with polycaprolactone which is frequently used at
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biological systems like tissue engineering and drug delivery systems. Besides the
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biodegradability, hydrophobic characteristic of polycaprolactone may enhances the
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interaction with non-ionic drugs. P. Dubois et al. modified polysaccharides with
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polycaprolactone and observed the improvement of mechanical and biodegradation
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ACCEPTED MANUSCRIPT properties [65]. Good film forming and mechanical properties of polyesters [14] were
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another reasons that lead us to modify hydrolyzed collagen with polycaprolactone.
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Therefore, ring opening polymerization of ε-caprolactone on hydrolyzed collagen using
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Sn(Oct)2 as the catalyst was performed. As shown in Scheme 1, hydrolyzed collagen
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powder was transformed to a viscose yellow liquid after grafting of polycaprolactone.
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The loading of PCL on the polymer backbone is approximately 50-60 % determined with
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weight difference between the starting material and the product. Co-polymerization of
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acrylic acid onto backbone of the modified collagen was carried out in the presence of
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MBA as a crosslinking agent.
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Scheme 1
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The water absorbency results of the nine hydrogels proposed by Taguchi’s model are
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listed in Table 3. Two batch of each sample were synthesized and presented as ES1 and
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ES2. Swelling measurements were repeated twice.
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Table 3
The optimum conditions analyzed with Minitab 16 software were obtained (4 mL of
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acrylic acid, 0.05 g of MBA and 0.1 g of APS). However, poor mechanical properties of
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this hydrogel was not suitable for our quests. Using the Taguchi’s model suggestion, we
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examined other conditions which possess near swilling ratio to the optimized value.
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Finally, a hydrogel (4 mL AA, 0.05 g APS, 0.1 g MBA) with both good water absorbency
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and mechanical properties was synthesized for our following studies.
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In order to achieve conductive hydrogel (CH), aniline monomers diffused through
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hydrogel pores and nanofibres of polyaniline formed due to in situ polymerization in the
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APS/HCl solution. In order to prevent clogging of the pores in the early stages of
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polymerization, the surface of hydrogels were washed thoroughly before incubation in the
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ACCEPTED MANUSCRIPT initiator solution. Appropriate swelling of the hydrogel make homogenous polymerization
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of aniline which is obvious from the shown cross section photograph (Scheme 1).
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3.2. Chemical characterization
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FT-IR and 1H NMR analyses were used for characterization of the synthesized C-PCL. In
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FT-IR spectrum (Fig. 1S), the peaks at 1120 and 1730 cm-1 correspond to the C-O and
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C=O stretching bands of polyester respectively and verify the presence of
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polycaprolactone in the C-PCL. Moreover, the intensity of alkyl C-H stretching band at
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2950 cm-1 has increased after modification. A representative 1H NMR spectra of
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hydrolyzed collagen and C-PCL are shown in Fig. 2S. The peaks were assigned to the
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corresponding hydrogen atoms of the polycaprolactone in Fig. 2Sb. The intensity of the
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hydrolyzed collagen hydrogen peaks at C-PCL has decreased in the presence of the
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polycaprolactone.
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Polyaniline incorporation into the hydrogel matrix was confirmed with comparision
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between FTIR spectra of polyaniline, non-conductive and conductive hydrogels (Fig. 3S).
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3.3. Thermal and mechanical characterization
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The TGA thermograms of C-PCL hydrogel (sample 1) and conductive hydrogel (CH)
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(sample 2) are plotted in Fig. 1. Sample 2 has the same degradation process as sample 1,
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which is attributed to the degradation of hydrolyzed collagen and polycaprolactone
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moiety. The slight difference in the thermal stability of the two samples in the range of
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350 to 600 ºC is due to the presence of polyaniline segment in the conductive hydrogel
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(sample 2).
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Figure 1
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ACCEPTED MANUSCRIPT The mechanical strength of the hydrogels are dependent on the amount of absorbent
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water. The more water absorbance, the lower the mechanical strength. Incorporation of
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polycaprolactone and polyaniline decreases the water absorbency and consequently
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improves the mechanical strength of the hydrogel. The Young’s modulus of proposed
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hydrogels for tissue engineering applications is 600 to 1600 KPa [23]. Both hydrolyzed
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collagen (sample A) and modified collagen (sample B) hydrogels were examined for
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mechanical tensile analysis (Fig. 2). The Young’s modulus of sample B had a huge
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increase compared to the sample A (from 93 to 340 MPa) with higher break elongation.
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This suggests that polyester moiety could effectively improves the mechanical strength of
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the hydrogel at wet state which is very important for biomedical application. This is
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probably related to the good mechanical property and low water absorbency of the
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polycaprolactone segment.
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3.4. Morphological studies
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In order to correlate the current-stimuli drug release behavior of the hydrogel to its
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physical characteristics changes, SEM and AFM experiments were carried out before and
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after applying potential.
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As illustrated in Fig. 3, AFM images exhibit morphological changes of the hydrogel after
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applying potential in 10 minutes. AFM images of the CH before (Fig. 3a) and after (Fig.
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3b) electrical field applying demonstrate that the surface roughness changed and more
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smoothness along with pore size expansion performed as illustrated in Fig. 3c. Through
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current flow, destructions take place in the surface of the hydrogel and size expansion of
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pores causes the drug release.
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ACCEPTED MANUSCRIPT Moreover, SEM images of CH films before (Fig. 4a) and after (Fig. 4b) applying
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potential demonstrate the effect of electrical field on the surface of the hydrogel. Surface
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porosity has changed through electrical field due to the polyaniline reduction and
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hydrogel backbone destruction. Hydrogel pore size before the potential applying is
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around of 10-50 nm (Fig. 4S). Formation of large pore size (approximately 50-100 nm) in
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the presence of electrical field have been shown in Fig. 4b.
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3.5. Electrical characteristics of conductive hydrogel
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Cyclic voltammetry graph of the polyaniline segment in phosphate buffer solution is
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presented in Fig. 5, which working electrode was in contact with surface of the CH. It is
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clear that cyclic voltammetry behavior of polyaniline has been preserved at the hydrogel
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matrix. Two oxidation peaks at 0.25 and 0.65 V are related to the leucoemeraldine-
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emeraldine and emeraldine-pernigraniline equilibrium states of the polyaniline
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respectively. The reduction peak is attributed to the protonated emeraldine and
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leucoemeraldine equilibrium states of the polyaniline.
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Figure 5
The surface resistivity of CH films were measured in both sides before and after
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performing 3 V potential in 10 minutes. Experiments have provided evidences that semi-
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conductive films with surface resistivity of 1.5×10-4 to 3×10-5 S.cm2 are acceptable for
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biomedical applications [2]. As illustrated in Table 4, the surface resistivity of the films
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increased after potential applying. This result could be related to the proton exchange
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through polyaniline chains due to current flow which increases the conductivity of the
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film. In the other related studies, volume resistivity have been measured instead of
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surface resistivity usually. For example, the volume conductivity of 0.3-5 mS/cm for
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collagen/CNT [17] and 1.1×10-3 mS/cm [3] or 10-6 [24] for hydrogel/polyaniline
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composites were reported. Table 4
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For UV-Vis measurement, CH was cut into small portions and then immersed in DMSO.
