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Polyvinyl alcohol composite nanofibres containing conjugated levofloxacin-chitosan for controlled drug release
Materials Science and Engineering C 73 (2017) 440–446
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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Polyvinyl alcohol composite nanofibres containing conjugated levofloxacin-chitosan for controlled drug release Javid Jalvandi a,b,⁎, Max White b, Yuan Gao a, Yen Bach Truong a, Rajiv Padhye b, Ilias Louis Kyratzis a a b
CSIRO, Manufacturing Flagship, Bayview Ave, Clayton, Victoria 3168, Australia School of Fashion and Textiles, College of Design and Social Context, RMIT University, 25 Dawson Street, Brunswick, Victoria 3056, Australia
a r t i c l e
i n f o
Article history: Received 29 July 2016 Received in revised form 7 December 2016 Accepted 20 December 2016 Available online 23 December 2016 Keywords: Controlled release Electrospinning Polyvinyl alcohol Chitosan Levofloxacin Conjugation
a b s t r a c t A range of biodegradable drug-nanofibres composite mats have been reported as drug delivery systems. However, their main disadvantage is the rapid release of the drug immediately after application. This paper reports an improved system based on the incorporation of drug conjugated-chitosan into polyvinyl alcohol (PVA) nanofibers. The results showed that controlled release of levofloxacin (LVF) could be achieved by covalently binding LVF to low molecular weight chitosan (CS) via a cleavable amide bond and then blending the conjugated CS with polyvinyl alcohol (PVA) nanofibres prior to electrospinning. PVA/LVF and PVA-CS/LVF nanofibres were fabricated as controls. The conjugated CS-LVF was characterized by FTIR, DSC, TGA and 1H NMR. Scanning electron microscopy (SEM) showed that the blended CS-PVA nanofibres had a reduced fibre diameter compared to the controls. Drug release profiles showed that burst release was decreased from 90% in the control PVA/LVF electrospun mats to 27% in the PVA/conjugated CS-LVF mats after 8 h in phosphate buffer at 37 °C. This slower release is due to the cleavable bond between LVF and CS that slowly hydrolysed over time at neutral pH. The results indicate that conjugation of the drug to the polymer backbone is an effective way of minimizing burst release behaviour and achieving sustained release of the drug, LVF. © 2016 Elsevier B.V. All rights reserved.
1. Introduction Controlled slow release systems for drugs have been widely explored due to their advantages compared to other methods of drug administration [1,2]. The advantages include reducing the side effects by delivering the drug at a controlled release rate, improved therapeutic efficiency, continuous sustained release of small amounts of drug instead of several large doses and improved of patient compliance [3]. In the past decade several drug delivery systems have been investigated for a controlled release of drugs from nanoparticles [4] and biodegradable nanofibres [5]. Nanofibres are made by a variety of methods such as melt blowing, drawing, phase separation, self-assembly and electrospinning [6–8]. Electrospinning is a simple and convenient technique to fabricate nanofibres from a large variety of polymers [9,10]. In the process of electrospinning, the polymer solution in a syringe is fed through the spinneret at a constant pump rate. When a high voltage is applied, the electrostatic field charges the solution drop at the nozzle of a spinneret and forms the electrified jet that can be attracted by the
⁎ Corresponding author at: School of Fashion and Textiles, College of Design and Social Context, RMIT University, 25 Dawson Street, Brunswick, Victoria 3056, Australia E-mail addresses: [email protected] (J. Jalvandi), [email protected] (M. White), [email protected] (Y. Gao), [email protected] (Y.B. Truong), [email protected] (R. Padhye), [email protected] (I.L. Kyratzis).
http://dx.doi.org/10.1016/j.msec.2016.12.112 0928-4931/© 2016 Elsevier B.V. All rights reserved.
