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A novel multilayer hydrogel wound dressing for antibiotic release
Journal of Drug Delivery Science and Technology 58 (2020) 101536
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A novel multilayer hydrogel wound dressing for antibiotic release Emel Tamahkar
a,b,∗,1
, Bengi Özkahraman
, Aysun Kılıç Süloğlu , Neslihan İdil , Işık Perçin
c,∗∗,1
d
d
T d
a
Hitit University, Department of Chemical Engineering, Çorum, 19030, Turkey Balıkesir University, Department of Food Engineering, Balıkesir, 10145, Turkey Hitit University, Department of Polymer Engineering, Çorum, 19030, Turkey d Hacettepe University, Department of Biology, Ankara, 06800, Turkey b c
A R T I C LE I N FO
A B S T R A C T
Keywords: Multilayer hydrogels Wound dressing Antibiotic release Antibacterial activity
In this study, the novel multilayer (ML) hydrogels were developed as antibacterial wound dressings. ML hydrogels were prepared as four layers using carboxylated polyvinyl alcohol (PVA-C), gelatin (G), hyaluronic acid (HA) and gelatin respectively. The upper layers (PVA-C and G) provide the moist control and physical barrier for microorganisms. The HA based middle layer was designed as an antibiotic-loaded layer. The lower layer serves as the controlling membrane for antibiotic release and provides the removal of excess exudate from the wound site. The ML hydrogels were characterized with FT-IR, SEM, DSC, swelling tests and hydrolytic degradation tests. Cell viability assay was also performed in L929 fibroblast cells in vitro. The in-vitro drug release profile of ML hydrogels was investigated at pH 7.4 at 37 °C and it was found that 63% of the antibiotic was released after 7 days.
1. Introduction
adhesion onto the wound surface and support for new tissue formation [6,7]. Multilayer hydrogels consisting a drug-loaded layer provide controlled drug release for long time providing mass transfer limitations for drug molecules throughout the polymeric matrix [8]. Layer-by-layer (LBL) approach can be applied to broad range of polymers enabling the design of functional biomaterials with desired properties without any additional chemicals. The preparation of multilayer hydrogels via LBL technique depends on the adsorption of the polymers having oppositely charged groups layer by layer providing the incorporation of them into multilayers [9]. Multilayer hydrogels fabricated by LBL self assembly method comprised of different polymeric layers present great attention for drug delivery with their facile preparation procedure, tunable morphological features and high biocompatibility providing the drug molecules released from the matrix in a controlled manner [10]. ReyesOrtega et al. produced a bilayer wound dressing based on gelatin-HA that consisted on a first layer of cross-linked gelatin-HA loaded with proadrenomedullin N-terminal 20 peptide and a second layer of polyurethane derivative that was loaded with nanoparticles of bemiparin. The bilayer hydrogels were investigated using in-vivo mouse models. The results showed that the bilayer wound dressing presents promising alternative since they reduced the contraction and the inflammation with the aid of the controlled release of the two drugs [11]. Ng and Tan synthesized a bilayer film based on alginate and gelatin containing
Wound healing including the phases of hemostasis, inflammation, proliferation and remodelling is a complex and long process since the remodelling proceeds for 21 days - 1 year forming suitable conditions for bacterial growth [1]. Selection of wound dressing that depends on many factors such as type and depth of wound, amount of exudate etc. becomes significant for an effective treatment. An ideal wound dressing should provide moist environment that facilitates healing process [2]. It should allow adequate gas permeability and remove excess exudate from the wound surface while keeping it moist. Also, it should serve as a barrier for microorganisms providing an antibacterial medium. Furthermore it should be biocompatible, biodegradable, cheap, easy-toapply without requiring frequent changing [3]. Recently, multilayer wound dressings enabling better healing process than single layer dressings have become promising alternatives [4]. Multilayer hydrogels consisting of more than two layers are designed to meet the requirements of wound healing by the choice of chemical and physical properties of the constituent biomaterials. Multilayer hydrogel wound dressings integrates the advantages of each constituents. The upper layer serves as a barrier for bacterial transition and as a controlling membrane for moist environment [5]. The lower layer with spongy structure is utilized for the absorption of the excess exudate, ∗
Corresponding author. Hitit University, Department of Chemical Engineering, Çorum, 19030, Turkey. Corresponding author. E-mail addresses: [email protected] (E. Tamahkar), [email protected] (B. Özkahraman). 1 These authors contributed equally to this work. ∗∗
https://doi.org/10.1016/j.jddst.2020.101536 Received 13 December 2019; Received in revised form 9 January 2020; Accepted 21 January 2020 Available online 23 January 2020 1773-2247/ © 2020 Published by Elsevier B.V.
Journal of Drug Delivery Science and Technology 58 (2020) 101536
E. Tamahkar, et al.
25 °C, and precipitated with acetone. The product was dried for 24 h under vacuum at 25 °C.
hydroxytyrosol for topical chemotherapy. Glycerol and poly propylene glycol were used as plasticizers. The drug-free bilayer films were prepared by changing plasticizers’ ratio and they were characterized for their such as physical, mechanical and rheological properties. The invitro drug release results showed that bilayer films had lower drug release rate than single films [12]. Ding et al. developed a multilayer hydrogel film using hyaluronic acid and PVA for insulin release. It was determined that 90% of insulin was released from the resultant multilayer hydrogel films that were produced via vacuum-drying after 12 h since the release of 90% of insulin was reached after 6 h with the prepared multilayer hydrogel films synthesized by lyophilization [13]. Although there exist significant improvements of the fabrication of the wound dressings, the development of novel antibacterial multilayer wound dressings with transparency, cost-efficiency, high healing capacity is still of great interest [14]. Infection is one of the major factors, which delays the healing process. Topical administration of an antibiotic drug at the wound site is one of the most important aspects for the prevention of infection during healing process. Controlled drug release present great advantages when applied to wound dressings providing delivery of the drug molecules through the wound site in a controlled manner for long time periods without need for the frequent replacement of the dressing material [15]. Also the utilization of drug-loaded wound dressings enable delivering the antibiotics in a determined dosage avoiding high systemic dose [16]. There are many reports about drug-loaded wound dressings with sustained antibiotic release [17,18]. The aim of this study was to develop multilayer hydrogel wound dressings composed of natural polymers with water-based approach for effective wound healing. The multilayer hydrogel film was composed of 4 layers of which PVA-based first layer, gelatin-based second layer, ampicillin loaded HA-based third layer and gelatin-based fourth layer. For the best knowledge of the authors, this is the first report concerning the development of multilayer hydrogel films consisting of PVA-C, gelatin and hyaluronic acid as a new system for wound healing and drug delivery. The drug-free and drug-loaded multilayer films were synthesized via solvent evaporation and characterized by FT-IR, SEM and DSC measurements. Also the swelling degree and hydrolytic degradation ratio of the hydrogels was investigated. The results of in-vitro drug release tests showed that 63% of the loaded drug was released from the multilayer hydrogels after 5 days. On the other hand, antimicrobial performance of novel polymeric ampicillin loaded ML hydrogels were investigated against both oxacillin resistant and sensitive Staphylococcus aureus (S. aureus) and Escherichia coli (E. coli). Previous studies have shown that dressing material itself can induce toxicity on cells due to monomers or antimicrobials used. Therefore it is necessary to show that the novel dressing material is not toxic to the cells which would be included in healing process [19]. Thus, the in-vitro cytotoxicity of the multilayer hydrogels was also investigated.
2.3. Synthesis of multilayer hydrogel film A multilayer hydrogel film consisted of four layers. For first layer, PVA-C hydrogel film was prepared using the solvent-casting process. Polymer solution of PVA-C (10 wt %) was stirred under mechanical agitation for 3 h at 80 °C and casted into Teflon petri dishes. The first layer of wound healing films was crosslinked at 40 °C for 48 h. To fabricate second layer, gelatin solution (5 wt %) was prepared by dissolving at 50 °C and then casted on the first layer. The second layer was left at 25 °C for 3 h and subsequently dried under vacuum at 40 °C for 48 h. For preparation of the third layer, HA solution (5 wt %) was poured onto the existent two layer and then was left at room temperature for 3 h with subsequent vacuum drying at 40 °C for 48 h. Finally, the last layer was prepared by dissolving gelatin solution (5 wt %) at 50 °C and then it was casted on resultant three layered hydrogels. Ampicillin-loaded multilayer hydrogels (ML-D) was prepared in the same way as blank multilayer hydrogels (ML) by adding ampicillin (2 wt %) directly to the HA solution. The codes of the multilayer hydrogels prepared in this study were listed in Table 1.