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The color of solvent changed due to polyaniline dissolving after 30 minutes. The UV-Vis
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spectrum (Fig. 6 solid line) demonstrates that polyaniline has preserved its optoelectronic
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characteristics in the hydrogel. Peaks at 320 and 600 nm are assigned to the π-π*
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transition of the benzene ring and the benzenoid to quinoid excitonic transition
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respectively. UV-Vis spectrum of the polyaniline alone is shown with dash line as
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control.
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Figure 6
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3.6. Swelling behavior of the hydrogels in various conditions
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The swelling behavior of the hydrogels (conductive and non-conductive form) in the PBS
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were measured through 3 hours (Fig. 7). The swelling degree of the hydrogels is
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dependent on the incorporation of conductive polyaniline or applying the electrical field.
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The swelling ratio of B, D and E samples were determined by immersing the pre-weight
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dry hydrogels in buffer solution and applying electric potential every 5 minutes.
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It is clear that non-conductive hydrogels swell rapidly than conductive hydrogels because
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of hydrophobic characteristic of polyaniline. Potential applying had no notable influence
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on the swelling behavior of the non-conductive hydrogels (A and B). On the contrary, the
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swelling behavior of the CHs were dependent on the current flow as seen in Fig. 7. As a
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result of current flow, polyaniline reduction and expansion of pores takes place causing
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the water uptake increasing. By applying higher potential, the swelling ratio of the CHs
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have increased. But why current flow have no influence on swelling of the non-
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ACCEPTED MANUSCRIPT conductive hydrogels? There are two possible reasons: Firstly, non-conductive hydrogels
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have their maximum swelling degree and secondly, in the CHs, current flow makes
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changes in the oxidation states of polyaniline which is the reason of pore size expansion
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and increasing of swelling ratio. But in the non-conductive hydrogel matrix, ions have the
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role of current transition which have no effect on the hydrogel morphology. Figure 7
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3.7. Drug release study
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Drug entrapment efficiency is one of the most important considerations in drug delivery
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systems. As illustrated in Fig. 8, the drug entrapment percent was calculated for
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conductive and non-conductive hydrogels seventh times. The average of these values are
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59.3 and 74.45 % for conductive and non-conductive hydrogels respectively. Lower
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swelling ratio of the CH compared to the non-conductive hydrogel due to hydrophobic
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characteristics of polyaniline causes the reduction of drug entrapment efficiency.
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Figure 8
For drug release studies, the non-ionic hydrocortisone was chosen as a model drug and a
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simple two electrode method was carried out for applying electric potential in the saline
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buffer solution (pH=5.5). The concentration of the released drug was followed by UV-Vis
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spectroscopy. Drug loading into the CH was carried out by swelling of the hydrogel in
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the drug solution with appropriate concentration (100 mg of drug per 5 g of dry hydrogel)
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and subsequently immersed in the fresh buffer solution to detach the adsorbent drugs on
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the surface of the hydrogel. The drug release was performed at 37 ºC in buffer solution
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and potential applied through two electrode (2 cm height and 2 mm diameter). For
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acceptable comparison, the hydrogel films were cut into the identical rectangular shape
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ACCEPTED MANUSCRIPT (1.5×0.5 cm approximately 0.25 g).
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mechanism, we examined five release methods through one hour.
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In the first method (Release 1 Fig. 9a), there was no current flow, and the drug release
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was performed via swelling of the hydrogel in buffer solution leading to exit of the drug
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from the matrix pores. There was slight increase of the drug release percent at early
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stages, but its maximum value reached to 40 % after 1 day incubation.
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In the next release method, electrical potential with various strength and duration time
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were applied. At the second experiment (Release 2, Fig. 9b), 3 V potential was applied
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between two electrodes for three minutes in each interval times. However, at the first
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electric pulse complete destruction of the hydrogel took place and 90% of the drug
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released from the CH. Potential of 1.5 V for 3 minutes duration times (Release 3 Fig. 9c)
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made more controllable drug release approach. The whole drug removed from the
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hydrogel after one hour, in the other words, 15 % of the drug released in each step
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approximately. At Release 4 (Fig. 9d), through applying potential difference of 1.5 V for
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1 minute 60 % of the drug released after 1 hour; i.e. 7 % in each step. Due to potential
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difference of 1.5 V between two electrodes (Release 5, Fig. 9e), 3 % of the drug released
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per one minutes. (27 % after one hour).
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For accurate investigation of the drug release
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Figure 9
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If drug interacts strongly with hydrogel scaffold through electrostatic forces, its release
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without any driving force will perform slowly. The proposed mechanism of drug release
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in current stimuli drug delivery systems, like our case study, is direct expansion of the
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pore size and increasing the free volume in the hydrogel matrix. It happens when the
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oxidation state of the conductive polymer changes (such as polyaniline) and consequently
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leads the drugs to the electrodes with different charges or to the hydrogel pores. Pore size
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ACCEPTED MANUSCRIPT expansion of the hydrogels may takes place due to the migration of water or charged ions
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from the hydrogel to the electrodes [25]. The influence of electrical field strength on the
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drug release is dependent on the mechanism of drug release as well as interaction of drug
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with hydrogel scaffold.
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Hydrocortisone, which known as a non-ionic drug can interacts with hydrogel scaffold
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via two routes: hydrogen bonding with polar functional groups and van der Waals
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interaction with PCL. Experiments demonstrated that, electrical signals were able to
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speed up the release of hydrocortisone. Upon using this technique, drug release can be
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regulated by strength and duration of applied electrical field. Both morphological studies
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(AFM and SEM) and drug release experiments confirmed that meanwhile the expansion
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of pore size of the hydrogel due to electrical signals, the drug has penetrated through the
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surface and finally, released.
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3.8. Cell viability
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Cell viability was measured using MTT assay and results are represented in Fig. 10. The
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cell viability numbers over 80 % for the hydrogels at non-conductive and conductive
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form (Sample 1 and Sample 2 respectively) indicated no toxicity of the hydrogels. The
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cell viability results based on the in vitro MTT method are shown in two different
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concentrations (Sample 1' and 2' are 10 times diluted than Sample 1 and 2). Therefore,
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polycaprolactone modification in Sample 1 and polyaniline incorporation in Sample 2
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have no harmful effects on the hydrolyzed collagen-based hydrogels, which is a primary
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concern for their in vivo drug release study.