collectors with the opposite electrical charge. As it travels to the collector the solvent is evaporated, the solid polymer nanofibre is deposited as a nonwoven mat on the surface of the collector [11]. Electrospun nanofibres have gained considerable attention as potential drug delivery systems in the last decade [12,13]. Various biodegradable nanofibres have been investigated for controlled release of drugs [14–16]. However, the major drawback of drug loaded nanofibres is the undesired burst release behaviour of the drug [17–20]. For example, ciprofloxacin loaded nanofibres of PVA have been reported [20] with high initial burst release (80%) of the ciprofloxacin content in the first 40 h. The burst release of salicylic acid (SA) from PVA nanofibres with different SA concentration has also been reported [21]. The release profiles showed very high burst release within the first hour with 40–65% of the total drug released. Such burst release behaviour of these drugs may be due to a large proportion of the drug being on the nanofibres' surface and/or rapid drug diffusion out of the polymeric matrix because of fast polymer degradation [20,22,23]. It is important, therefore, to devise systems which overcome the initial burst release. One approach for achieving a controlled release of drug is covalent conjugation of drugs to the polymer backbones [24– 26]. This requires the presence of functional groups on both the drug and polymer. For example, docetaxel (DTX) was covalently attached to low molecular weight CS via a cleavable amide linker. Prolonged slow release of DTX was obtained in simulated intestinal fluid
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(pH 7.5) which resulted in higher bioavailability along with prolonged circulation time in the blood stream than the intravenously injected DTX [27]. In another study, composite nanofibres of poly(lactide-coglycolide) (PLGA) and poly(ethylene glycol)-g-chitosan-ibuprofen (PEG-g-CS-IBU) conjugates were fabricated by covalently linking IBU to the amine groups of CS via an amide bond [28]. There it was observed that presence of conjugated CS-IBU in the matrix reduced the burst release due to the cleavable amide bond between CS and ibuprofen. Another report includes synthesis of doxorubicin conjugated stearic acidg-chitosan polymeric micelles (DOX–CS–SA) [29]. The release profile showed that release behaviour of DOX is pH-dependent where the release rate of DOX increased with the reduction of the pH for release medium from 7.2 to 5.0. PVA is widely used in biomedical applications due to its well-known characteristics of biocompatibility, biodegradability, water solubility, nontoxicity, swelling capability in an aqueous medium, and appropriate mechanical properties [30–32]. It has been reported that by incorporating a second polymer component such as CS into the PVA solution, leads to an improvement in biocompatibility and mechanical properties (e.g. solubility) of the nanofibres. Chuang et al. [33] fabricated PVA/CS blended membranes for cell culture. They observed that the PVA/CS blended membrane was more favourable for the cell culture with better spreading of cultured cells than the pure PVA membrane. The aim of this work was to design a controlled drug delivery system by conjugating the carboxylic groups in LVF to the amine groups of CS [34,35], followed by mixing the conjugated system with PVA prior to electrospinning in order to reduce the burst release of drug from highly water soluble PVA nanofibres [20,21]. The release profiles of PVA/CSLVF (conjugated) electrospun composite mats were measured and compared to the control PVA/LVF and PVA-CS/LVF (non-conjugated) electrospun composite mats. 2. Materials and methods 2.1. Materials PVA (MW 85,000–124,000, 98–99% hydrolysed), low molecular weight CS (deacetylation ≥ 75%), LVF (98.0%) and acetic acid were sourced from Sigma-Aldrich. 1-Ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) crosslinker was supplied by Pierce, Dimethyl sulfoxide (DMSO) from Fisher Chemical and sodium hydroxide from Chem Supply. All other chemicals were of reagent grade. 2.2. Methods 2.2.1. Synthesis of CS-LVF The conjugation of LVF and CS was based on carbodiimide chemistry (Fig. 1) [34,35]. CS (1 g) was dissolved in 0.1 M acetic acid solution (500 mL). A mixture solution of LVF and EDC (1:1 M ratio) was prepared