2.4. Characterization studies The hydrogels were characterized by a FTIR spectrophotometer (Agilent Technologies Inc., USA). The morphological characterization of multi-layer hydrogels was taken by a FEQ Scanning electron microscope (SEM). Differential scanning calorimetry (DSC) was utilized for thermal analysis of the multilayer hydrogels. The heating profile was established from 25 °C to 300 °C at a rate of 5 °C/min.
2.5. Swelling studies The swelling study of the multilayer wound healing hydrogels was determined gravimetrically using different buffer solutions (pH 5.5 and pH 7.4) during 24 h. The multilayer hydrogel was immersed in these solutions at 37 °C. The swollen multilayer hydrogel was weighed at regular time intervals until equilibrium was reached after the excess water was wiped off from the hydrogel film surface. The swelling ratio (SR %) was calculated from the expression (1):
SR% = (Ws − Wd/ Wd ) ∗ 100
(1)
where Ws and Wd are the weights of the swollen hydrogels at equilibrium and dry hydrogels, respectively.
2. Experimental
2.6. In-vitro hydrolytic degradation
2.1. Chemicals
In vitro hydrolytic degradability of the multilayer hydrogel (W0) was investigated using PBS at 37 °C to provide a similar environment with wound exudate. At determined time intervals, the remaining hydrogel was taken out from the solution and dried for 24 h (Wt). The weight loss percentage was calculated with the equation (2) as follows:
Polyvinyl alcohol (PVA) (Mw: 30000-70000), ampicillin sodium salt, 4-dimethylaminopyridine (DMAP) and phosphate buffer tablet were obtained from Sigma-Aldrich (St. Louis, MO). Gelatin (G, type A) was obtained Biomatik (Wilmington, Delaware, USA). Sodium hyaluronate was purchased from Across Organic (Geel, Belgium) products. All other reagents were of analytical grade.
Weight loss (%) = (Wo − Wt / Wt ) ∗ 100
(2)
Table 1 The codes of the multi-layer hydrogels.
2.2. Synthesis of carboxylated PVA PVA (10 g) powder was dissolved in deionized water (100 mL) at 90 °C. DMAP (2.77 g) was mixed with the solution and stirred for 1 h at 65 °C. Succinic anhydride (23 g) was mixed with the reaction medium and stirred for 24 h at 65 °C. The PVA-COOH solution was cooled to 2
Polymer code
First layer
Second layer
Third layer
Fourth layer
ML ML-D
PVA-C PVA-C
G G
HA HA-Amp
G G
Journal of Drug Delivery Science and Technology 58 (2020) 101536
E. Tamahkar, et al.
2.7. Drug release studies The in vitro drug release profile of ML-D hydrogel was evaluated in pH 7.4 at 37 °C for one week. At scheduled time intervals, 2 mL was taken from the release medium and replaced with same volume of fresh buffer solutions. The amount of released drug from the drug-loaded hydrogels was calculated with aid of calibration curves plotted using ampicillin solutions (pH 7.4). The amount of ampicillin release was determined utilizing T80+ UV/VIS Spectrometer (PG Instruments Ltd.) at 275 nm.
2.8. Antibacterial performances of multilayer hydrogels The antibacterial performances of multilayer hydrogels were examined against both oxacillin resistant and sensitive S. aureus and E. coli. The identification of these bacterial strains and the determination of their antibacterial susceptibilities were performed via automated system (Vitek 2 System, bioMérieux, USA). Antibacterial performances of multilayer hydrogels were investigated by agar disc diffusion assay. In the first step, 100 mL of Luria Bertani (LB) broth in a 250 mL Erlenmayer flask was prepared. Then, bacterial strains were inoculated into LB broth and incubated at 37 °C for 18 h to obtain fresh bacterial cultures. In the second step, bacterial culture suspensions at the exponential growth phase were adjusted according to 0.5 McFarland (1.5×108 CFU/mL). In the third step, 100 μL of prepared culture suspensions were inoculated onto LB agar plates. Ampicillin loaded-multilayer hydrogels, ampicillin free-multilayer hydrogel (negative control) and ampicillin (10 μg/disc, positive control) were placed onto the LB agar plates which were cultivated at 37 °C for 18 h. The antimicrobial performances of multilayer hydrogels were evaluated by measuring the diameters of the inhibition zones in mm. The antibacterial performance assays were done in triplicate.
2.9. Cell viability assay The cell viability properties of ML and ML-D hydrogels were investigated. L929 mouse fibroblast cells were obtained from the American Type Culture Collection. Cells were cultured in DMEM (Dulbecco's modified Eagle's medium) supplemented with L-glutamine, 10% (v/v) fetal bovine serum (FBS) (Cegrogen biotech, Germany) and 1% penicillin-streptomycin (Gibco, Life Technologies, USA) solution. Incubation was carried out in an oven with humidified atmosphere of 5% CO2, at 37 °C. L929 cell line (Passage 17) was studied. The medium was changed every two days. Cells were trypsinized when they reached 90% confluence. All hydrogel samples were left to sterilize at UV light in the cabin for 45 min. The parafilms which prevent cells to attach surface were cut for the size of 48 well plate and were waited for 20 min at 70% alcohol. Then the parafilms were placed into sterile distilled water and engrafted on well plate and lefted at UV light for 45 min. Before seeding of the cells, hydrogels were conditioned on DMEM medium with serum for 30 min. Cells were cultivated on two 75 cm2 flasks and seeded at two different stages on hydrogel samples. Cells were trypsinized and santrifuged at 800 rpm for 5 min. Pellet was divided into fifteen for 24, 48 and 72 h. For each well, the number of cells was 3 × 103 per well. MTT [3-(4,5-dimetiltiazol-2-il)-2,5-difeniltetrazolium bromid] solution (MTT, Sigma, St. Louis, MO, USA) was added to wells. Cells were lefted for 3 h incubation. After incubation, medium was aspired. DMSO/NH3 solution was prepared (5% NH3-95% DMSO), 200 μl per well was added to samples. Samples were transferred to non-sterile 96 well plate. Plates were lefted on shaker for 10–15 min. The OD value was measured using an ELISA plate reader of absorbance at 550 nm (Bio Tek Instruments Inc., Winooski, VT, USA). Same procedure was repeated for 48 and 72 h.
Fig. 1. FT-IR spectrums of PVA, PVA-C, gelatin, hyaluronic acid and multilayer hydrogel (A) and ampicillin, free ML hydrogel and ampicillin-loaded ML hydrogel (ML-D) (B).
3. Results and discussion In this study, ML hydrogels were prepared via layer-by-layer self assembly through electrostatic interactions between polymeric layers. We supposed that positively charged gelatin that has isoelectric point at around 7-9 would fabricate a stable layer when poured onto the negatively charged carboxylated PVA layer. Hyaluronic acid layer which has a negative charge was added to form an antibiotic-loaded layer as a drug carrier reservoir. Lastly, gelatin layer was introduced onto the resultant three layers. 3.1. Characterization studies Fig. 1A shows ATR-FTIR spectra of carboxylated PVA, HA, gelatin and ML hydrogel. The carboxylation of PVA was verified by FTIR measurements. The sharp new peak at around 1700 cm−1 that was assigned to carbonyl stretching mode of esters was observed after the esterification. The main characteristic bands of gelatin were 3
Journal of Drug Delivery Science and Technology 58 (2020) 101536
E. Tamahkar, et al.
Fig. 2. The cross-sectional SEM images (A) and optic photograph of multilayer hydrogels (B).