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4. Conclusion
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ACCEPTED MANUSCRIPT In this study, a semi-conductive hydrogel film was investigated as a current stimuli drug
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delivery system. Hydrolyzed collagen as a biodegradable scaffold was modified with
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polycaprolactone and then, through co-polymerization of acrylic acid and subsequently in
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situ polymerization of polyaniline the conductive hydrogel was produced. The various
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characterization methods and drug release experiments well illustrated that this
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nanocomposite hydrogel could be used as a promising current stimuli drug delivery
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system.
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Figure captions:
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Figure 1: TGA curves of a) C-PCL based hydrogel b) CH
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Figure 2: stress/elongation curves of a) hydrolyzed collagen based hydrogel b) C-PCL
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based hydrogel with 20% humidity at room temperature
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Figure 3: AFM images of CH films a) before and b) after applying potential of 3 V for
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10 minutes c) a representative for morphological change of the hydrogel through
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application of 3 V potential for 10 minutes
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Figure 4: SEM images of the CH films a) before and b) after application of 3 V potential
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for 10 minutes
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Figure 5: Cyclic voltammograms of conductive hydrogel in aqueous buffer solutions
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pH=7. Scan rate: 10 mV/s.
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Figure 6: UV-Vis spectrum of the CH immersed in DMSO (polyaniline alone as a
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control represented with dashed line)
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Figure 7: Swelling ratio in terms of time for the non-conductive (A and B) and
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conductive (C-E) hydrogels with different condition: No electrical signals (A and C);
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applying 3 V potential (B and E); applying 1.5 V potential (D)
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Figure 8: Drug entrapment values of a) non-conductive and b) conductive forms of the
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hydrogel
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Figure 9: drug release study in phosphate buffer pH=5.5, 37 ºC, a) No electrical
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potential. In the electric potential of b) 3 V for 3 min c) 1.5 V for 3 min d) 3 V for 1 min
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e) 1 V for 1 min.
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conductive hydrogel (Sample 2)
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[13] Indermuna S, Choonara YE, Kumara P, du Toita LC, Modib G, Luttgec R, Pillaya
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polyester/polysaccharide blends. Macromol Symp 2003, 198; 233-243.
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[16] Denning D, Abu-Rub MT, Zeugolis DI, Habelitz S, Pandit A, Fertala A, Rodriguez
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hydrogels. Acta Biomaterialia 2012, 8; 3073-3079.
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the electrical conductivity of fibroblast-seeded collagen hydrogels. Acta Biomaterialia
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[18] Stejskal J. Polyaniline. Preparation of a conducting polymer. Pure Appl Chem 2002,
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polyaniline polymers in tissue: Biomaterial surface interactions. In: AIChE Annual
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[20] de Jesus MC, Weiss RA. Synthesis of Conductive Nanocomposites by Selective In
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Situ Polymerization of Pyrrole within the Lamellar Microdomains of a Block Copolymer.
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Macromolecules 1998, 31; 2230-2235.
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[21] Luong ND, Korhonen JT, Soininen AJ, Ruokolainen J, Johansson LS, Seppala J.
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Processable polyaniline suspensions through in situ polymerization onto nanocellulose.
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Eur Polym J 2013, 49; 335-344.
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[22] Bardajee GR, Pourjavadi A, Soleyman R. Irradiation synthesis of biopolymer-based
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superabsorbent hydrogel: optimization using the Taguchi method and investigation of its
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swelling behavior. Adv Polym Tech 2009, 28; 131–140.
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[23] You JO, Rafat M, Ye GC, Auguste DT. Nanoengineering the heart: conductive
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scaffolds enhance connexin 43 expression. Nano Lett 2011, 11; 3643–3648.
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[24] Zhao W, Glavas L, Odelius K, Edlund U, Albertsson A. A robust pathway to
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electrically conductive hemicellulose hydrogels with high and controllable swelling
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behavior. Polymer 2014, 55; 2967-2976.
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ACCEPTED MANUSCRIPT [25] Qiu Y, Park K. Environment-sensitive hydrogels for drug delivery. Adv Drug
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Deliver Rev 2001, 53; 321-339.
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Table 1: Experimental control factor and their levels Terial A B C 1 1 1 1 2 1 2 2 3 1 3 3 4 2 1 2 5 2 2 3 6 2 3 1 7 3 1 3 8 3 2 1 9 3 3 2
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Table 2: Experimental layouts of an L9 orthogonal array according to Taguchi’s suggestion Control factor Level 1 Level 2 Level 3 Acrylic Acid (mL) 3 4 5 APS (g) 0.03 0.05 0.1 MBA (g) 0.05 0.1 0.2 500
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3.89 4.05 4.44
11.34 5.57 7.046
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ES2(g/g)
22.38 18.53 17.67
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Table 3: Water absorbency of nine trials of hydrogels Trial 1 2 3 4 5 6 6.92 3.54 13.34 4.77 33.38 ES1(g/g) 24.51 5.28 6.57 6.38
27.26 29.4 22.52
7 5.37 5.18 5.9 6.09
8 12.23 12.38 10.83 10.76
Table 4: The surface resistivity of films (S.cm2): 1) Before applying potential 2) After applying potential of 3V for 10 minutes Surface 1 Surface 2 18×10-6 8×10-6 1 -3 1.82×10 1.456×10-3 2 503 504 505
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9 5.29 5.43 5.18 5.26
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Figure 1: TGA curves of a) C-PCL based hydrogel b) CH
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Figure 2: stress/elongation curves of a) hydrolyzed collagen based hydrogel b) C-PCL based
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hydrogel with 20% humidity at room temperature
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Figure 3: AFM images of CH films a) before and b) after applying potential of 3 V for 10 minutes c) a representative for morphological change of the hydrogel through application of
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3 V potential for 10 minutes
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Figure 4: SEM images of the CH films a) before and b) after application of 3 V potential for
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10 minutes
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Figure 5: Cyclic voltammograms of conductive hydrogel in aqueous buffer solutions pH=7.
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Scan rate: 10 mV/s.
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Figure 6: UV-Vis spectrum of the CH immersed in DMSO (polyaniline alone as a control
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represented with dashed line)
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Figure 7: Swelling ratio in terms of time for the non-conductive (A and B) and conductive (C-E) hydrogels with different condition: No electrical signals (A and C); applying 3 V
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potential (B and E); applying 1.5 V potential (D)
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Figure 8: Drug entrapment values of a) non-conductive and b) conductive forms of the
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hydrogel
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Figure 9: drug release study in phosphate buffer pH=5.5, 37 ºC, a) No electrical potential. In
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the electric potential of b) 3 V for 3 min c) 1.5 V for 3 min d) 3 V for 1 min e) 1 V for 1 min.
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Figure 10: The cell viability of the polycaprolactone modified hydrogel (Sample 1) and
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conductive hydrogel (Sample 2)
ACCEPTED MANUSCRIPT 1) Hydrolyzed collagen was modified with polycaprolactone. 2) The mechanical properties of composites based on modified collagen improved greatly. 3) A conductive hydrogel was synthesized through in situ polymerization of aniline.
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and applied for in vitro hydrocortisone release.