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by dissolving LVF and EDC in 10 mL anhydrous dimethyl sulfoxide (DMSO). Then, the mixture solution was added dropwise to the CS solution under continuous stirring for 12 h. The conjugated CS was then coagulated by adding 1 M NaOH. CS-LVF was collected by centrifugation (15 min, 16,000 g), washed by DI water and then dried at room temperature for 24 h. 2.2.2. Preparation of composite electrospun mats Electrospun mats were prepared containing 1 wt.% of LVF, and 5 wt.% of CS-LVF (with respect to PVA). For polymer solutions 20 mg LVF, 100 mg CS-LVF were added to two separate vials containing 20 g of 10 wt.% PVA (dissolved in water) and stirred continuously until a homogeneous solution was obtained. A solution of 10 wt.% PVA-CS (9:1) was also prepared with 1 wt.% LVF (with respect to PVA-CS). The basic setup used for electrospinning is shown in Fig. 2. The setup contained a spinneret needle (23 G Terumo), syringe pumps (model NE-1000 New Era Pump Systems, Inc., Farmingdale, NY, USA), a high voltage supply (Spellman high voltage DC power supply, USA) and a 5 mL syringes (Terumo). The electrospinning distance was varied between 0 and 300 mm (optimised distance of 150 mm) and the applied voltage varied between 0 and 40 (optimised value of 12 kV). Electrospun mats were collected for 10 h at 0.6 mL/h pump rate on a drum collector (100 mm diameter) driven by a 12 V motor rotating at 60 mms−1 and a moving horizontal oscillating speed of 10 mms−1. 2.3. Characterization IR spectra were collected in the transmission mode using a Thermo Nicolet 6700 FTIR Spectrometer. DSC analysis of samples was carried out at a heating rate of 10 °C/min under a nitrogen purge of 40 mL/min using a Netzsch STA 449 F1 Jupiter DSC Thermal Analyser. TGA was performed using a Netzsch STA 449 F1 Jupiter TGA Thermal Analyser at a heating rate of 10 °C/min under a nitrogen purge of 40 mL/min. 1H NMR spectra of the samples were recorded on a Bruker Av400 NMR spectrometers. Scanning electron microscopy (SEM) analyses were carried out using a Philips XL30 field microscope. Fibre diameters were measured using microstructure measurement software. For each sample, 100 measurements were taken to obtain an average fibre diameter. A Varian UV–Vis spectrophotometer was used to measure the LVF concentration at 290 nm. 2.4. In vitro drug release The electrospun mats were cut into 50 mm squares, weighed and then placed in 5 mL of phosphate buffer saline solution (PBS, 100 mM, pH 7.4 and pH 5.0) and shaken continuously for 3 days at 150 rpm at 37 °C. At each measurement time point 1 mL of release medium was removed and replenished with 1 mL of fresh buffer. The amount of LVF released at each time point was determined by UV spectrophotometer.
Fig. 1. Preparation of CS-LVF.
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Fig. 2. Schematic diagram of preparation of CS-LVF and PVA/CS-LVF nanofibres.
The UV absorbance of LVF in buffer solutions was measured at 290 nm and converted to the LVF concentration according to a calibration curve. LVF release profiles were calculated from the LVF concentration at each time point. 3. Results and discussion In this study we report a novel drug delivery system which improves the release profile from highly soluble PVA nanofibres. The system is based on conjugating LVF to CS, combining them with PVA and then fabricating composite nanofibres by electrospinning. Again LVF was used as the model drug for its functional carboxylic acid group and ease of analytical detection by absorption at its peak at 290 nm. 3.1. FTIR FTIR spectra of LVF, CS and CS-LVF were obtained and shown in Fig. 3. Characteristic FTIR transmission peaks of LVF can be observed at
Fig. 3. FTIR spectra of LVF, CS and CS-LVF.