demonstrated at 1628, 1526, 1238 cm−1 corresponding to amide I, II and III respectively. The spectra of the ML hydrogel were characterized due to the amide bands at 1630, 1530 and 1239 cm−1. The incorporation of HA into the ML hydrogel structure was verified by the new peak appeared at 1039 cm−1 which attributes to C-O-C stretching of hyaluronic acid. Also the strength of the band at 3275 cm−1 of ML hydrogels assigned to hydroxyl group decreased with respect to the other constituents indicating the chemical interactions between the polymeric layers. The ATR-FTIR spectrum of ML-D hydrogels was displayed in Fig. 1B. The loading of ampicillin molecules into the HA layer of ML-D hydrogels was determined due to new peak observed at 2087 cm−1 referred to bending of S-C of ampicillin. Also the strength of the amide bands (1631, 1526 and 1238 cm−1) and the hydroxyl band (3275 cm−1) decreased with the addition of drug molecules throughout the polymeric matrix. The SEM images of multilayer hydrogels corresponding to crosssections performed through the thickness of the hydrogels were presented in Fig. 2. The images indicate the multilayered structure consisting of four layers fabricated with no obvious distinction between layers [20]. The upper layers were involved of carboxylated-PVA and gelatin; the drug-loaded layer was located at the middle; the lower layer was composed of gelatin. The upper layers were formed to control the water vapor since the lower layer that adheres to the wound surface was designed to serve as the controlling layer of the drug release process [21]. The drug-loaded layer that was composed of HA was determined to have macroporous structure with high interconnectivity [22]. This network structure could facilitate the swelling of the hydrogels. Fig. 2B shows the transparency of ML hydrogels. DSC thermograms of ML hydrogels and the physical blend of the used polymers (ML-Blend) were shown in Fig. 3. Tg values of ML and ML-Blend were found as 59.29 °C and 49.93 °C respectively. The first endothermic peaks of ML and ML-Blend were observed at 63.7 °C and 53 °C respectively. The second endothermic peaks of ML and ML-Blend were found at 147 °C and 92.5 °C respectively. It was determined that the endothermic peaks of the polymer blend were shifted to the higher temperatures with the preparation of multilayer hydrogels implying the change in the chemical structure of the polymers. These results offer the
Fig. 3. DSC thermograms of ML hydrogels and the physical blend of the used polymers.
fabrication of a more stable structure with respect to the polymerpolymer interactions between the layers of the multilayer hydrogels [23]. The in-vitro swelling tests were evaluated at pH 5.5 and pH 7.4. The equilibrium swelling ratios of ML hydrogels at pH 5.5 and 7.4 were calculated gravimetrically and found as 517.8% and 339.4% respectively. The mechanism of the swelling of the ML hydrogels depends on the pH of the medium and thus the ionization state of the carboxylic acid and amino groups that are located at the side chains of the polymers. When the pH of the medium increases above the pKa of side groups of the polymeric structure, the hydrogels become swollen due to the ionic repulsion. Therefore, the lower swelling degree observed at pH 7.4 may be explained due to the lack of the ionization state of the 4
Journal of Drug Delivery Science and Technology 58 (2020) 101536
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D hydrogels was occurred within the first 6 h. Then the release rate decreases exponentially during 3 days. Lastly, the release rate became constant until 7 days. The release profile of ampicillin loaded ML hydrogels presented a burst release to prevent the risk of the bacterial infection, which can be occurred during the initial injury [26]. The fast release rate of the drug molecules within the first 6 h can be due to their location in the polymeric structure since the slower release rate of ampicillin can be attributed to the need of time to reach the diffusion barriers. The controlled drug release profile was due to the presence of the extra layers acting as the controlling membranes for drug diffusion [27]. Also, the interaction between negatively charged COO− and positively charged NH3+, Na+, H+ could reduce the release rate. The wound dressing with prolonged antibiotic release could be used for longer time periods thus resolving the need for repeated dressing changes. The prepared wound dressing in this study exhibited significant potential by showing antibiotic release in a controlled manner for prolong duration of time eliminating the requirement of continual dressing change. Kamoun et al. developed polyvinyl alcohol-alginate hydrogels for ampicillin delivery. The hydrogels showed a burst release of 38–45% within 15 min and the total release of ampicillin was obtained after 6 h [28]. Kenawy et al. prepared polyvinyl alcohol-hydroxyethyl starch (HES) hydrogels with freeze thawing method for controlled release of ampicillin. The release amount of ampicillin was increased with the increasing amount of HES and the release equilibrium was attained after 6 h [29]. Bako et al. synthesized nanogel with poly-gamma-glutamic acid nanoparticles via photopolymerization. The loaded ampicillin into the resultant nanogels was released during 24 h with a burst release in the first 4 [30]. Poonguzhali et al. prepared alginate based film with the incorporation of nanocellulose for ampicillin release. The release of ampicillin molecules from nanocomposite film obtained equilibrium after 500 min [31]. All these literature for ampicillin release show the production of biomaterials with fast release kinetics. MLD hydrogels having four layer enable the release of ampicillin for prolonged time periods. In order to describe the release profile of ampicillin from ML-D hydrogels, zero order, first order, Higuchi model and Korsmeyer-Peppas model were applied to release data. The mathematical expressions of zero order model (1a), first order model (2b), Higuchi model (3) and Korsmeyer-Peppas model (4) are described as follows:
Fig. 4. The in-vitro hydrolytic degradation of ML hydrogels.
side groups. The in-vitro swelling degree is one of the major properties of the wound dressings indicating the capability of the absorption of the excess exudate form the wound site which, causes possible bacterial infection. It has been reported that the typical wound (10 cm2) produces 5 mL of exudate in 24 h [24]. The ML hydrogels prepared in this study have swelling degree of 5.17 g/g (pH 5.5) and 3.39 g/g (pH 7.4) implying the suitability of ML hydrogels as a wound dressing. The in-vitro hydrolytic degradation of the ML hydrogels was tested by quantifying the mass loss after incubating the dressings in PBS at 37 °C. The hydrolytic degradation ratios were shown in Fig. 4. It was shown that the mass loss increased with increasing time and the degradation was completed after 15 days exhibiting a long-term degradation profile with good stability thus makes the prepared hydrogels a good alternative as wound dressings. The in-vitro drug release from ampicillin loaded multilayer hydrogels was shown in Fig. 5. Almost 65% of ampicillin was released within 7 days indicates that a long-term release profile could be achieved using the prepared multilayer hydrogels. The mechanism of the drug release from ML-D hydrogels was occurred such as hydration of the polymeric structure, swelling of the hydrogel and diffusion of the drug molecules throughout the swollen matrix [25]. The release profile can be separated into three main phases. The initial burst release (34.5%) from ML-
Qt = Q0 + k 0 t
(1a)
lnQt = lnQ0 − k1 t
(2b)
Qt = kH t
(3)
Qt = kKP t n Qeq
(4)
where, Qt is the drug released amount at time t, Q0 is the initial drug amount of the release medium, Qeq is the amount of drug released at equilibrium, k0 is the rate constant of the zero order model, k1 is the rate constant of the first order model, kH the rate constant of the Higuchi kinetic model, kKP is the rate constant of the Korsmeyer-Peppas kinetic model and t is the release time, n is the release exponent. The release parameters of ML hydrogels were listed in Table 2. The zero order model was found the best fit to the release data due to larger regression coefficient than that of the other models. For zero order kinetic model, the cumulative release of drug depended on only time and it was independent of drug concentration making it an ideal model. 3.2. Antibacterial performance assay Antibacterial performances of ampicillin loaded multilayer hydrogels were determined using agar disc diffusion assay. The results of antibacterial performances of ampicillin loaded and free multilayer hydrogels were given in Table 3. No inhibition zone was observed for
