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BIOMAC 5190 1–6

International Journal of Biological Macromolecules xxx (2015) xxx–xxx

Contents lists available at ScienceDirect

International Journal of Biological Macromolecules journal homepage: www.elsevier.com/locate/ijbiomac

Injectable biopolymer based hydrogels for drug delivery applications

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Sadia Atta a , Shaista Khaliq a,b , Atif Islam b,∗ , Irtaza Javeria b,d , Tahir Jamil b , Muhammad Makshoof Athar a , Muhammad Imtiaz Shafiq c , Abdul Ghaffar d a

Institute of Chemistry, University of the Punjab, Lahore, Pakistan Department of Polymer Engineering and Technology, University of the Punjab, Quaid-e-Azam campus, Lahore, Pakistan Institute of Biochemistry and Biotechnology, University of the Punjab, Lahore, Pakistan d Department of Chemistry, University of Engineering and Technology, Lahore, Pakistan b c

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Article history: Received 5 May 2015 Received in revised form 22 June 2015 Accepted 24 June 2015 Available online xxx

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Keywords: Injectable Biopolymer Hydrogels Drug delivery

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1. Introduction

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Biopolymer based pH-sensitive hydrogels were prepared using chitosan (CS) with polyethylene glycol (PEG) of different molecular weights in the presence of silane crosslinker. The incorporated components remain undissolved in different swelling media as they are connected by siloxane linkage which was confirmed by Fourier transform infrared spectroscopy. The swelling in water was enhanced by the addition of higher molecular weight PEG. The swelling behaviour of the hydrogels against pH showed high swelling in acidic and basic pH, whereas, low swelling is examined at pH 6 and 7. This characteristic pH responsive behaviour at neutral pH made them suitable for injectable controlled drug delivery. The controlled release analysis of Cefixime (CFX) (model drug) loaded CS/PEG hydrogel exhibited that the entire drug was released in 30 min in simulated gastric fluid (SGF) while in simulated intestinal fluid (SIF), 85% of drug was released in controlled manner within 80 min. This inferred that the developed hydrogels can be an attractive biomaterial for injectable drug delivery with physiological pH and other biomedical applications. © 2015 Published by Elsevier B.V.

Recently, much attention has been given to utilize biopolymer based hydrogels in many applications including medicine, cosmetics, agriculture and biotechnology. Polymer hydrogels have received significant attention for years. Polymer hydrogels usually consist of stretchy cross-linked networks and interstitial spaces (fluid-filled) of network. They can modify their shape and volume reversibly depending on various environmental stimuli i.e. solvent composition, temperature, ionic strength, pH, electric and magnetic fields [1,2]. Hydrogels are formed by chemical or physical crosslinking of polymers. Consequently, they propose the possibility of several advanced functional polymers [3]. The properties of hydrogels, for example, hydrophilicity, superabsorbency, low interfacial tension, expandability, softness and selective permeability can be used in various biomedical and other applications [4–7]. Hydrogels have been studied widely for tissue engineering, drug delivery [8], cell encapsulation [9], etc. because of their excellent biocompatibility [10], biodegradability, no toxicity etc. [3,11].

∗ Corresponding author. Tel.: +92 300 6686 506. E-mail address: dratifi[email protected] (A. Islam).

Moreover, they have received notable attention as they cause minimal thrombosis, inflammatory responses and tissue damage [10]. Chitosan is a biocompatible [12], biodegradable [13], non-toxic, cationic polymer which is obtained by partial deacetylation of naturally occurring chitin [14,15]. Chitin is known to be the second most abundant natural polymer after cellulose in nature. Chitin is the key component of exoskeleton of crustaceans [16]. It is composed of N-acetylglucosamine repeating units with 1,4 linkage. The reactivity, processability and solubility of chitin have been enhanced in the form of chitosan by chemical modification (deacetylation). Chitosan is found to be readily processable into membranes and films than chitin [17]. It consists of N-acetyl-dglucosamine and d-glucosamine units with 1,4 linkage. Because of remarkable properties, chitosan and its derivatives have been getting ever growing attention in various fields including biotechnology, pharmaceutics, textile, food, cosmetics, biomedical and other industries [18–20]. Because of the cationic properties, chitosan has been broadly investigated by numerous investigators in gene delivery, drug delivery and biomedical areas [21]. Moreover, chitosan own an assortment of biological activities such as antifungal, antibacterial and antioxidant activities. It has instinctive antimicrobial activity due to the fact that chitosan molecules

http://dx.doi.org/10.1016/j.ijbiomac.2015.06.044 0141-8130/© 2015 Published by Elsevier B.V.