1294 cm−1 (assigned to stretching of amines), 1620 cm−1 for aromatic C\\C, 1720 cm−1 due to carboxyl C_O and a broad peak at 3265 cm−1 representing O\\H of carboxylic acid. The peak assignment of CS shows characteristic bands at 1643 cm− 1 due to NH2 scissoring vibration, 1590 cm−1 for carbonyl asymmetric stretching, and at 1028 cm−1 due to the C\\O of the pyranose ring [36,37]. The conjugated CS-LVF spectrum shows a peak at 1652 cm− 1 due to a secondary amide band (C\\O stretch of acetyl group); and the peak at 1575 cm−1 most likely due to N\\H stretch of amide bond, confirming the formation of amide bond [35,37–39]. 3.2. DSC DSC analysis was conducted to determine the interaction of CS with LVF. Thermograms of LVF, CS, and CS-LVF are presented in Fig. 4. The DSC thermogram of pure CS indicated one exothermic transition at 309.6 °C along with one endothermic transition at 101.9 °C. The exothermic transition is indicative of CS degradation at 309.6 °C and the occurrence of an endothermic transition at 101.9 °C could be attributed to the moisture present in samples or the presence of acetic acid moieties [40–42]. The DSC thermogram of LVF showed endothermic transitions at 94.2 °C and 237.2 °C due to the decomposition of LVF. The CS-LVF conjugate showed a new sharp endothermic transition at 250.5 °C. This endothermic transition at 250.5 °C may have arisen due to the carboxylic
Fig. 4. DSC thermograms of LVF, CS and CS-LVF.
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ascribed to dehydration of the sugar rings, depolymerization and decomposition of the acetylated and deacetylated units of the polymer [44]. The TGA spectra of the conjugated CS-LVF demonstrated slow weight loss at about 90 °C due to evaporation of water and moisture content in the polysaccharide. The second decomposition stage starts at about 195 °C which is lower than that of CS. The fast process of weight loss appears from 195 to 310 °C. This is attributed to the thermal destabilization of the CS chains due to the conjugation of LVF in a polymer chain and related dissociation of the strong intra-molecular bonding [43,45]. 3.4. 1H NMR
Fig. 5. TGA spectra of LVF, CS and conjugated CS-LVF.
linkage between the –COO− moieties of LVF and –NH3+ moieties of CS and formation of an amide bond between LVF and CS [41].
The formation of an amide linkage between CS and LVF was confirmed by 1H NMR spectroscopy. As shown in Fig. 6, CS-LVF gave rise to a series of peaks at 4.60, 3.11, 3.90 and 3.72 ppm corresponding to 1 H linked to the carbon atoms in the sugar ring (Ha, Hb, Hc, Hd, respectively) [39,46]. A new signal appeared at δ = 2.5–2.7 ppm corresponding to Ha due to the formation of an amide bond in between LVF and CS chains [46–48]. 3.5. Morphology of the nanofibres
3.3. TGA The thermogram of CS exhibits two distinct stages of weight loss (Fig. 5). One is in the range of 30–107 °C due to the loss of adsorbed and bound water and decomposition of low molecular weight polymers. The second stage starts at 240 °C and continues up to 370 °C due to the degradation of CS biopolymer [43]. This second stage is
The morphology and fibre diameters of the electrospun mats with diameter sizes distributions are shown in Fig. 7A–C. PVA electrospun nanofibres have a uniform structure and fibre diameters are similar to the PVA/CS composite. The addition of CS in the electrospinning solution greatly influenced the morphology of the electrospun nanofibres [49]. The fibre diameter distribution of the PVA/LVF nanofibres followed
Fig. 6. 1H NMR spectra of LVF, CS and CS-LVF.