Fig. 5. The in-vitro drug release profile of ML-D hydrogels. 5
Journal of Drug Delivery Science and Technology 58 (2020) 101536
E. Tamahkar, et al.
Table 2 The release parameters of ML-D hydrogels. ML-D
Zero order model
First order model 2
k0
R
0.023
0.96
Higuchi model 2
k1
R
0.0039
0.88
free multilayer hydrogels and therefore, it has been proven that free multilayer hydrogels do not have any antibacterial activity alone. However, it is capable of releasing the antibacterial agent loaded. S. aureus strains were included in this study as Gram positive bacterial strains. Furthermore, antibacterial performances of ampicillin loaded multilayer hydrogels against both oxacillin resistant and sensitive S. aureus to evaluate the difference of inhibition zone formation. Oxacillin is one of the member of beta-lactam group antibiotics similarly ampicillin. It is well known that if the microorganism is resistant to one of the beta-lactam group antibiotics, it becomes also resistant to all other antibiotics included to this group. It was observed that ampicillin loaded multilayer hydrogels indicated excellent antibacterial performance against oxacillin sensitive S. aureus strains through markedly inhibition of bacterial growth. As can be seen from Fig. 6, growth of this bacterial strain decreased around ampicillin loaded multilayer hydrogels much the similar as the performance of commercial ampicillin disc. It is remarkable to mention the unique antibacterial performance of ML hydrogels against S. aureus which is one of the most common causative agents of wound infections. Dermal wounds could be easily colonized with S. aureus via cross-contamination in healthcare settings, especially in hospitalized patients. In this respect, ampicillin loaded multilayer hydrogels have been presented as functional antibacterial materials in order to block wound infections caused by beta-lactam sensitive bacteria. In addition, it should be emphasized that these hydrogels are introduced as promising materials for wound dressings with preserving their antibiotic releasing capabilities. On the other hand, recent studies have shown that multi-drug resistance has been increased in microorganisms. Unconscious antibiotic use and inappropriate prescribing are the major reasons given for developing resistance. The obtained results pointed out that, antibiotics which may be the potential treatment of choice for infections caused by multi-drug resistant bacterial strains, could also be loaded into the hyaluronic acid layer of multilayer hydrogels. As well as can be expected, ampicillin loaded multilayer hydrogels showed any antibacterial performance against oxacillin resistant S. aureus and ampicillin resistant E. coli. E. coli strains was included in this study as Gram negative bacterial strains and showed high ampicillin resistance in recent years.
Korsmeyer-Peppas model 2
kH
R
0.58
0.90
kKP
n
R2
0.035
0.41
0.91
Fig. 6. (1): Antimicrobial activities of ampicillin disc (10 μg/disc), ML-D and ML against oxacillin sensitive S. aureus, Line A: Inhibition zone diameter of ampicillin disc (10 μg/disc), Line B: Inhibition zone diameter of ML-D, (2): Antimicrobial activities of ampicillin disc (10 μg/disc), ML-D and ML against E. coli.
Fig. 7. L929 cells were incubated for 24 h, 48 h and 72 h on ML or ML-D hydrogels and the cells were assayed using the MTT assay.
was no cytotoxic effect of ML and ML-D hydrogels on L929 cells (Fig. 7). Cell viability was highest at 72 h incubation for ML group, on the other hand in ML-D group the cell viability was stable in all incubation times, indicating the antibiotic elution did not cause severe effect on fibroblast proliferation. Similar to our results, Onat et al. showed that, multilayer Tannic Acid-Ciprofloxacin containing and poly(N-vinyl caprolactam coassembled chitosan/poly(ethylene glycol) hydrogel materials increased the human fibroblast proliferation compared to bare hydrogels [33].
3.3. Cell viability assay 4. Conclusions In several studies dressing material itself induced cellular cytotoxicity [32]. After showing the bacterial inhibition of dressing composite developed, it is also necessary to evaluate that it is not toxic to fibroblast cells in order to support its applicative usage. According to MTT assay showing mitochondrial enzyme activity of alive cells, there
In this study, ML hydrogels having four layers of carboxylated PVA, gelatin, hyaluronic acid and gelatin prepared via layer-by-layer self assembly technique were developed as a novel antibiotic eluting wound dressing. ML hydrogels loaded with ampicillin showed antibiotic
Table 3 Antibacterial performances of multilayer hydrogels against test bacterial strains. Bacterial strains
oxacillin sensitive Staphylococcus aureus oxacillin resistant Staphylococcus aureus Escherichia coli
Zone of inhibition (mm) Ampicillin (10 μg/disc)
ML-D
ML
19 ± 1.0 -
14 ± 2.0 -
-
ML-D: ampicillin loaded multilayer hydrogels, ML: multilayer hydrogels without drug molecules. 6
Journal of Drug Delivery Science and Technology 58 (2020) 101536
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release for 7 days with a moderate burst release of 34.5% within 6 h. The ML hydrogels showed antibacterial activity against oxacillin sensitive S. aureus and showed no toxic effect on cultured fibroblasts, indicating that the novel wound dressings represent an effective option for selective treatment of bacterial infections.
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CRediT authorship contribution statement Emel Tamahkar: Project administration, Data curation, Writing original draft. Bengi Özkahraman: Project administration, Data curation, Writing - original draft. Aysun Kılıç Süloğlu: Investigation, Methodology, Writing - review & editing. Neslihan İdil: Investigation, Methodology, Writing - review & editing. Işık Perçin: Investigation, Methodology, Writing - review & editing. Declaration of competing interest The author declares that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest. References [1] A.C.d.O. Gonzalez, T.F. Costa, Z.d.A. Andrade, A.R.A.P. Medrado, Wound healing a literature review, An. Bras. Dermatol. 91 (2016) 614–620, https://doi.org/10. 1590/abd1806-4841.20164741. [2] E. Rezvani Ghomi, S. Khalili, S. Nouri Khorasani, R. Esmaeely Neisiany, S. Ramakrishna, Wound dressings: current advances and future directions, J. Appl. Polym. Sci. 136 (2019) 47738, https://doi.org/10.1002/app.47738. [3] D. Zhang, W. Zhou, B. Wei, X. Wang, R. Tang, J. Nie, J. Wang, Carboxyl-modified poly(vinyl alcohol)-crosslinked chitosan hydrogel films for potential wound dressing, Carbohydr. Polym. 125 (2015) 189–199, https://doi.org/10.1016/j.carbpol. 2015.02.034. [4] J. Tavakoli, S. Mirzaei, Y. Tang, Cost-effective double-layer hydrogel composites for wound dressing applications, Polymers 10 (2018), https://doi.org/10.3390/ polym10030305. [5] L. Ding, X. Shan, X. Zhao, H. Zha, X. Chen, J. Wang, C. Cai, X. Wang, G. Li, J. Hao, G. Yu, Spongy bilayer dressing composed of chitosan–Ag nanoparticles and chitosan–Bletilla striata polysaccharide for wound healing applications, Carbohydr. Polym. 157 (2017) 1538–1547, https://doi.org/10.1016/j.carbpol.2016.11.040. [6] S.G. Priya, A. Gupta, E. Jain, J. Sarkar, A. Damania, P.R. Jagdale, B.P. Chaudhari, K.C. Gupta, A. Kumar, Bilayer cryogel wound dressing and skin regeneration grafts for the treatment of acute skin wounds, ACS Appl. Mater. Interfaces 8 (2016) 15145–15159, https://doi.org/10.1021/acsami.6b04711. [7] Y. Guo, S. Pan, F. Jiang, E. Wang, L. Miinea, N. Marchant, M. Cakmak, Anisotropic swelling wound dressings with vertically aligned water absorptive particles, RSC Adv. 8 (2018) 8173–8180, https://doi.org/10.1039/C7RA13764H. [8] M. Shemesh, M. Zilberman, Structure–property effects of novel bioresorbable hybrid structures with controlled release of analgesic drugs for wound healing applications, Acta Biomater. 10 (2014) 1380–1391, https://doi.org/10.1016/j.actbio. 2013.11.025. [9] G. Liu, Z. Ding, Q. Yuan, H. Xie, Z. Gu, Multi-layered hydrogels for biomedical applications, Front. Chem. 6 (2018), https://doi.org/10.3389/fchem.2018.00439 439-439. [10] P.T. Hammond, Engineering materials layer-by-layer: challenges and opportunities in multilayer assembly, AIChE J. 57 (2011) 2928–2940, https://doi.org/10.1002/ aic.12769. [11] F. Reyes-Ortega, A. Cifuentes, G. Rodríguez, M.R. Aguilar, Á. González-Gómez, R. Solis, N. García-Honduvilla, J. Buján, J. García-Sanmartin, A. Martínez, J.S. Román, Bioactive bilayered dressing for compromised epidermal tissue regeneration with sequential activity of complementary agents, Acta Biomater. 23 (2015) 103–115, https://doi.org/10.1016/j.actbio.2015.05.012. [12] S.-F. Ng, S.-L. Tan, Development and in vitro assessment of alginate bilayer films containing the olive compound hydroxytyrosol as an alternative for topical chemotherapy, Int. J. Pharm. 495 (2015) 798–806, https://doi.org/10.1016/j.ijpharm. 2015.09.057. [13] J. Ding, R. He, G. Zhou, C. Tang, C. Yin, Multilayered mucoadhesive hydrogel films based on thiolated hyaluronic acid and polyvinylalcohol for insulin delivery, Acta
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Contents lists available at ScienceDirect
Journal of Drug Delivery Science and Technology journal homepage: www.elsevier.com/locate/jddst
A novel multilayer hydrogel wound dressing for antibiotic release Emel Tamahkar
a,b,∗,1
, Bengi Özkahraman
, Aysun Kılıç Süloğlu , Neslihan İdil , Işık Perçin
c,∗∗,1
d
d
T d
a
Hitit University, Department of Chemical Engineering, Çorum, 19030, Turkey Balıkesir University, Department of Food Engineering, Balıkesir, 10145, Turkey Hitit University, Department of Polymer Engineering, Çorum, 19030, Turkey d Hacettepe University, Department of Biology, Ankara, 06800, Turkey b c
A R T I C LE I N FO
A B S T R A C T
Keywords: Multilayer hydrogels Wound dressing Antibiotic release Antibacterial activity
In this study, the novel multilayer (ML) hydrogels were developed as antibacterial wound dressings. ML hydrogels were prepared as four layers using carboxylated polyvinyl alcohol (PVA-C), gelatin (G), hyaluronic acid (HA) and gelatin respectively. The upper layers (PVA-C and G) provide the moist control and physical barrier for microorganisms. The HA based middle layer was designed as an antibiotic-loaded layer. The lower layer serves as the controlling membrane for antibiotic release and provides the removal of excess exudate from the wound site. The ML hydrogels were characterized with FT-IR, SEM, DSC, swelling tests and hydrolytic degradation tests. Cell viability assay was also performed in L929 fibroblast cells in vitro. The in-vitro drug release profile of ML hydrogels was investigated at pH 7.4 at 37 °C and it was found that 63% of the antibiotic was released after 7 days.