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4) A current-stimuli drug delivery system proposed by Taguchi method was designed
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Supplemental Material
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Supplemental Figure 1: FTIR spectra of a) hydrolyzed collagen and b) modified collagen
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Supplemental Figure 2: H NMR spectra of a) hydrolyzed collagen and b) C-PCL in DMSO Supplemental Figure 1: FTIR spectra of a) hydrolyzed collagen and b) modified collagen
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Supplemental Figure 3: FTIR spectra of a) Polyaniline b) non-conductive hydrogel and c)
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conductive hydrogel.
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Supplemental Figure 4: SEM image of the hydrogel before potential applying
S0032-3861(15)30069-0
DOI:
10.1016/j.polymer.2015.06.050
Reference:
JPOL 17948
To appear in:
Polymer
Received Date: 7 April 2015 Revised Date:
21 June 2015
Accepted Date: 25 June 2015
Please cite this article as: Pourjavadi A, Doroudian M, Synthesis and characterization of semiconductive nanocomposite based on hydrolyzed collagen and in vitro electrically controlled drug release study, Polymer (2015), doi: 10.1016/j.polymer.2015.06.050. This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.
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ACCEPTED MANUSCRIPT 1
Synthesis and characterization of semi-conductive nanocomposite based on
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hydrolyzed collagen and in vitro electrically controlled drug release study
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Ali Pourjavadi∗, a, Mohadeseh Doroudian a
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Polymer Research Laboratory, Department Of Chemistry, Sharif University of Technology, Azadi Avenue,P.O.Box11365-9516, Tehran, Iran
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Abstract
∗
Corresponding author E- mail address: [email protected]
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ACCEPTED MANUSCRIPT In this study, a semi-conductive nanocomposite for electrically controlled drug delivery is
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introduced. Hydrolyzed collagen known as a naturally abundant polypeptide was modified
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with polycaprolactone. This modification changed the mechanical properties of the
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hydrolyzed-collagen. A hydrogel compound was synthesized through radical co-
24
polymerization of acrylic acid on the backbone of this biocompatible polymer in the presence
25
of a crosslinker. The reaction parameters affecting the water absorbency of the hydrogel were
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optimized using Taguchi method. In situ polymerization of aniline, incorporated conductive
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nanofiber pathways throughout the hydrogel matrix. 1H NMR, TGA, AFM, SEM, FTIR, UV-
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Vis, cyclic voltammetry and conductivity measurements were used for the characterization of
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this system. Moreover, in vitro conductive-stimuli drug release of hydrocortisone as a model
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drug was investigated. MTT assay showed no cytotoxicity for the conductive and non-
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conductive hydrogels. Results suggest that this nanocomposite acts as an appropriate
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externally controlled drug delivery system that can be tailored to match physiological
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processes.
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Keywords: semi-conductive hydrogel; drug delivery; polyaniline.
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1. Introduction
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Recently, conductive scaffolds systems have been developed for both in vitro and in vivo
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biomedical applications. Theses conductive scaffolds such as current-stimuli polymeric
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structures have successfully introduced extra controlling drug delivery systems which provide
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switchable release profile. Nowadays, drug release from conductive polymeric systems by
40
applying desired voltage at precise duration time is accessible.
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Inherently conductive polymers have wide applications including in microelectronic industry,
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electrode production, optoelectronic or photovoltaic devices and recently in biotechnological
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systems such as nerve regeneration [1] and drug delivery systems. The major problem of the
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ACCEPTED MANUSCRIPT conductive polymers is their insolubility in common solvents that makes difficulties in film
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forming or production of three dimensional structures. On the other hand the biodegradability
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of the conductive polymers is still undetermined.
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Three dimensional hydrogels demonstrate a distinct resemblance to the physiological
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conditions because of their unique properties such as swelling-deswelling behavior at
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different environmental situations, good mechanical properties and high hydration level.
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Therefore, there are numerous studies that remark application of hydrogels in transdermal,
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ocular and subcutaneous delivery systems. However, because of low electrical conductivity
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of the hydrogels their response to the electrical signal is very slow.
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Considering properties of both conductive polymers and hydrogels, their composite could
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have positively synergic effects. Elastomeric nature and compatibility with physiological
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conditions of hydrogels and high sensitivity of conductive polymers to the electrical signals;
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provide appropriate three dimensional conductive systems for biomedical quests.
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Conductive hydrogels have been used frequently in biotechnological systems including nerve
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regeneration [2-3], actuators [4] and transdermal drug delivery systems. Swelling or
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shrinking of the conductive hydrogels due to electrical signals provides an on-off switching
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system for release of appropriate amount of drug in accurate time and place. For example, the
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release of growth hormones, insulin and estrogen were investigated using conductive
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hydrogels [5]. Therefore, synthesis of conductive hydrogels have been studied by several
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research groups. J. O. You et al. prepared a conductive hydrogel containing Au nanoparticles.
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The conductivity of this hydrogel changes at narrow range of pH which is very important for
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physiological applications [6]. C. Lau et al. used carbon nanotube filler for the fabrication of
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conductive composite with chitosan [7]. C. Hou et al. synthesized a hydrogel with graphene
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and investigated its volume changes in terms of external electrical field [8]. P. M. George et
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ACCEPTED MANUSCRIPT al. used polypyrrole doped biotin for current-controlled drug delivery [9]. S. Niamlang et al.
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examined salicylic acid release via electrical pulse from conductive poly (p-phenylene
70
vinylene)/polyacrylamide hydrogel [10]. N. Paradee et al. prepared a conductive drug
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delivery system based on alginate and investigated the effects of crosslinker type and
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electrical field on drug release profile [11]. S. H. Takahashi synthesized a conductive
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hydrogel from polypyrrole and polyacrylic acid. This composite showed zero-order release
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profile of safranin through electrical signals [12]. S. S. Indermun et al. examined release of
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indomethacin from conductive PVA-vinyl imidazole hydrogel [13]. However, it remains a
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challenge to consider all structural factors of the hydrogels [14]. Therefore, target application
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of hydrogels must be cleared in order to fast introduction of them into clinical formulation.
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However, there are some studies about in vivo drug release from conductive hydrogels. For
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example, methyl methacrylate-CNT hydrogel was used for in vivo release of
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[15].
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In this study, we synthesized an electro-responsive hydrogel containing biodegradable and
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available
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physiological systems of body, researchers have used collagen at tissue engineering
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frequently [16-17]. We found that the mechanical properties and film forming of hydrolyzed
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collagen improve via polycaprolactone modification. It has been well documented that
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polyaniline possesses good conductivity at emeraldine salt state as well as environment
87
stability [18]. Studies have shown that polyaniline has no destructive effects on tissue cells
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but there are concerns about chronic inflammation due to its non-biodegradability. K.
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Sengothi et al. showed that polyaniline have no bad or carcinogen effect on Sprague-Dawley
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rat cells after two years of implantation [19]. There are various methods for incorporation of
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conductive polymers into an insulator material. One of the most studied methods is in situ
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polymerization of monomer on the pores of the hydrogels [20-21]. Through polymerization
collagen
polypeptide.