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Fig. 7. SEM images of electrospun PVA/LVF (A), PVA-CS/LVF (B) and PVA/CS-LVF (C); and their relative diameter distribution.
a normal distribution. Incorporation of CS reduced the fibre diameter by approximately 50%. The average fibre diameter decreased from 200 to 300 nm to 100–200 nm. This reduction in diameter is due to the presence of CS a cationic polysaccharide with amino groups, which are
ionizable under acidic conditions. It has been reported that the addition of cationic and anionic polyelectrolytes affects the conductivity of polymer solution thus resulting in a thinner fibre [50,51]. 3.6. Cumulative release of LVF
Fig. 8. The cumulative release of LVF from: PVA/LVF, PVA-CS/LVF and PVA/conjugated CSLVF electrospun mats (pH 5.0 and pH 7.4).
Fig. 8 shows the release profiles of the LVF at two pH levels from the controls and two composite nanofibres; one with CS/LVF in the composition and the other with drug covalently attached to CS (Conjugated CS-LVF). The uniform PVA/LVF electrospun mat showed an extreme burst nature, releasing 90% of the drug in the first 8 h. This very fast release behaviour of LVF was attributed to the swelling properties and dissolution of the PVA matrix within the medium. It has been reported that as soon as the PVA matrix began to swell, LVF molecules were leached out from the matrix very quickly [52,53]. A similar phenomena has been reported by Kenawy et al. [54]; where they observed a fast burst release of ketoprofen from PVA with a release of about 85.24% of the drug content within the first 2 h. The amount of LVF released increased with decreasing pH value. As shown in Fig. 8 the addition of CS in to PVA matrix only slightly reduced the burst release to 82% after 8 h in pH 7.4. This indicated that addition of CS to the PVA matrix may have enhanced the swelling and dissolution properties of PVA [55]. The release profile of LVF from PVA-CS/LVF at
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pH 5.0 shows a faster release than those at pH 7.4 due to faster dissolution of CS in the acidic medium. In other words, incorporation of hydrophobic CS into PVA nanofibres decreased the diffusion and therefore the release from the nanofibres at pH 7.4. A recent report shows that composite nanofibres of PVA and polyvinyl acetate could reduce the burst release of ciprofloxacin to 55% after 24 h from 82% in control PVA nanofibres [20]. On the other hand, the release profile of the PVA/conjugated CS-LVF nanofibres at pH 7.4 was significantly different to the other two nanofibres. The release was 27% after 8 h and gradually increased to 59% after 72 h. At pH 5.0 it released 32% of LVF content after 8 h at pH 5 and reached to 85% after 72 h. This is due to the covalently grafted (via an amide bond) to CS. The release occurs first by CS dissociation and hydrolysis of the amide bond from composite nanofibres and then by simple diffusion. The cleavage of the amide bond depends on the pH value, the lower the pH value the faster hydrolysis of amide bond and higher drug release from the composite nanofibres containing conjugated CS [56]. These results indicate the notable stability of the amide bond at pH 7.4 which resulted in limiting the diffusion rate of LVF in the matrix. A similar drug release profile has been reported by Jiang et al. [28]. Composite nanofibres of poly(lactide-co-glycolide) (PLGA) and poly(ethylene glycol)-g-chitosan (PEG-g-CS) with ibuprofen (IBU) conjugated to CS were fabricated. They observed that the burst release of IBU from PLGA membrane dropped from 85% to 20% in PLGA-PEG-g-CS (IBU was covalently conjugated to the CS via amide bond) after 4 days. This slow release is due to the amide bond between IBU and CS. Another report shows the slow release behaviour of IBU from polyethylene glycol hydroxyethyl amide (PEG-HEA) due to the hydrolysis of amide bond between IBU and hydroxyl ethyl amide [57]. 