1. Introduction
adhesion onto the wound surface and support for new tissue formation [6,7]. Multilayer hydrogels consisting a drug-loaded layer provide controlled drug release for long time providing mass transfer limitations for drug molecules throughout the polymeric matrix [8]. Layer-by-layer (LBL) approach can be applied to broad range of polymers enabling the design of functional biomaterials with desired properties without any additional chemicals. The preparation of multilayer hydrogels via LBL technique depends on the adsorption of the polymers having oppositely charged groups layer by layer providing the incorporation of them into multilayers [9]. Multilayer hydrogels fabricated by LBL self assembly method comprised of different polymeric layers present great attention for drug delivery with their facile preparation procedure, tunable morphological features and high biocompatibility providing the drug molecules released from the matrix in a controlled manner [10]. ReyesOrtega et al. produced a bilayer wound dressing based on gelatin-HA that consisted on a first layer of cross-linked gelatin-HA loaded with proadrenomedullin N-terminal 20 peptide and a second layer of polyurethane derivative that was loaded with nanoparticles of bemiparin. The bilayer hydrogels were investigated using in-vivo mouse models. The results showed that the bilayer wound dressing presents promising alternative since they reduced the contraction and the inflammation with the aid of the controlled release of the two drugs [11]. Ng and Tan synthesized a bilayer film based on alginate and gelatin containing
Wound healing including the phases of hemostasis, inflammation, proliferation and remodelling is a complex and long process since the remodelling proceeds for 21 days - 1 year forming suitable conditions for bacterial growth [1]. Selection of wound dressing that depends on many factors such as type and depth of wound, amount of exudate etc. becomes significant for an effective treatment. An ideal wound dressing should provide moist environment that facilitates healing process [2]. It should allow adequate gas permeability and remove excess exudate from the wound surface while keeping it moist. Also, it should serve as a barrier for microorganisms providing an antibacterial medium. Furthermore it should be biocompatible, biodegradable, cheap, easy-toapply without requiring frequent changing [3]. Recently, multilayer wound dressings enabling better healing process than single layer dressings have become promising alternatives [4]. Multilayer hydrogels consisting of more than two layers are designed to meet the requirements of wound healing by the choice of chemical and physical properties of the constituent biomaterials. Multilayer hydrogel wound dressings integrates the advantages of each constituents. The upper layer serves as a barrier for bacterial transition and as a controlling membrane for moist environment [5]. The lower layer with spongy structure is utilized for the absorption of the excess exudate, ∗
Corresponding author. Hitit University, Department of Chemical Engineering, Çorum, 19030, Turkey. Corresponding author. E-mail addresses: [email protected] (E. Tamahkar), [email protected] (B. Özkahraman). 1 These authors contributed equally to this work. ∗∗
https://doi.org/10.1016/j.jddst.2020.101536 Received 13 December 2019; Received in revised form 9 January 2020; Accepted 21 January 2020 Available online 23 January 2020 1773-2247/ © 2020 Published by Elsevier B.V.
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25 °C, and precipitated with acetone. The product was dried for 24 h under vacuum at 25 °C.
hydroxytyrosol for topical chemotherapy. Glycerol and poly propylene glycol were used as plasticizers. The drug-free bilayer films were prepared by changing plasticizers’ ratio and they were characterized for their such as physical, mechanical and rheological properties. The invitro drug release results showed that bilayer films had lower drug release rate than single films [12]. Ding et al. developed a multilayer hydrogel film using hyaluronic acid and PVA for insulin release. It was determined that 90% of insulin was released from the resultant multilayer hydrogel films that were produced via vacuum-drying after 12 h since the release of 90% of insulin was reached after 6 h with the prepared multilayer hydrogel films synthesized by lyophilization [13]. Although there exist significant improvements of the fabrication of the wound dressings, the development of novel antibacterial multilayer wound dressings with transparency, cost-efficiency, high healing capacity is still of great interest [14]. Infection is one of the major factors, which delays the healing process. Topical administration of an antibiotic drug at the wound site is one of the most important aspects for the prevention of infection during healing process. Controlled drug release present great advantages when applied to wound dressings providing delivery of the drug molecules through the wound site in a controlled manner for long time periods without need for the frequent replacement of the dressing material [15]. Also the utilization of drug-loaded wound dressings enable delivering the antibiotics in a determined dosage avoiding high systemic dose [16]. There are many reports about drug-loaded wound dressings with sustained antibiotic release [17,18]. The aim of this study was to develop multilayer hydrogel wound dressings composed of natural polymers with water-based approach for effective wound healing. The multilayer hydrogel film was composed of 4 layers of which PVA-based first layer, gelatin-based second layer, ampicillin loaded HA-based third layer and gelatin-based fourth layer. For the best knowledge of the authors, this is the first report concerning the development of multilayer hydrogel films consisting of PVA-C, gelatin and hyaluronic acid as a new system for wound healing and drug delivery. The drug-free and drug-loaded multilayer films were synthesized via solvent evaporation and characterized by FT-IR, SEM and DSC measurements. Also the swelling degree and hydrolytic degradation ratio of the hydrogels was investigated. The results of in-vitro drug release tests showed that 63% of the loaded drug was released from the multilayer hydrogels after 5 days. On the other hand, antimicrobial performance of novel polymeric ampicillin loaded ML hydrogels were investigated against both oxacillin resistant and sensitive Staphylococcus aureus (S. aureus) and Escherichia coli (E. coli). Previous studies have shown that dressing material itself can induce toxicity on cells due to monomers or antimicrobials used. Therefore it is necessary to show that the novel dressing material is not toxic to the cells which would be included in healing process [19]. Thus, the in-vitro cytotoxicity of the multilayer hydrogels was also investigated.
2.3. Synthesis of multilayer hydrogel film A multilayer hydrogel film consisted of four layers. For first layer, PVA-C hydrogel film was prepared using the solvent-casting process. Polymer solution of PVA-C (10 wt %) was stirred under mechanical agitation for 3 h at 80 °C and casted into Teflon petri dishes. The first layer of wound healing films was crosslinked at 40 °C for 48 h. To fabricate second layer, gelatin solution (5 wt %) was prepared by dissolving at 50 °C and then casted on the first layer. The second layer was left at 25 °C for 3 h and subsequently dried under vacuum at 40 °C for 48 h. For preparation of the third layer, HA solution (5 wt %) was poured onto the existent two layer and then was left at room temperature for 3 h with subsequent vacuum drying at 40 °C for 48 h. Finally, the last layer was prepared by dissolving gelatin solution (5 wt %) at 50 °C and then it was casted on resultant three layered hydrogels. Ampicillin-loaded multilayer hydrogels (ML-D) was prepared in the same way as blank multilayer hydrogels (ML) by adding ampicillin (2 wt %) directly to the HA solution. The codes of the multilayer hydrogels prepared in this study were listed in Table 1.