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hydrogel. Here, the optimized hydrogel synthesis with Taguchi method was carried out and
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its potential as an extra controlled drug release systems examined with hydrocortisone
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hydrochloride as a model drug.
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2. Experimental
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2.1. Materials
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Hydrolyzed collagen type II (from Parvar Novin-E Tehran, Iran), acrylic acid (distilled under
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reduced pressure), N,N´-methylene bisacrylamide (MBA from Loba chemie, India),
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ammonium persulfate (APS from Fulka), aniline (distilled under reduced pressure), stannous
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octoate (Sn(Oct)2), ε-caprolactone (from Merck) and Hydrocortisone (from Jaber Ebne
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Hayyan pharmaceutical Co., Iran) were used without further purification.
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The vibrational spectra were measured using FTIR spectrometer (MBB Bommem, MB-100)
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in the range of 400 to 4000 cm-1. Samples were compressed into a KBr disk for FTIR
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measurements. Thermal characteristics of the samples were studied using thermal gravimetric
108
analysis (TGA). TGA was carried out under a nitrogen atmosphere with a heating rate of 10
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ºC/min and the scan range from 30 to 600 ºC. Veeco’s Autoprobe (CP-research) atomic force
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microscope in non-contact mode and field-emission scanning electron microscope (FE-SEM,
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S-4160) were used for morphological studies. Voltammetry measurement was carried out
112
using an electrochemical cell made up of gold wire working electrode, platinum wire counter
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electrode and Ag/AgCl reference electrode immersed in phosphate buffer solution with
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pH=7. UV-Vis spectra were measured from 200 to 800 nm in DMSO using PerkinElmer
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(Lambda 25) UV-Vis spectrometer. A Bruker (DRX-500 Avanes) NMR was used to record
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used for mechanical measurements of the hydrogels under tensile loading force 1 kN.
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2.3. Modification of the hydrolyzed collagen with polycaprolactone (C-PCL)
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3 g of hydrolyzed collagen was dissolved in 7 mL of distilled water and separated from
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insoluble phosphate salts. 3 mL of ε-caprolactone (0.027 mol) was added into the
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collagen solution and the mixture stirred in 80 ºC for one hour. Then, 200 µL of Sn(Oct)2
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was added to start ring opening polymerization of ε-caprolactone. The reaction was
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allowed to continue at 110 ºC for 12 hours and then, the unreacted monomers and
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homopolymers of ε-caprolactone were separated by liquid extraction with toluene at 60
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ºC for 24 hours. The modified hydrolyzed collagen (C-PCL) in aqueous phase was dried
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at 50 ºC for three days.
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2.4. Experimental design
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Repeatability is an important characteristic that must be considered in order to achieve
129
appropriate target drug delivery systems. It is understood that swelling of the hydrogels
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influences both uptake and release of the drugs and thus, all of structural parameters
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which are known to be important must take into account. Taguchi model as a powerful
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statistic method is able to provide an approach to optimize these structural parameters
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[22].
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Accordingly, an orthogonal array for three parameters (initiator, crosslinker and
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monomer) in three levels was designed as presented in Table 1 and Table 2.
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Table 1
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Table 2
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and contribution of each factor.
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2.5. Synthesis of conductive hydrogel (CH)
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In general, 1.00 g of modified collagen was dissolved in 40.0 mL of distilled water and
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stirred at 80 ºC. After 2 minutes certain amount of APS (0.03-0.1) was added for macro
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radical initiator formation. After 2 minutes acrylic acid (3-5 mL) and several minutes later
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MBA (0.05-0.2) were added respectively and finally, hydrogels formed during 15-60
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minutes. The hydrogels were placed in distilled water for 2 days to remove unreacted
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monomers and then dried at 50 ºC for 3 days. The conductive polyaniline insertion was
149
done through in situ polymerization of aniline. In a typical procedure the hydrogels were
150
swollen in HCl solution (1 M) containing aniline (1 mL per 5 g of dry hydrogel). After
151
one day the hydrogels were removed and washed with water then, placed in 1.1 M APS
152
solution. In situ polymerization of soaked aniline and color change from yellow to green
153
was performed gradually in the hydrogel matrix during one hour.
154
Swelling measurements were done by the gravimetric method. The Equilibrium Swelling
155
(ES) capacity was calculated at room temperature using the following formula (Eq. 1):
156
ES (g/g) = (W2 – W1)/W1
157
where W1 and W2 are the weights of dry and swollen gel, respectively.
158
2.6. Drug loading (DL) and release profiles
159
100.0 mg of hydrocortisone was dissolved in saline phosphate buffer (PBS) at pH=7.
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5.00 g of dry conductive hydrogel (CH) (1.5×0.5×0.2 cm pieces) was placed in the drug
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solution for one day. The swollen hydrogels were raised and placed in the fresh buffer in
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order to remove the weak adsorbent drugs in the surface of the hydrogels.
(Eq. 1)
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and placed in 20 mL of phosphate buffer with pH=7 for 3 days. The concentration of drug
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was determined using UV-Vis absorption peak at 245 nm. Eq. 2 was used for calculation
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of actual amount of the drug:
167
DL=
(Eq. 2)
× 100
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where A is the actual quantity of hydrocortisone which determined with UV-Vis
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spectroscopy and A0 is the theoretical amount of the drug.
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For drug release investigation, PBS (NaCl: 0.8 g, KCl: 0.02 g, Na2HPO4: 0.178 g,
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KH2PO4: 0.027 g) with pH=5.5 was chosen for simulation of skin conditions. Five
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different release profiles were investigated. At the first release experiment, samples were
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placed in 30.0 mL of PBS at 37 ºC and 3.00 mL of buffer solution were withdrawn at
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selected times (5, 8, 18, 21, 31, 34, 44, 47, 57 and 60 minutes after incubation) for drug
175
concentration analysis and returned to the mother solution (Release 1).
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At the second release experiment, samples were placed in 30.0 mL of PBS at 37 ºC and
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potential difference (3 V) applied to the hydrogel by using two electrodes for 3 minutes in
178
the selected times (5, 18, 31, 44, 57 minutes after incubation). 1 mL of PBS was
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withdrawn with the same time intervals as Release 1 and diluted to 10 mL for
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determination of the drug concentration with UV-Vis analysis (Release 2).
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The third release experiment was performed quite similar to the Release 2 profile except
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that potential difference between two electrodes was 1.5 V (Release 3).
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The fourth and fifth experiments were done similar to the Release 2 and 3 respectively,
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expect that duration time of the applied potential was 1 minute (Release 4 and 5). UV-
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minutes after incubation). All reported data are the average of three experiments.