4. Conclusion This paper describes a new composite nanofibres for achieving sustained drug release. The novel nanofibres combines covalently binding the carboxyl group of the LVF to the amine group of CS through the formation of an amide bond and confining the conjugated LVF into PVA nanofibres using electrospinning. Such an approach significantly reduced the burst release of the drug from uniform PVA nanofibres after 8 h and led to a sustained release over 72 h in phosphate buffer. The findings show that chemically binding a drug to the polymer backbone of CS and distribution of the conjugated polymer in nanofibres provides an effective means of achieving sustained drug release efficacy. The improved drug delivery, with approximately 50% of the drug was delivered over 72 h, is due to the cleavable bond between the drug and the nanocarrier. Acknowledgement This work has been supported by CSIRO Manufacturing. References [1] V.V. Ranade, J.B. Cannon, Drug Delivery Systems, CRC Press, 2011. [2] K. Jain, Drug delivery systems—an overview, in: K. Jain (Ed.), Drug Delivery Systems, Humana Press 2008, pp. 1–50. [3] D. Paolino, P. Sinha, M. Fresta, M. Ferrari, Drug delivery systems, Encyclopedia of Medical Devices and Instrumentation, 2006. [4] I.I. Slowing, J.L. Vivero-Escoto, C.-W. Wu, V.S.-Y. Lin, Mesoporous silica nanoparticles as controlled release drug delivery and gene transfection carriers, Adv. Drug Deliv. Rev. 60 (11) (2008) 1278–1288. [5] P. Karuppuswamy, J. Reddy Venugopal, B. Navaneethan, A. Luwang Laiva, S. Ramakrishna, Polycaprolactone nanofibers for the controlled release of tetracycline hydrochloride, Mater. Lett. 141 (2015) 180–186. [6] M. Naghibzadeh, M. Adabi, Evaluation of effective electrospinning parameters controlling gelatin nanofibers diameter via modelling artificial neural networks, Fibers Polym. 15 (4) (2014) 767–777. [7] C. Ribeiro, V. Sencadas, J.L.G. Ribelles, S. Lanceros-Méndez, Influence of processing conditions on polymorphism and nanofiber morphology of electroactive poly(vinylidene fluoride) electrospun membranes, Soft Mater. 8 (3) (2010) 274–287.
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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Polyvinyl alcohol composite nanofibres containing conjugated levofloxacin-chitosan for controlled drug release Javid Jalvandi a,b,⁎, Max White b, Yuan Gao a, Yen Bach Truong a, Rajiv Padhye b, Ilias Louis Kyratzis a a b
CSIRO, Manufacturing Flagship, Bayview Ave, Clayton, Victoria 3168, Australia School of Fashion and Textiles, College of Design and Social Context, RMIT University, 25 Dawson Street, Brunswick, Victoria 3056, Australia
a r t i c l e
i n f o
Article history: Received 29 July 2016 Received in revised form 7 December 2016 Accepted 20 December 2016 Available online 23 December 2016 Keywords: Controlled release Electrospinning Polyvinyl alcohol Chitosan Levofloxacin Conjugation
a b s t r a c t A range of biodegradable drug-nanofibres composite mats have been reported as drug delivery systems. However, their main disadvantage is the rapid release of the drug immediately after application. This paper reports an improved system based on the incorporation of drug conjugated-chitosan into polyvinyl alcohol (PVA) nanofibers. The results showed that controlled release of levofloxacin (LVF) could be achieved by covalently binding LVF to low molecular weight chitosan (CS) via a cleavable amide bond and then blending the conjugated CS with polyvinyl alcohol (PVA) nanofibres prior to electrospinning. PVA/LVF and PVA-CS/LVF nanofibres were fabricated as controls. The conjugated CS-LVF was characterized by FTIR, DSC, TGA and 1H NMR. Scanning electron microscopy (SEM) showed that the blended CS-PVA nanofibres had a reduced fibre diameter compared to the controls. Drug release profiles showed that burst release was decreased from 90% in the control PVA/LVF electrospun mats to 27% in the PVA/conjugated CS-LVF mats after 8 h in phosphate buffer at 37 °C. This slower release is due to the cleavable bond between LVF and CS that slowly hydrolysed over time at neutral pH. The results indicate that conjugation of the drug to the polymer backbone is an effective way of minimizing burst release behaviour and achieving sustained release of the drug, LVF. © 2016 Elsevier B.V. All rights reserved.