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(having positive charge) interact with negatively charged bacteria which in turn causing disruption onto the cell [22]. Along with numerous noteworthy properties, chitosan has some drawbacks i.e. applications of chitosan are restricted by poor solubility. It is only soluble in such acidic aqueous solutions where protonated amino groups are present [16]. Furthermore, chitosan can be deformed through external pressure, which in turn causes intense swelling in acidic aqueous solutions [23]. Therefore, it is hard to use chitosan for some biomedical applications because of its poor mechanical properties. Hence, many researchers tried to chemically modify its properties for further use in various constructive applications [24,25]. Biocompatible polyethylene glycol (PEG) has gained a great deal of consideration in recent years [17–20]. PEG is non-biodegradable polymer and exhibits properties such as immunogenicity, low toxicity and protein resistance. Furthermore, it has been shown that PEGs are able to preserve their biological properties by repealing the immunogenicity of proteins [26,27]. One of the most significant areas of research for PEG is its conjugation to peptides, non-peptide drugs and proteins. The conjugates are usually stable towards degradative enzymes, less immunogenic and less antigenic [28]. PEGs have been selected frequently as drug carriers owing to their competing properties i.e. minimal toxicity, biocompatibility, good solubility in water and some other common solvents. Likewise, with linear aliphatic polyesters such as poly (lactic acid), PEGs can be co-polymerized for use in tissue engineering and drug delivery systems and also for enhancing the biocompatibility of polymers [29–31]. The application of injectable hydrogels as compare to implantation of synthesized hydrogels has been found to be more attractive in biomedical applications because it can reduce patient discomfort, infection risk as well as operation cost [32,33]. In addition injectable hydrogels have received substantial in drug delivery applications since many years as they can be used as injectable controlled release vehicles for confined drug delivery. These injectable delivery systems acquire various advantages including ease of application, localized delivery in case of site-specific action, extended delivery periods, enhanced patient comfort and reduced body drug dosage by lessening unwanted side effects that may occur in systemic delivery [34,35]. Additionally, the reagents used in synthesizing hydrogels must be non-toxic, and after curing disease, the formed hydrogels should be degraded [36]. The objective of present work was to synthesize a novel pH responsive blend hydrogels derived from biodegradable polymers chitosan and poly (ethylene glycol) crosslinked with non-toxic silane crosslinker for controlled drug delivery. Tetraethoxyorthosilicate (TEOS) was selected as crosslinking agent due to its nontoxic nature and it is well accepted in biomaterials applications. At acidic pH, chitosan (CS)/PEG hydrogels exhibited maximum swelling while low swelling was observed at neutral as well as basic pH. The pH sensitivity made these hydrogels an appropriate candidate for injectable drug delivery. The swelling experiments were performed in distilled water and other swelling solvents. The most appropriate hydrogel was selected to achieve goal of work. Antimicrobial tests of synthesized hydrogels were conducted using disk diffusion and well diffusion method against bacterial strains of Staphylococcus aureus (S. aureus) and Escherichia coli (E. coli). S. aureus (Gram+) and E. coli (Gram−) were used as bacteria model. Cefixime (CFX) is widely used cephalosporin antibiotic having anti-bactericidal activity and used to cure bacterial infections (gonorrhoea, pharyngitis, etc.). The injectable release of CFX (model drug) was evaluated in simulated gastric fluid (SGF) and simulated intestinal fluid (SIF). The synthesized drug loaded sample showed controlled release of drug in SIF. The amount of CFX was measured by ultraviolet spectroscopy.