2.4. Characterization studies The hydrogels were characterized by a FTIR spectrophotometer (Agilent Technologies Inc., USA). The morphological characterization of multi-layer hydrogels was taken by a FEQ Scanning electron microscope (SEM). Differential scanning calorimetry (DSC) was utilized for thermal analysis of the multilayer hydrogels. The heating profile was established from 25 °C to 300 °C at a rate of 5 °C/min.
2.5. Swelling studies The swelling study of the multilayer wound healing hydrogels was determined gravimetrically using different buffer solutions (pH 5.5 and pH 7.4) during 24 h. The multilayer hydrogel was immersed in these solutions at 37 °C. The swollen multilayer hydrogel was weighed at regular time intervals until equilibrium was reached after the excess water was wiped off from the hydrogel film surface. The swelling ratio (SR %) was calculated from the expression (1):
SR% = (Ws − Wd/ Wd ) ∗ 100
(1)
where Ws and Wd are the weights of the swollen hydrogels at equilibrium and dry hydrogels, respectively.
2. Experimental
2.6. In-vitro hydrolytic degradation
2.1. Chemicals
In vitro hydrolytic degradability of the multilayer hydrogel (W0) was investigated using PBS at 37 °C to provide a similar environment with wound exudate. At determined time intervals, the remaining hydrogel was taken out from the solution and dried for 24 h (Wt). The weight loss percentage was calculated with the equation (2) as follows:
Polyvinyl alcohol (PVA) (Mw: 30000-70000), ampicillin sodium salt, 4-dimethylaminopyridine (DMAP) and phosphate buffer tablet were obtained from Sigma-Aldrich (St. Louis, MO). Gelatin (G, type A) was obtained Biomatik (Wilmington, Delaware, USA). Sodium hyaluronate was purchased from Across Organic (Geel, Belgium) products. All other reagents were of analytical grade.
Weight loss (%) = (Wo − Wt / Wt ) ∗ 100
(2)
Table 1 The codes of the multi-layer hydrogels.
2.2. Synthesis of carboxylated PVA PVA (10 g) powder was dissolved in deionized water (100 mL) at 90 °C. DMAP (2.77 g) was mixed with the solution and stirred for 1 h at 65 °C. Succinic anhydride (23 g) was mixed with the reaction medium and stirred for 24 h at 65 °C. The PVA-COOH solution was cooled to 2
Polymer code
First layer
Second layer
Third layer
Fourth layer
ML ML-D
PVA-C PVA-C
G G
HA HA-Amp
G G
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2.7. Drug release studies The in vitro drug release profile of ML-D hydrogel was evaluated in pH 7.4 at 37 °C for one week. At scheduled time intervals, 2 mL was taken from the release medium and replaced with same volume of fresh buffer solutions. The amount of released drug from the drug-loaded hydrogels was calculated with aid of calibration curves plotted using ampicillin solutions (pH 7.4). The amount of ampicillin release was determined utilizing T80+ UV/VIS Spectrometer (PG Instruments Ltd.) at 275 nm.
2.8. Antibacterial performances of multilayer hydrogels The antibacterial performances of multilayer hydrogels were examined against both oxacillin resistant and sensitive S. aureus and E. coli. The identification of these bacterial strains and the determination of their antibacterial susceptibilities were performed via automated system (Vitek 2 System, bioMérieux, USA). Antibacterial performances of multilayer hydrogels were investigated by agar disc diffusion assay. In the first step, 100 mL of Luria Bertani (LB) broth in a 250 mL Erlenmayer flask was prepared. Then, bacterial strains were inoculated into LB broth and incubated at 37 °C for 18 h to obtain fresh bacterial cultures. In the second step, bacterial culture suspensions at the exponential growth phase were adjusted according to 0.5 McFarland (1.5×108 CFU/mL). In the third step, 100 μL of prepared culture suspensions were inoculated onto LB agar plates. Ampicillin loaded-multilayer hydrogels, ampicillin free-multilayer hydrogel (negative control) and ampicillin (10 μg/disc, positive control) were placed onto the LB agar plates which were cultivated at 37 °C for 18 h. The antimicrobial performances of multilayer hydrogels were evaluated by measuring the diameters of the inhibition zones in mm. The antibacterial performance assays were done in triplicate.
2.9. Cell viability assay The cell viability properties of ML and ML-D hydrogels were investigated. L929 mouse fibroblast cells were obtained from the American Type Culture Collection. Cells were cultured in DMEM (Dulbecco's modified Eagle's medium) supplemented with L-glutamine, 10% (v/v) fetal bovine serum (FBS) (Cegrogen biotech, Germany) and 1% penicillin-streptomycin (Gibco, Life Technologies, USA) solution. Incubation was carried out in an oven with humidified atmosphere of 5% CO2, at 37 °C. L929 cell line (Passage 17) was studied. The medium was changed every two days. Cells were trypsinized when they reached 90% confluence. All hydrogel samples were left to sterilize at UV light in the cabin for 45 min. The parafilms which prevent cells to attach surface were cut for the size of 48 well plate and were waited for 20 min at 70% alcohol. Then the parafilms were placed into sterile distilled water and engrafted on well plate and lefted at UV light for 45 min. Before seeding of the cells, hydrogels were conditioned on DMEM medium with serum for 30 min. Cells were cultivated on two 75 cm2 flasks and seeded at two different stages on hydrogel samples. Cells were trypsinized and santrifuged at 800 rpm for 5 min. Pellet was divided into fifteen for 24, 48 and 72 h. For each well, the number of cells was 3 × 103 per well. MTT [3-(4,5-dimetiltiazol-2-il)-2,5-difeniltetrazolium bromid] solution (MTT, Sigma, St. Louis, MO, USA) was added to wells. Cells were lefted for 3 h incubation. After incubation, medium was aspired. DMSO/NH3 solution was prepared (5% NH3-95% DMSO), 200 μl per well was added to samples. Samples were transferred to non-sterile 96 well plate. Plates were lefted on shaker for 10–15 min. The OD value was measured using an ELISA plate reader of absorbance at 550 nm (Bio Tek Instruments Inc., Winooski, VT, USA). Same procedure was repeated for 48 and 72 h.
Fig. 1. FT-IR spectrums of PVA, PVA-C, gelatin, hyaluronic acid and multilayer hydrogel (A) and ampicillin, free ML hydrogel and ampicillin-loaded ML hydrogel (ML-D) (B).
3. Results and discussion In this study, ML hydrogels were prepared via layer-by-layer self assembly through electrostatic interactions between polymeric layers. We supposed that positively charged gelatin that has isoelectric point at around 7-9 would fabricate a stable layer when poured onto the negatively charged carboxylated PVA layer. Hyaluronic acid layer which has a negative charge was added to form an antibiotic-loaded layer as a drug carrier reservoir. Lastly, gelatin layer was introduced onto the resultant three layers. 3.1. Characterization studies Fig. 1A shows ATR-FTIR spectra of carboxylated PVA, HA, gelatin and ML hydrogel. The carboxylation of PVA was verified by FTIR measurements. The sharp new peak at around 1700 cm−1 that was assigned to carbonyl stretching mode of esters was observed after the esterification. The main characteristic bands of gelatin were 3
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Fig. 2. The cross-sectional SEM images (A) and optic photograph of multilayer hydrogels (B).