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2.7. Cell viability
188
To evaluate cytotoxicity of the conductive hydrogel, MTT colorimetric assay based on
189
general methods was used. For extract of products, sterilized samples (8 cm3) using UV
190
irradiation were placed in individual wells and added cell culture media (EMEM (EBSS)
191
+ 1% NEAA + Bovine insulin-optional +10% FBS) in each well and incubated at 37 ºC
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for one week. This media were removed for MTT assay. A431 (human skin cell
193
carcinoma) were seeded at a density of 1×104 cells/mL on a 96-well tissue culture
194
polystyrene plate and incubated at 37 ºC for 24 hours. The extract of the products (at
195
various concentration) were added to the wells and cells cultured for more 24 hours. Then
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the culture medium was removed and 100 µL of MTT solution diluted in PBS (0.5
197
mg/mL) added to the wells. Plate incubated at 37 ºC for 4 h. after removal of the medium,
198
isopropanol solution was added to each well in order to complete dissolving of the
199
produced blue crystals. The absorbance was measured at 545 nm using spectrometric
200
microplate reader. The cell viability was calculated by normalization of optical density of
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the samples to the control. . The wells with more alive cells show higher optical density
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(OD) therefor, the cell viability could be calculated using equation (3) and (4).
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Toxicity % = (1 – (mean OD of sample / mean OD of control) × 100
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Viability % = 100 – Toxicity % Eq. (4)
Eq. (3)
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3. Result and Discussion
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3.1. Synthesis of the conductive hydrogel
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In this study, a conductive hydrogel based on hydrolyzed collagen was synthesized.
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Hydrolyzed collagen was modified with polycaprolactone which is frequently used at
209
biological systems like tissue engineering and drug delivery systems. Besides the
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biodegradability, hydrophobic characteristic of polycaprolactone may enhances the
211
interaction with non-ionic drugs. P. Dubois et al. modified polysaccharides with
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polycaprolactone and observed the improvement of mechanical and biodegradation
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214
another reasons that lead us to modify hydrolyzed collagen with polycaprolactone.
215
Therefore, ring opening polymerization of ε-caprolactone on hydrolyzed collagen using
216
Sn(Oct)2 as the catalyst was performed. As shown in Scheme 1, hydrolyzed collagen
217
powder was transformed to a viscose yellow liquid after grafting of polycaprolactone.
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The loading of PCL on the polymer backbone is approximately 50-60 % determined with
219
weight difference between the starting material and the product. Co-polymerization of
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acrylic acid onto backbone of the modified collagen was carried out in the presence of
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MBA as a crosslinking agent.
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Scheme 1
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The water absorbency results of the nine hydrogels proposed by Taguchi’s model are
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listed in Table 3. Two batch of each sample were synthesized and presented as ES1 and
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ES2. Swelling measurements were repeated twice.
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Table 3
The optimum conditions analyzed with Minitab 16 software were obtained (4 mL of
228
acrylic acid, 0.05 g of MBA and 0.1 g of APS). However, poor mechanical properties of
229
this hydrogel was not suitable for our quests. Using the Taguchi’s model suggestion, we
230
examined other conditions which possess near swilling ratio to the optimized value.
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Finally, a hydrogel (4 mL AA, 0.05 g APS, 0.1 g MBA) with both good water absorbency
232
and mechanical properties was synthesized for our following studies.
233
In order to achieve conductive hydrogel (CH), aniline monomers diffused through
234
hydrogel pores and nanofibres of polyaniline formed due to in situ polymerization in the
235
APS/HCl solution. In order to prevent clogging of the pores in the early stages of
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polymerization, the surface of hydrogels were washed thoroughly before incubation in the
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238
of aniline which is obvious from the shown cross section photograph (Scheme 1).
239
3.2. Chemical characterization
240
FT-IR and 1H NMR analyses were used for characterization of the synthesized C-PCL. In
241
FT-IR spectrum (Fig. 1S), the peaks at 1120 and 1730 cm-1 correspond to the C-O and
242
C=O stretching bands of polyester respectively and verify the presence of
243
polycaprolactone in the C-PCL. Moreover, the intensity of alkyl C-H stretching band at
244
2950 cm-1 has increased after modification. A representative 1H NMR spectra of
245
hydrolyzed collagen and C-PCL are shown in Fig. 2S. The peaks were assigned to the
246
corresponding hydrogen atoms of the polycaprolactone in Fig. 2Sb. The intensity of the
247
hydrolyzed collagen hydrogen peaks at C-PCL has decreased in the presence of the
248
polycaprolactone.
249
Polyaniline incorporation into the hydrogel matrix was confirmed with comparision
250
between FTIR spectra of polyaniline, non-conductive and conductive hydrogels (Fig. 3S).
251
3.3. Thermal and mechanical characterization
252
The TGA thermograms of C-PCL hydrogel (sample 1) and conductive hydrogel (CH)
253
(sample 2) are plotted in Fig. 1. Sample 2 has the same degradation process as sample 1,
254
which is attributed to the degradation of hydrolyzed collagen and polycaprolactone
255
moiety. The slight difference in the thermal stability of the two samples in the range of
256
350 to 600 ºC is due to the presence of polyaniline segment in the conductive hydrogel
257
(sample 2).
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Figure 1
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ACCEPTED MANUSCRIPT The mechanical strength of the hydrogels are dependent on the amount of absorbent
260
water. The more water absorbance, the lower the mechanical strength. Incorporation of
261
polycaprolactone and polyaniline decreases the water absorbency and consequently
262
improves the mechanical strength of the hydrogel. The Young’s modulus of proposed
263
hydrogels for tissue engineering applications is 600 to 1600 KPa [23]. Both hydrolyzed
264
collagen (sample A) and modified collagen (sample B) hydrogels were examined for
265
mechanical tensile analysis (Fig. 2). The Young’s modulus of sample B had a huge
266
increase compared to the sample A (from 93 to 340 MPa) with higher break elongation.
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This suggests that polyester moiety could effectively improves the mechanical strength of
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the hydrogel at wet state which is very important for biomedical application. This is
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probably related to the good mechanical property and low water absorbency of the
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polycaprolactone segment.
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Figure 2
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3.4. Morphological studies
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In order to correlate the current-stimuli drug release behavior of the hydrogel to its
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physical characteristics changes, SEM and AFM experiments were carried out before and
275
after applying potential.
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As illustrated in Fig. 3, AFM images exhibit morphological changes of the hydrogel after
277
applying potential in 10 minutes. AFM images of the CH before (Fig. 3a) and after (Fig.
278
3b) electrical field applying demonstrate that the surface roughness changed and more
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smoothness along with pore size expansion performed as illustrated in Fig. 3c. Through
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current flow, destructions take place in the surface of the hydrogel and size expansion of
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pores causes the drug release.
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ACCEPTED MANUSCRIPT Moreover, SEM images of CH films before (Fig. 4a) and after (Fig. 4b) applying
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potential demonstrate the effect of electrical field on the surface of the hydrogel. Surface
285
porosity has changed through electrical field due to the polyaniline reduction and
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hydrogel backbone destruction. Hydrogel pore size before the potential applying is
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around of 10-50 nm (Fig. 4S). Formation of large pore size (approximately 50-100 nm) in
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the presence of electrical field have been shown in Fig. 4b.