1. Introduction Controlled slow release systems for drugs have been widely explored due to their advantages compared to other methods of drug administration [1,2]. The advantages include reducing the side effects by delivering the drug at a controlled release rate, improved therapeutic efficiency, continuous sustained release of small amounts of drug instead of several large doses and improved of patient compliance [3]. In the past decade several drug delivery systems have been investigated for a controlled release of drugs from nanoparticles [4] and biodegradable nanofibres [5]. Nanofibres are made by a variety of methods such as melt blowing, drawing, phase separation, self-assembly and electrospinning [6–8]. Electrospinning is a simple and convenient technique to fabricate nanofibres from a large variety of polymers [9,10]. In the process of electrospinning, the polymer solution in a syringe is fed through the spinneret at a constant pump rate. When a high voltage is applied, the electrostatic field charges the solution drop at the nozzle of a spinneret and forms the electrified jet that can be attracted by the
⁎ Corresponding author at: School of Fashion and Textiles, College of Design and Social Context, RMIT University, 25 Dawson Street, Brunswick, Victoria 3056, Australia E-mail addresses: [email protected] (J. Jalvandi), [email protected] (M. White), [email protected] (Y. Gao), [email protected] (Y.B. Truong), [email protected] (R. Padhye), [email protected] (I.L. Kyratzis).
http://dx.doi.org/10.1016/j.msec.2016.12.112 0928-4931/© 2016 Elsevier B.V. All rights reserved.
collectors with the opposite electrical charge. As it travels to the collector the solvent is evaporated, the solid polymer nanofibre is deposited as a nonwoven mat on the surface of the collector [11]. Electrospun nanofibres have gained considerable attention as potential drug delivery systems in the last decade [12,13]. Various biodegradable nanofibres have been investigated for controlled release of drugs [14–16]. However, the major drawback of drug loaded nanofibres is the undesired burst release behaviour of the drug [17–20]. For example, ciprofloxacin loaded nanofibres of PVA have been reported [20] with high initial burst release (80%) of the ciprofloxacin content in the first 40 h. The burst release of salicylic acid (SA) from PVA nanofibres with different SA concentration has also been reported [21]. The release profiles showed very high burst release within the first hour with 40–65% of the total drug released. Such burst release behaviour of these drugs may be due to a large proportion of the drug being on the nanofibres' surface and/or rapid drug diffusion out of the polymeric matrix because of fast polymer degradation [20,22,23]. It is important, therefore, to devise systems which overcome the initial burst release. One approach for achieving a controlled release of drug is covalent conjugation of drugs to the polymer backbones [24– 26]. This requires the presence of functional groups on both the drug and polymer. For example, docetaxel (DTX) was covalently attached to low molecular weight CS via a cleavable amide linker. Prolonged slow release of DTX was obtained in simulated intestinal fluid
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(pH 7.5) which resulted in higher bioavailability along with prolonged circulation time in the blood stream than the intravenously injected DTX [27]. In another study, composite nanofibres of poly(lactide-coglycolide) (PLGA) and poly(ethylene glycol)-g-chitosan-ibuprofen (PEG-g-CS-IBU) conjugates were fabricated by covalently linking IBU to the amine groups of CS via an amide bond [28]. There it was observed that presence of conjugated CS-IBU in the matrix reduced the burst release due to the cleavable amide bond between CS and ibuprofen. Another report includes synthesis of doxorubicin conjugated stearic acidg-chitosan polymeric micelles (DOX–CS–SA) [29]. The release profile showed that release behaviour of DOX is pH-dependent where the release rate of DOX increased with the reduction