2. Materials and methods

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2.1. Materials

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Chitosan (viscosity = 165 centipoise; molecular weight = 179,268 g/mol; deacetylation degree or DAC = 90.28%) was received from MP Biomedicals, France. PEG of mol. wt. 1000, 4000 and formic acid (HCOOH) (90%) were purchased from BDH laboratory; England. PEG 1500 g/mol (hydroxyl No. 70–80) was provided by Scharlau, Spain. Similarly, PEG of mol. wt. 6000 g/mol was obtained by Merck, Germany. Tetraethoxysilane (TEOS) (98.5%) was supplied by DAEJUNG chemical & metals Co. Ltd., Korea. All chemicals were of analytical grade and used as received. CFX was obtained locally as a model drug. 2.2. Synthesis of hydrogels

2.3. Swelling experiments

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Swelling experiments of hydrogels were performed in different swelling solvents (distilled water and ionic and buffer solutions) by using following procedure. The prepared hydrogels were cut into small piece (∼30.0 mg), weighed and immersed into 40 mL of their respective solvents at room temperature. After specified time intervals, the excess surface solution was gently wiped with tissue paper and swollen weight of the specimen was determined. Same procedure was repeated with all the samples until the equilibrium state was achieved. Data presented in this research work was the mean values of the triplicate measurements. Swelling results were calculated using following equation: (1)

where, Ws is the swollen hydrogel weight and Wd is the dry hydrogel weight at time t [13]. The pH-dependent swellings experiments were performed in buffer solutions (pH 2, 4, 6, 7, 8 and 10) while the swelling response in saline solutions at different concentrations i.e. 0.1, 0.3, 0.5, 0.7, 0.9 and 1 mol/L were evaluated. 2.4. Characterization 2.4.1. IR analysis IR spectra of hydrogels are recorded on Shimadzu IR Prestige21 equipped with (HATR) horizontal attenuated total reflectance kit having zinc selenide (ZnSe) crystal. Spectra were collected in transmittance mode at 4000–400 cm−1 scanning range with 70 scans per spectrum at a resolution of 4 cm−1 . 2.4.2. Thermogravimetric analysis Thermal degradation of synthesized hydrogels was studied by thermo gravimetric analyzer (TA instruments SDTQ600) under

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0.75 g of CS was dissolved in 25 mL of 1% formic acid solution by stirring on hot plate (Wisd laboratory Instruments; Model: MSH 20A). Similarly, PEG solutions of different mol. wts. (1000, 1500, 4000, 6000) were prepared by dissolving 0.1 g in 50 mL 1% formic acid. PEG solution was mixed with CS solution and magnetically stirred for 2 h. TEOS (25 ␮L) dissolved in ethanol (5 mL) was added drop wise in the respective blends and stirred for next 2 h. The crosslinked blends were poured on petri dish, oven dried and finally separated from petri dish. The fabricated blend hydrogels were assigned codes; CP10, CP15, CP40 and CP60, where, digits represents 1000, 1500, 4000, 6000 mol. wt. of PEGs. CT code was used to represent controlled hydrogel containing only CS and the crosslinker.

(Ws − Wd ) Swelling (g/g) = Wd

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Scheme 1. Intra-and inter-molecular hydrogen bonding as well as silane crosslinking between biopolymers (A/B; Intra-/inter-molecular H-bonding).

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nitrogen flow of 15 mL/min to eliminate unnecessary corrosive gas. Heating rate was maintained at 20 ◦ C/min from room temp to 600 ◦ C. 2.4.3. Antimicrobial activity Antimicrobial tests of synthesized hydrogels were conducted using disk diffusion method and well diffusion method against bacterial strains of S. aureus and E. coli. S. aureus (Gram+) and E. coli (Gram−) were used as bacteria model. The disk diffusion method involved inoculating disks of filter paper (1.5 cm in diameter) with the samples of interest (CP15 and CT). The disks were placed on surface of agar plate that was inoculated with an indicator organism (S. aureus and E. coli). The plate was incubated overnight at 35–37 ◦ C inside the circulation over [37]. In well diffusion method, bacterial strains were inoculated on the Brain Heart Infusion (BHI) agar surface. Then wells of about 5 mm in diameter were cut from agar and desired samples (CP15 and CT) were placed in the wells. For 18 h, the plates were incubated in oven at 37 ◦ C [38]. 2.5. Application of synthesized hydrogel for drug delivery

3. Results and discussion The proposed interactions of silane crosslinked blends of CS with PEG (different molecular weights) are presented in Scheme 1.