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3.5. Electrical characteristics of conductive hydrogel
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Cyclic voltammetry graph of the polyaniline segment in phosphate buffer solution is
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presented in Fig. 5, which working electrode was in contact with surface of the CH. It is
293
clear that cyclic voltammetry behavior of polyaniline has been preserved at the hydrogel
294
matrix. Two oxidation peaks at 0.25 and 0.65 V are related to the leucoemeraldine-
295
emeraldine and emeraldine-pernigraniline equilibrium states of the polyaniline
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respectively. The reduction peak is attributed to the protonated emeraldine and
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leucoemeraldine equilibrium states of the polyaniline.
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Figure 5
The surface resistivity of CH films were measured in both sides before and after
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performing 3 V potential in 10 minutes. Experiments have provided evidences that semi-
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conductive films with surface resistivity of 1.5×10-4 to 3×10-5 S.cm2 are acceptable for
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biomedical applications [2]. As illustrated in Table 4, the surface resistivity of the films
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increased after potential applying. This result could be related to the proton exchange
304
through polyaniline chains due to current flow which increases the conductivity of the
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film. In the other related studies, volume resistivity have been measured instead of
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surface resistivity usually. For example, the volume conductivity of 0.3-5 mS/cm for
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collagen/CNT [17] and 1.1×10-3 mS/cm [3] or 10-6 [24] for hydrogel/polyaniline
308
composites were reported. Table 4
310
For UV-Vis measurement, CH was cut into small portions and then immersed in DMSO.
311
The color of solvent changed due to polyaniline dissolving after 30 minutes. The UV-Vis
312
spectrum (Fig. 6 solid line) demonstrates that polyaniline has preserved its optoelectronic
313
characteristics in the hydrogel. Peaks at 320 and 600 nm are assigned to the π-π*
314
transition of the benzene ring and the benzenoid to quinoid excitonic transition
315
respectively. UV-Vis spectrum of the polyaniline alone is shown with dash line as
316
control.
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Figure 6
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3.6. Swelling behavior of the hydrogels in various conditions
319
The swelling behavior of the hydrogels (conductive and non-conductive form) in the PBS
320
were measured through 3 hours (Fig. 7). The swelling degree of the hydrogels is
321
dependent on the incorporation of conductive polyaniline or applying the electrical field.
322
The swelling ratio of B, D and E samples were determined by immersing the pre-weight
323
dry hydrogels in buffer solution and applying electric potential every 5 minutes.
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It is clear that non-conductive hydrogels swell rapidly than conductive hydrogels because
325
of hydrophobic characteristic of polyaniline. Potential applying had no notable influence
326
on the swelling behavior of the non-conductive hydrogels (A and B). On the contrary, the
327
swelling behavior of the CHs were dependent on the current flow as seen in Fig. 7. As a
328
result of current flow, polyaniline reduction and expansion of pores takes place causing
329
the water uptake increasing. By applying higher potential, the swelling ratio of the CHs
330
have increased. But why current flow have no influence on swelling of the non-
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ACCEPTED MANUSCRIPT conductive hydrogels? There are two possible reasons: Firstly, non-conductive hydrogels
332
have their maximum swelling degree and secondly, in the CHs, current flow makes
333
changes in the oxidation states of polyaniline which is the reason of pore size expansion
334
and increasing of swelling ratio. But in the non-conductive hydrogel matrix, ions have the
335
role of current transition which have no effect on the hydrogel morphology. Figure 7
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3.7. Drug release study
338
Drug entrapment efficiency is one of the most important considerations in drug delivery
339
systems. As illustrated in Fig. 8, the drug entrapment percent was calculated for
340
conductive and non-conductive hydrogels seventh times. The average of these values are
341
59.3 and 74.45 % for conductive and non-conductive hydrogels respectively. Lower
342
swelling ratio of the CH compared to the non-conductive hydrogel due to hydrophobic
343
characteristics of polyaniline causes the reduction of drug entrapment efficiency.
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Figure 8
For drug release studies, the non-ionic hydrocortisone was chosen as a model drug and a
346
simple two electrode method was carried out for applying electric potential in the saline
347
buffer solution (pH=5.5). The concentration of the released drug was followed by UV-Vis
348
spectroscopy. Drug loading into the CH was carried out by swelling of the hydrogel in
349
the drug solution with appropriate concentration (100 mg of drug per 5 g of dry hydrogel)
350
and subsequently immersed in the fresh buffer solution to detach the adsorbent drugs on
351
the surface of the hydrogel. The drug release was performed at 37 ºC in buffer solution
352
and potential applied through two electrode (2 cm height and 2 mm diameter). For
353
acceptable comparison, the hydrogel films were cut into the identical rectangular shape
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ACCEPTED MANUSCRIPT (1.5×0.5 cm approximately 0.25 g).
355
mechanism, we examined five release methods through one hour.
356
In the first method (Release 1 Fig. 9a), there was no current flow, and the drug release
357
was performed via swelling of the hydrogel in buffer solution leading to exit of the drug
358
from the matrix pores. There was slight increase of the drug release percent at early
359
stages, but its maximum value reached to 40 % after 1 day incubation.
360
In the next release method, electrical potential with various strength and duration time
361
were applied. At the second experiment (Release 2, Fig. 9b), 3 V potential was applied
362
between two electrodes for three minutes in each interval times. However, at the first
363
electric pulse complete destruction of the hydrogel took place and 90% of the drug
364
released from the CH. Potential of 1.5 V for 3 minutes duration times (Release 3 Fig. 9c)
365
made more controllable drug release approach. The whole drug removed from the
366
hydrogel after one hour, in the other words, 15 % of the drug released in each step
367
approximately. At Release 4 (Fig. 9d), through applying potential difference of 1.5 V for
368
1 minute 60 % of the drug released after 1 hour; i.e. 7 % in each step. Due to potential
369
difference of 1.5 V between two electrodes (Release 5, Fig. 9e), 3 % of the drug released
370
per one minutes. (27 % after one hour).
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For accurate investigation of the drug release
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Figure 9
372
If drug interacts strongly with hydrogel scaffold through electrostatic forces, its release
373
without any driving force will perform slowly. The proposed mechanism of drug release
374
in current stimuli drug delivery systems, like our case study, is direct expansion of the
375
pore size and increasing the free volume in the hydrogel matrix. It happens when the
376
oxidation state of the conductive polymer changes (such as polyaniline) and consequently
377
leads the drugs to the electrodes with different charges or to the hydrogel pores. Pore size
16
ACCEPTED MANUSCRIPT expansion of the hydrogels may takes place due to the migration of water or charged ions
379
from the hydrogel to the electrodes [25]. The influence of electrical field strength on the
380
drug release is dependent on the mechanism of drug release as well as interaction of drug
381
with hydrogel scaffold.