demonstrated at 1628, 1526, 1238 cm−1 corresponding to amide I, II and III respectively. The spectra of the ML hydrogel were characterized due to the amide bands at 1630, 1530 and 1239 cm−1. The incorporation of HA into the ML hydrogel structure was verified by the new peak appeared at 1039 cm−1 which attributes to C-O-C stretching of hyaluronic acid. Also the strength of the band at 3275 cm−1 of ML hydrogels assigned to hydroxyl group decreased with respect to the other constituents indicating the chemical interactions between the polymeric layers. The ATR-FTIR spectrum of ML-D hydrogels was displayed in Fig. 1B. The loading of ampicillin molecules into the HA layer of ML-D hydrogels was determined due to new peak observed at 2087 cm−1 referred to bending of S-C of ampicillin. Also the strength of the amide bands (1631, 1526 and 1238 cm−1) and the hydroxyl band (3275 cm−1) decreased with the addition of drug molecules throughout the polymeric matrix. The SEM images of multilayer hydrogels corresponding to crosssections performed through the thickness of the hydrogels were presented in Fig. 2. The images indicate the multilayered structure consisting of four layers fabricated with no obvious distinction between layers [20]. The upper layers were involved of carboxylated-PVA and gelatin; the drug-loaded layer was located at the middle; the lower layer was composed of gelatin. The upper layers were formed to control the water vapor since the lower layer that adheres to the wound surface was designed to serve as the controlling layer of the drug release process [21]. The drug-loaded layer that was composed of HA was determined to have macroporous structure with high interconnectivity [22]. This network structure could facilitate the swelling of the hydrogels. Fig. 2B shows the transparency of ML hydrogels. DSC thermograms of ML hydrogels and the physical blend of the used polymers (ML-Blend) were shown in Fig. 3. Tg values of ML and ML-Blend were found as 59.29 °C and 49.93 °C respectively. The first endothermic peaks of ML and ML-Blend were observed at 63.7 °C and 53 °C respectively. The second endothermic peaks of ML and ML-Blend were found at 147 °C and 92.5 °C respectively. It was determined that the endothermic peaks of the polymer blend were shifted to the higher temperatures with the preparation of multilayer hydrogels implying the change in the chemical structure of the polymers. These results offer the
Fig. 3. DSC thermograms of ML hydrogels and the physical blend of the used polymers.
fabrication of a more stable structure with respect to the polymerpolymer interactions between the layers of the multilayer hydrogels [23]. The in-vitro swelling tests were evaluated at pH 5.5 and pH 7.4. The equilibrium swelling ratios of ML hydrogels at pH 5.5 and 7.4 were calculated gravimetrically and found as 517.8% and 339.4% respectively. The mechanism of the swelling of the ML hydrogels depends on the pH of the medium and thus the ionization state of the carboxylic acid and amino groups that are located at the side chains of the polymers. When the pH of the medium increases above the pKa of side groups of the polymeric structure, the hydrogels become swollen due to the ionic repulsion. Therefore, the lower swelling degree observed at pH 7.4 may be explained due to the lack of the ionization state of the 4
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D hydrogels was occurred within the first 6 h. Then the release rate decreases exponentially during 3 days. Lastly, the release rate became constant until 7 days. The release profile of ampicillin loaded ML hydrogels presented a burst release to prevent the risk of the bacterial infection, which can be occurred during the initial injury [26]. The fast release rate of the drug molecules within the first 6 h can be due to their location in the polymeric structure since the slower release rate of ampicillin can be attributed to the need of time to reach the diffusion barriers. The controlled drug release profile was due to the presence of the extra layers acting as the controlling membranes for drug diffusion [27]. Also, the interaction between negatively charged COO− and positively charged NH3+, Na+, H+ could reduce the release rate. The wound dressing with prolonged antibiotic release could be used for longer time periods thus resolving the need for repeated dressing changes. The prepared wound dressing in this study exhibited significant potential by showing antibiotic release in a controlled manner for prolong duration of time eliminating the requirement of continual dressing change. Kamoun et al. developed polyvinyl alcohol-alginate hydrogels for ampicillin delivery. The hydrogels showed a burst release of 38–45% within 15 min and the total release of ampicillin was obtained after 6 h [28]. Kenawy et al. prepared polyvinyl alcohol-hydroxyethyl starch (HES) hydrogels with freeze thawing method for controlled release of ampicillin. The release amount of ampicillin was increased with the increasing amount of HES and the release equilibrium was attained after 6 h [29]. Bako et al. synthesized nanogel with poly-gamma-glutamic acid nanoparticles via photopolymerization. The loaded ampicillin into the resultant nanogels was released during 24 h with a burst release in the first 4 [30]. Poonguzhali et al. prepared alginate based film with the incorporation of nanocellulose for ampicillin release. The release of ampicillin molecules from nanocomposite film obtained equilibrium after 500 min [31]. All these literature for ampicillin release show the production of biomaterials with fast release kinetics. MLD hydrogels having four layer enable the release of ampicillin for prolonged time periods. In order to describe the release profile of ampicillin from ML-D hydrogels, zero order, first order, Higuchi model and Korsmeyer-Peppas model were applied to release data. The mathematical expressions of zero order model (1a), first order model (2b), Higuchi model (3) and Korsmeyer-Peppas model (4) are described as follows:
Fig. 4. The in-vitro hydrolytic degradation of ML hydrogels.
side groups. The in-vitro swelling degree is one of the major properties of the wound dressings indicating the capability of the absorption of the excess exudate form the wound site which, causes possible bacterial infection. It has been reported that the typical wound (10 cm2) produces 5 mL of exudate in 24 h [24]. The ML hydrogels prepared in this study have swelling degree of 5.17 g/g (pH 5.5) and 3.39 g/g (pH 7.4) implying the suitability of ML hydrogels as a wound dressing. The in-vitro hydrolytic degradation of the ML hydrogels was tested by quantifying the mass loss after incubating the dressings in PBS at 37 °C. The hydrolytic degradation ratios were shown in Fig. 4. It was shown that the mass loss increased with increasing time and the degradation was completed after 15 days exhibiting a long-term degradation profile with good stability thus makes the prepared hydrogels a good alternative as wound dressings. The in-vitro drug release from ampicillin loaded multilayer hydrogels was shown in Fig. 5. Almost 65% of ampicillin was released within 7 days indicates that a long-term release profile could be achieved using the prepared multilayer hydrogels. The mechanism of the drug release from ML-D hydrogels was occurred such as hydration of the polymeric structure, swelling of the hydrogel and diffusion of the drug molecules throughout the swollen matrix [25]. The release profile can be separated into three main phases. The initial burst release (34.5%) from ML-
Qt = Q0 + k 0 t
(1a)
lnQt = lnQ0 − k1 t
(2b)
Qt = kH t
(3)
Qt = kKP t n Qeq
(4)
where, Qt is the drug released amount at time t, Q0 is the initial drug amount of the release medium, Qeq is the amount of drug released at equilibrium, k0 is the rate constant of the zero order model, k1 is the rate constant of the first order model, kH the rate constant of the Higuchi kinetic model, kKP is the rate constant of the Korsmeyer-Peppas kinetic model and t is the release time, n is the release exponent. The release parameters of ML hydrogels were listed in Table 2. The zero order model was found the best fit to the release data due to larger regression coefficient than that of the other models. For zero order kinetic model, the cumulative release of drug depended on only time and it was independent of drug concentration making it an ideal model. 3.2. Antibacterial performance assay Antibacterial performances of ampicillin loaded multilayer hydrogels were determined using agar disc diffusion assay. The results of antibacterial performances of ampicillin loaded and free multilayer hydrogels were given in Table 3. No inhibition zone was observed for
Fig. 5. The in-vitro drug release profile of ML-D hydrogels. 5
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Table 2 The release parameters of ML-D hydrogels. ML-D
Zero order model
First order model 2
k0
R
0.023
0.96
Higuchi model 2
k1
R
0.0039
0.88
free multilayer hydrogels and therefore, it has been proven that free multilayer hydrogels do not have any antibacterial activity alone. However, it is capable of releasing the antibacterial agent loaded. S. aureus strains were included in this study as Gram positive bacterial strains. Furthermore, antibacterial performances of ampicillin loaded multilayer hydrogels against both oxacillin resistant and sensitive S. aureus to evaluate the difference of inhibition zone formation. Oxacillin is one of the member of beta-lactam group antibiotics similarly ampicillin. It is well known that if the microorganism is resistant to one of the beta-lactam group antibiotics, it becomes also resistant to all other antibiotics included to this group. It was observed that ampicillin loaded multilayer hydrogels indicated excellent antibacterial performance against oxacillin sensitive S. aureus strains through markedly inhibition of bacterial growth. As can be seen from Fig. 6, growth of this bacterial strain decreased around ampicillin loaded multilayer hydrogels much the similar as the performance of commercial ampicillin disc. It is remarkable to mention the unique antibacterial performance of ML hydrogels against S. aureus which is one of the most common causative agents of wound infections. Dermal wounds could be easily colonized with S. aureus via cross-contamination in healthcare settings, especially in hospitalized patients. In this respect, ampicillin loaded multilayer hydrogels have been presented as functional antibacterial materials in order to block wound infections caused by beta-lactam sensitive bacteria. In addition, it should be emphasized that these hydrogels are introduced as promising materials for wound dressings with preserving their antibiotic releasing capabilities. On the other hand, recent studies have shown that multi-drug resistance has been increased in microorganisms. Unconscious antibiotic use and inappropriate prescribing are the major reasons given for developing resistance. The obtained results pointed out that, antibiotics which may be the potential treatment of choice for infections caused by multi-drug resistant bacterial strains, could also be loaded into the hyaluronic acid layer of multilayer hydrogels. As well as can be expected, ampicillin loaded multilayer hydrogels showed any antibacterial performance against oxacillin resistant S. aureus and ampicillin resistant E. coli. E. coli strains was included in this study as Gram negative bacterial strains and showed high ampicillin resistance in recent years.