3.1. Swelling response in water Time dependent swelling response of CS/PEG hydrogels in distilled water is shown in Fig. 1. This figure shows that all the samples have different swelling behaviour with respect to each other. Initially, all the samples show a linear increase in swelling with time and reached their equilibrium state at different time intervals. CP60 has higher swelling than all other samples and attained the maximum swelling (58.0 g/g) after 100 min. Similarly, CP10 shows least swelling at all-time intervals, attained maximum swelling (13.66 g/g) after 40 min. This increase in swelling with increase in molecular weight of PEG is mainly due to the increase in hydrophilicity due to PEG with high molecular weight which confirms a direct relation between molecular weight and hydrophilicity of PEG. Moreover, higher the molecular weight of

2.5.1. Preparation of simulated solutions SGF (pH 1.2) was prepared by mixing sodium chloride (NaCl) (1 g) in hydrochloric acid (HCl) (3.5 mL) and diluted the solution upto 500 mL with the help of distilled water. SIF having pH 6.8 was made by preparing 0.2 M solutions of KH2 PO4 and NaOH and adding KH2 PO4 (250 mL) to NaOH (118 mL). 2.5.2. Drug loading procedure To prepare drug-loaded sample, CFX (25 mg/25 mL of methanol) was loaded into blend of CS (0.75 g) and PEG (0.1 g, Mw = 6000) before the addition of TEOS and stirred for 1 h (total volume of the blend was 80.875 mL). At 37 ◦ C, the drug loaded sample was placed in SGF and SIF solutions (100 mL each). 5 mL was taken from the beaker after every 10 min while the beaker was refilled with same volume of fresh solution (SGF or SIF) to make up the solution volume. Amount of released or unconfined cefixime was determined spectrophotometrically at 288 nm by using UV–vis spectrophotometer (spectronic 20). SIF and SGF were used as the reference standards. Amount of cefixime released by drug loaded hydrogel was calculated with reference standard solutions (100 ppm of drug in SGF and SIF).

Fig. 1. Time dependent swelling behaviour of CP60, CP40, CP15, CP10 and CT in water.

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Fig. 2. Swelling behaviour of CP60, CP40, CP15, CP10 and CT in buffer media of different pH.

Fig. 3. Swelling response of CS/PEG hydrogels in different molar concentrations of NaCl (solid) and CaCl2 (hollow) ionic solutions.

PEG, larger will be the network or pore size which results in lower crosslinking density of the hydrogels [39].

3.3. Swelling response in ionic solutions

3.2. pH-sensitivity of hydrogels Swelling medium and polymer properties are the two important factors that influence the swelling performance of hydrogels. The pH sensitivity of hydrogels is determined in buffer solutions of various pH i.e. 2, 4, 6, 7, 8 and 10. The influence of pH on swelling response of hydrogels is shown in Fig. 2. In all the samples, highest swelling at pH 2 while least at pH 7 is observed. However, the swelling of hydrogels is increased with increase in the molecular weight of PEG. The maximum swelling of 19.38 g/g is exhibited by CP60 at pH 2; its minimum value of 3.93 g/g is obtained at pH 7 while the swelling of CP60 at pH 2 is almost equal to that at pH 10. Hydrogels exhibited a decrease in swelling up to pH 7 and subsequently an increase in swelling values with the increase in pH up to 10. The swelling response of the hydrogels against pH indicated high swelling in acidic and basic pH, whereas, low swelling is observed at pH 6 and 7. It is well-known fact that in hydrogels, high concentration of ionic groups (charged) enhances swelling because of osmosis as well as charge repulsion. The influence of pH on swelling of hydrogels can be described on the basis of protonation of NH2 group of chitosan. In acidic buffer, hydrogels swelling process causes the ionization of amino groups in acid; the acid in turn would then be attached by the ionic bonds to hydrogels. In other words, the active amino groups (–NH2 ) get protonated and transformed into ammonium ions (–NH3 + ). Hence, maximum concentration of ionic groups (–NH3 + ) present in acidic medium enhances swelling of hydrogels as a result of charge repulsion and osmosis. Consequently, weight of the samples is increased in corresponding buffer medium [40–42]. On the other hand, deprotonation of amino groups takes place at higher pH which in turn causes reduction in the repulsion of polymer chains allowing the shrinkage of hydrogels as a result of decrease in the water content. Additionally, when the ionization degree of ionic groups is reduced, swelling also reduced. Therefore, in neutral or basic pH, less swelling of hydrogels is observed [43]. This pH-sensitive swelling behaviour is the most important stuff for hydrogel in order to study the drug release analysis.