382
Hydrocortisone, which known as a non-ionic drug can interacts with hydrogel scaffold
383
via two routes: hydrogen bonding with polar functional groups and van der Waals
384
interaction with PCL. Experiments demonstrated that, electrical signals were able to
385
speed up the release of hydrocortisone. Upon using this technique, drug release can be
386
regulated by strength and duration of applied electrical field. Both morphological studies
387
(AFM and SEM) and drug release experiments confirmed that meanwhile the expansion
388
of pore size of the hydrogel due to electrical signals, the drug has penetrated through the
389
surface and finally, released.
390
3.8. Cell viability
391
Cell viability was measured using MTT assay and results are represented in Fig. 10. The
392
cell viability numbers over 80 % for the hydrogels at non-conductive and conductive
393
form (Sample 1 and Sample 2 respectively) indicated no toxicity of the hydrogels. The
394
cell viability results based on the in vitro MTT method are shown in two different
395
concentrations (Sample 1' and 2' are 10 times diluted than Sample 1 and 2). Therefore,
396
polycaprolactone modification in Sample 1 and polyaniline incorporation in Sample 2
397
have no harmful effects on the hydrolyzed collagen-based hydrogels, which is a primary
398
concern for their in vivo drug release study.
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400
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4. Conclusion
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ACCEPTED MANUSCRIPT In this study, a semi-conductive hydrogel film was investigated as a current stimuli drug
402
delivery system. Hydrolyzed collagen as a biodegradable scaffold was modified with
403
polycaprolactone and then, through co-polymerization of acrylic acid and subsequently in
404
situ polymerization of polyaniline the conductive hydrogel was produced. The various
405
characterization methods and drug release experiments well illustrated that this
406
nanocomposite hydrogel could be used as a promising current stimuli drug delivery
407
system.
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ACCEPTED MANUSCRIPT 408
Figure captions:
410
Figure 1: TGA curves of a) C-PCL based hydrogel b) CH
411
Figure 2: stress/elongation curves of a) hydrolyzed collagen based hydrogel b) C-PCL
412
based hydrogel with 20% humidity at room temperature
413
Figure 3: AFM images of CH films a) before and b) after applying potential of 3 V for
414
10 minutes c) a representative for morphological change of the hydrogel through
415
application of 3 V potential for 10 minutes
416
Figure 4: SEM images of the CH films a) before and b) after application of 3 V potential
417
for 10 minutes
418
Figure 5: Cyclic voltammograms of conductive hydrogel in aqueous buffer solutions
419
pH=7. Scan rate: 10 mV/s.
420
Figure 6: UV-Vis spectrum of the CH immersed in DMSO (polyaniline alone as a
421
control represented with dashed line)
422
Figure 7: Swelling ratio in terms of time for the non-conductive (A and B) and
423
conductive (C-E) hydrogels with different condition: No electrical signals (A and C);
424
applying 3 V potential (B and E); applying 1.5 V potential (D)
425
Figure 8: Drug entrapment values of a) non-conductive and b) conductive forms of the
426
hydrogel
427
Figure 9: drug release study in phosphate buffer pH=5.5, 37 ºC, a) No electrical
428
potential. In the electric potential of b) 3 V for 3 min c) 1.5 V for 3 min d) 3 V for 1 min
429
e) 1 V for 1 min.
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ACCEPTED MANUSCRIPT Figure 10: The cell viability of the polycaprolactone modified hydrogel (Sample 1) and
431
conductive hydrogel (Sample 2)
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Table 1: Experimental control factor and their levels Terial A B C 1 1 1 1 2 1 2 2 3 1 3 3 4 2 1 2 5 2 2 3 6 2 3 1 7 3 1 3 8 3 2 1 9 3 3 2
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Table 2: Experimental layouts of an L9 orthogonal array according to Taguchi’s suggestion Control factor Level 1 Level 2 Level 3 Acrylic Acid (mL) 3 4 5 APS (g) 0.03 0.05 0.1 MBA (g) 0.05 0.1 0.2 500
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3.89 4.05 4.44
11.34 5.57 7.046
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Table 3: Water absorbency of nine trials of hydrogels Trial 1 2 3 4 5 6 6.92 3.54 13.34 4.77 33.38 ES1(g/g) 24.51 5.28 6.57 6.38
27.26 29.4 22.52
7 5.37 5.18 5.9 6.09
8 12.23 12.38 10.83 10.76
Table 4: The surface resistivity of films (S.cm2): 1) Before applying potential 2) After applying potential of 3V for 10 minutes Surface 1 Surface 2 18×10-6 8×10-6 1 -3 1.82×10 1.456×10-3 2 503 504 505
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9 5.29 5.43 5.18 5.26
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Figure 1: TGA curves of a) C-PCL based hydrogel b) CH
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Figure 2: stress/elongation curves of a) hydrolyzed collagen based hydrogel b) C-PCL based
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hydrogel with 20% humidity at room temperature
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Figure 3: AFM images of CH films a) before and b) after applying potential of 3 V for 10 minutes c) a representative for morphological change of the hydrogel through application of
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3 V potential for 10 minutes
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Figure 4: SEM images of the CH films a) before and b) after application of 3 V potential for
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10 minutes
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Figure 5: Cyclic voltammograms of conductive hydrogel in aqueous buffer solutions pH=7.
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Scan rate: 10 mV/s.
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Figure 6: UV-Vis spectrum of the CH immersed in DMSO (polyaniline alone as a control
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represented with dashed line)
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Figure 7: Swelling ratio in terms of time for the non-conductive (A and B) and conductive (C-E) hydrogels with different condition: No electrical signals (A and C); applying 3 V
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potential (B and E); applying 1.5 V potential (D)
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Figure 8: Drug entrapment values of a) non-conductive and b) conductive forms of the
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hydrogel
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Figure 9: drug release study in phosphate buffer pH=5.5, 37 ºC, a) No electrical potential. In
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the electric potential of b) 3 V for 3 min c) 1.5 V for 3 min d) 3 V for 1 min e) 1 V for 1 min.
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Figure 10: The cell viability of the polycaprolactone modified hydrogel (Sample 1) and
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conductive hydrogel (Sample 2)
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and applied for in vitro hydrocortisone release.
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4) A current-stimuli drug delivery system proposed by Taguchi method was designed
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Supplemental Material
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Supplemental Figure 1: FTIR spectra of a) hydrolyzed collagen and b) modified collagen
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Supplemental Figure 2: H NMR spectra of a) hydrolyzed collagen and b) C-PCL in DMSO Supplemental Figure 1: FTIR spectra of a) hydrolyzed collagen and b) modified collagen
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Supplemental Figure 3: FTIR spectra of a) Polyaniline b) non-conductive hydrogel and c)
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conductive hydrogel.
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Supplemental Figure 4: SEM image of the hydrogel before potential applying