Korsmeyer-Peppas model 2
kH
R
0.58
0.90
kKP
n
R2
0.035
0.41
0.91
Fig. 6. (1): Antimicrobial activities of ampicillin disc (10 μg/disc), ML-D and ML against oxacillin sensitive S. aureus, Line A: Inhibition zone diameter of ampicillin disc (10 μg/disc), Line B: Inhibition zone diameter of ML-D, (2): Antimicrobial activities of ampicillin disc (10 μg/disc), ML-D and ML against E. coli.
Fig. 7. L929 cells were incubated for 24 h, 48 h and 72 h on ML or ML-D hydrogels and the cells were assayed using the MTT assay.
was no cytotoxic effect of ML and ML-D hydrogels on L929 cells (Fig. 7). Cell viability was highest at 72 h incubation for ML group, on the other hand in ML-D group the cell viability was stable in all incubation times, indicating the antibiotic elution did not cause severe effect on fibroblast proliferation. Similar to our results, Onat et al. showed that, multilayer Tannic Acid-Ciprofloxacin containing and poly(N-vinyl caprolactam coassembled chitosan/poly(ethylene glycol) hydrogel materials increased the human fibroblast proliferation compared to bare hydrogels [33].
3.3. Cell viability assay 4. Conclusions In several studies dressing material itself induced cellular cytotoxicity [32]. After showing the bacterial inhibition of dressing composite developed, it is also necessary to evaluate that it is not toxic to fibroblast cells in order to support its applicative usage. According to MTT assay showing mitochondrial enzyme activity of alive cells, there
In this study, ML hydrogels having four layers of carboxylated PVA, gelatin, hyaluronic acid and gelatin prepared via layer-by-layer self assembly technique were developed as a novel antibiotic eluting wound dressing. ML hydrogels loaded with ampicillin showed antibiotic
Table 3 Antibacterial performances of multilayer hydrogels against test bacterial strains. Bacterial strains
oxacillin sensitive Staphylococcus aureus oxacillin resistant Staphylococcus aureus Escherichia coli
Zone of inhibition (mm) Ampicillin (10 μg/disc)
ML-D
ML
19 ± 1.0 -
14 ± 2.0 -
-
ML-D: ampicillin loaded multilayer hydrogels, ML: multilayer hydrogels without drug molecules. 6
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release for 7 days with a moderate burst release of 34.5% within 6 h. The ML hydrogels showed antibacterial activity against oxacillin sensitive S. aureus and showed no toxic effect on cultured fibroblasts, indicating that the novel wound dressings represent an effective option for selective treatment of bacterial infections.
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CRediT authorship contribution statement Emel Tamahkar: Project administration, Data curation, Writing original draft. Bengi Özkahraman: Project administration, Data curation, Writing - original draft. Aysun Kılıç Süloğlu: Investigation, Methodology, Writing - review & editing. Neslihan İdil: Investigation, Methodology, Writing - review & editing. Işık Perçin: Investigation, Methodology, Writing - review & editing. Declaration of competing interest The author declares that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest. References [1] A.C.d.O. Gonzalez, T.F. Costa, Z.d.A. Andrade, A.R.A.P. Medrado, Wound healing a literature review, An. Bras. Dermatol. 91 (2016) 614–620, https://doi.org/10. 1590/abd1806-4841.20164741. [2] E. Rezvani Ghomi, S. Khalili, S. Nouri Khorasani, R. Esmaeely Neisiany, S. Ramakrishna, Wound dressings: current advances and future directions, J. Appl. Polym. Sci. 136 (2019) 47738, https://doi.org/10.1002/app.47738. [3] D. Zhang, W. Zhou, B. Wei, X. Wang, R. Tang, J. Nie, J. Wang, Carboxyl-modified poly(vinyl alcohol)-crosslinked chitosan hydrogel films for potential wound dressing, Carbohydr. Polym. 125 (2015) 189–199, https://doi.org/10.1016/j.carbpol. 2015.02.034. [4] J. Tavakoli, S. Mirzaei, Y. Tang, Cost-effective double-layer hydrogel composites for wound dressing applications, Polymers 10 (2018), https://doi.org/10.3390/ polym10030305. [5] L. Ding, X. Shan, X. Zhao, H. Zha, X. Chen, J. Wang, C. Cai, X. Wang, G. Li, J. Hao, G. Yu, Spongy bilayer dressing composed of chitosan–Ag nanoparticles and chitosan–Bletilla striata polysaccharide for wound healing applications, Carbohydr. Polym. 157 (2017) 1538–1547, https://doi.org/10.1016/j.carbpol.2016.11.040. [6] S.G. Priya, A. Gupta, E. Jain, J. Sarkar, A. Damania, P.R. Jagdale, B.P. Chaudhari, K.C. Gupta, A. Kumar, Bilayer cryogel wound dressing and skin regeneration grafts for the treatment of acute skin wounds, ACS Appl. Mater. Interfaces 8 (2016) 15145–15159, https://doi.org/10.1021/acsami.6b04711. [7] Y. Guo, S. Pan, F. Jiang, E. Wang, L. Miinea, N. Marchant, M. Cakmak, Anisotropic swelling wound dressings with vertically aligned water absorptive particles, RSC Adv. 8 (2018) 8173–8180, https://doi.org/10.1039/C7RA13764H. [8] M. Shemesh, M. Zilberman, Structure–property effects of novel bioresorbable hybrid structures with controlled release of analgesic drugs for wound healing applications, Acta Biomater. 10 (2014) 1380–1391, https://doi.org/10.1016/j.actbio. 2013.11.025. [9] G. Liu, Z. Ding, Q. Yuan, H. Xie, Z. Gu, Multi-layered hydrogels for biomedical applications, Front. Chem. 6 (2018), https://doi.org/10.3389/fchem.2018.00439 439-439. [10] P.T. Hammond, Engineering materials layer-by-layer: challenges and opportunities in multilayer assembly, AIChE J. 57 (2011) 2928–2940, https://doi.org/10.1002/ aic.12769. [11] F. Reyes-Ortega, A. Cifuentes, G. Rodríguez, M.R. Aguilar, Á. González-Gómez, R. Solis, N. García-Honduvilla, J. Buján, J. García-Sanmartin, A. Martínez, J.S. Román, Bioactive bilayered dressing for compromised epidermal tissue regeneration with sequential activity of complementary agents, Acta Biomater. 23 (2015) 103–115, https://doi.org/10.1016/j.actbio.2015.05.012. [12] S.-F. Ng, S.-L. Tan, Development and in vitro assessment of alginate bilayer films containing the olive compound hydroxytyrosol as an alternative for topical chemotherapy, Int. J. Pharm. 495 (2015) 798–806, https://doi.org/10.1016/j.ijpharm. 2015.09.057. [13] J. Ding, R. He, G. Zhou, C. Tang, C. Yin, Multilayered mucoadhesive hydrogel films based on thiolated hyaluronic acid and polyvinylalcohol for insulin delivery, Acta
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