Ionic solutions of NaCl and CaCl2 have been used to study the swelling response of hydrogels. Both have same anion (Cl−1 ) but different charge and cation. Effect of ionic concentration of these electrolytes on the swelling behaviour of hydrogels is shown in Fig. 3. This figure exhibits that the swelling capacity of hydrogels is significantly decreased with increase in the concentration of electrolytes. As the concentration of electrolytes in the swelling medium is increased, a charge screening effect of additional electrolytes is achieved, which in turn decreased the osmotic pressure difference developed between the hydrogels and the external salt solution. Therefore, diffusion of solvent into hydrogels decreased which resulted in less swelling. Moreover, in NaCl solution, overall higher swelling is observed as compared to other salt solution. This decrease in swelling is dependent strongly on kind of added salt to swelling medium. Swelling behaviour is affected by the type of cation, depending on charge and different radius of cations. With increase in cation charge, crosslinking degree is increased which consequently causes decrease in swelling [13,44].

3.4. Structural analysis Structural analysis is conducted by Fourier transform infrared spectroscopy. IR spectra of controlled, CP60, CP40 and CP10 are shown in Fig. 4. A band at 1652 cm−1 and band ranges from 1540 to 1570 cm−1 correspond to amide-I and amide-II groups, respectively, while cis amide-III band is observed at 1342 cm−1 . These are the characteristic bands of chitosan. Similarly, the bands at 1151 and 890 cm−1 implied glycosidic linkages [11,13,16,45]. The bands at 2882 and 2920 cm−1 represent stretching vibrations of –CH3 and –CH2 , respectively. The asymmetric stretching bands in the range of 1110–1000 cm−1 indicated –Si–O–C and –Si–O–Si, which confirmed the presence of crosslinking due to TEOS. Furthermore, a stretching band at 1110 cm−1 is due to acyclic –C–O–C (PEG) and at 1230 cm−1 is due to cyclic –C–O–C (chitosan) are observed [46]. A band in the range of 3570–3200 cm−1 leads to –O–H stretching vibrations. By increasing molecular weight of PEG, intensity of this band is increased. Silanol (–Si–OH) stretching vibration is observed at 3725 cm−1 .

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Table 1 Thermogravimetric data of different CS/PEG hydrogels. Sample

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weight loss percent drops significantly by increasing the molecular weight of PEG [47]. That is why CP10 hydrogel with more weight Q3 loss percent is most stable (Table 1). 3.6. Antibacterial activities

Fig. 4. IR spectra of CP60, CP40, CP10 and CT in the range of 4000–500 cm−1 .

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3.5. Thermogravimetric analysis The thermograms of controlled and other hydrogels are shown in Fig. 5. Thermogravimetric analysis is conducted to study the effect of molecular weight of PEG on thermal stability of hydrogels. Figure demonstrates that in CT hydrogel, weight loss from 50 to 270 ◦ C is owing to the moisture removal, loss of bound water and dehydration. Degradation of the main backbone is started around 270 ◦ C which is the onset of degradation of hydrogel. The degradation completed around 350 ◦ C which corresponds to the offset of degradation. The CT hydrogel is composed of chitosan and crosslinker (TEOS) lacking the PEG molecule. Though, in hydrogels, blended with PEG, onset of degradation is observed around 270 ◦ C but the offset of degradation is appeared at temperature around 440 ◦ C which is due to the addition of PEG into hydrogels. TG curves indicated that more temperature is required to degrade the crosslinked CS/PEG hydrogels. It is evident from figure that CP10 is thermally more stable than all other PEG modified samples. The 10% weight loss of CP10 is at 67 ◦ C while it is decreased to 53 ◦ C in CP60 which is thermally least stable. Molecular weight of the chain is the most important factor in order to study the effect of PEG. It seems that the complete approach of larger PEG molecule is hampered by steric hindrance, thus impeding the quantitative reaction and equimolar loadings of different PEG. Hence, it is concluded that

Fig. 5. Thermograms of CP10, CP15, CP40, CP60 and CT.

The antibacterial activities of selective samples (CP15 and CT) were estimated by disc diffusion and well diffusion method. Results indicated that selective samples exhibited antibacterial activity against gram negative bacteria E. coli and gram positive bacteria S. aureus. The interaction between bacterial cell membranes (negatively charged) with chitosan molecules (positively charged) is the most appropriate mechanism behind the antibacterial activity. The electrostatic forces among the electronegative charges of bacterial cell surfaces and NH3 + groups (protonated) of chitosan are basically responsible for this interaction. This electrostatic interaction promotes changes in bacterial membrane wall, thus causing internal osmotic imbalances which in turn inhibiting the bacterial growth [48]. Another proposed mechanism that may occur is the chitosan binding with the bacterial DNA, inhibiting the synthesis of protein and mRNA by penetrating into bacterial nuclei. The incorporated crosslinker moieties and hydrophilic group (PEG and TEOS) kept chitosan chains away from each other by decreasing their intermolecular hydrogen bonding and increase in their solubility. As a result, synthesized hydrogels could easily penetrate into bacterial cell and inhibiting their growth by preventing the DNA transformation into RNA [49]. 3.7. Controlled release examination of CFX CP60 was loaded with model drug (CFX) and its release mechanism was investigated in SIF with respect to time t (Fig. 6). In SGF, the entire loaded drug was released in half an hour as it was observed in swelling analysis (Fig. 2) that hydrogels swollen more at acidic pH. According to this result, synthesized hydrogel could not be used for oral administration. In SIF, CFX was released in a

Fig. 6. Percent release of CFX in SIF (pH 6.8) with time.

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controlled manner and 85% of the total amount was released within 80 min. The remaining amount of drug could not be measured easily because the hydrogel crumbled into large fragments as well as entrapment of drug in the hydrogel. Therefore, the controlled release behaviour was exhibited by hydrogel which showed that CFX can be used for intravenous (IV) medications. Consequently, the injectable hydrogels will be administrated directly into the venous circulations via a syringe. 4. Conclusions

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A novel crosslinked CS/PEG blend with varying molecular weight of PEG has been synthesized and. IR spectra confirmed the presence of crosslinked network suggested that the interaction between incorporated components increased by using high molecular weight PEG. CP10 showed maximum thermal stability as compare to the other samples. CP60 showed maximum swelling (58.0 g/g) in distilled water exhibited the effect of molecular weight of PEG. The pH response demonstrated maximum swelling in acidic and basic pH while low swelling at neutral pH which made these hydrogels appropriate for injectable controlled release behaviour. This mechanism showed the release of drug (CFX) for injectable cefixime dose. As in SIF, sustained amount of drug (CFX) is released (85%) during 80 min. Also, the synthesized hydrogels penetrated into bacterial cell and inhibited their growth by avoiding the DNA conversion into RNA.

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Acknowledgements

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The author is highly obliged to the Department of Polymer Engineering and Technology and Institute of Chemistry, University of the Punjab, Lahore for providing lab facilities.

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Injectable biopolymer based hydrogels for drug delivery applications.

Biopolymer based pH-sensitive hydrogels were prepared using chitosan (CS) with polyethylene glycol (PEG) of different molecular weights in the presenc...
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