of the pH for release medium from 7.2 to 5.0. PVA is widely used in biomedical applications due to its well-known characteristics of biocompatibility, biodegradability, water solubility, nontoxicity, swelling capability in an aqueous medium, and appropriate mechanical properties [30–32]. It has been reported that by incorporating a second polymer component such as CS into the PVA solution, leads to an improvement in biocompatibility and mechanical properties (e.g. solubility) of the nanofibres. Chuang et al. [33] fabricated PVA/CS blended membranes for cell culture. They observed that the PVA/CS blended membrane was more favourable for the cell culture with better spreading of cultured cells than the pure PVA membrane. The aim of this work was to design a controlled drug delivery system by conjugating the carboxylic groups in LVF to the amine groups of CS [34,35], followed by mixing the conjugated system with PVA prior to electrospinning in order to reduce the burst release of drug from highly water soluble PVA nanofibres [20,21]. The release profiles of PVA/CSLVF (conjugated) electrospun composite mats were measured and compared to the control PVA/LVF and PVA-CS/LVF (non-conjugated) electrospun composite mats. 2. Materials and methods 2.1. Materials PVA (MW 85,000–124,000, 98–99% hydrolysed), low molecular weight CS (deacetylation ≥ 75%), LVF (98.0%) and acetic acid were sourced from Sigma-Aldrich. 1-Ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) crosslinker was supplied by Pierce, Dimethyl sulfoxide (DMSO) from Fisher Chemical and sodium hydroxide from Chem Supply. All other chemicals were of reagent grade. 2.2. Methods 2.2.1. Synthesis of CS-LVF The conjugation of LVF and CS was based on carbodiimide chemistry (Fig. 1) [34,35]. CS (1 g) was dissolved in 0.1 M acetic acid solution (500 mL). A mixture solution of LVF and EDC (1:1 M ratio) was prepared
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by dissolving LVF and EDC in 10 mL anhydrous dimethyl sulfoxide (DMSO). Then, the mixture solution was added dropwise to the CS solution under continuous stirring for 12 h. The conjugated CS was then coagulated by adding 1 M NaOH. CS-LVF was collected by centrifugation (15 min, 16,000 g), washed by DI water and then dried at room temperature for 24 h. 2.2.2. Preparation of composite electrospun mats Electrospun mats were prepared containing 1 wt.% of LVF, and 5 wt.% of CS-LVF (with respect to PVA). For polymer solutions 20 mg LVF, 100 mg CS-LVF were added to two separate vials containing 20 g of 10 wt.% PVA (dissolved in water) and stirred continuously until a homogeneous solution was obtained. A solution of 10 wt.% PVA-CS (9:1) was also prepared with 1 wt.% LVF (with respect to PVA-CS). The basic setup used for electrospinning is shown in Fig. 2. The setup contained a spinneret needle (23 G Terumo), syringe pumps (model NE-1000 New Era Pump Systems, Inc., Farmingdale, NY, USA), a high voltage supply (Spellman high voltage DC power supply, USA) and a 5 mL syringes (Terumo). The electrospinning distance was varied between 0 and 300 mm (optimised distance of 150 mm) and the applied voltage varied between 0 and 40 (optimised value of 12 kV). Electrospun mats were collected for 10 h at 0.6 mL/h pump rate on a drum collector (100 mm diameter) driven by a 12 V motor rotating at 60 mms−1 and a moving horizontal oscillating speed of 10 mms−1. 2.3. Characterization IR spectra were collected in the transmission mode using a Thermo Nicolet 6700 FTIR Spectrometer. DSC analysis of samples was carried out at a heating rate of 10 °C/min under a nitrogen purge of 40 mL/min using a Netzsch STA 449 F1 Jupiter DSC Thermal Analyser. TGA was performed using a Netzsch STA 449 F1 Jupiter TGA Thermal Analyser at a heating rate of 10 °C/min under a nitrogen purge of 40 mL/min. 1H NMR spectra of the samples were recorded on a Bruker Av400 NMR spectrometers. Scanning electron microscopy (SEM) analyses were carried out using a Philips XL30 field microscope. Fibre diameters were measured using microstructure measurement software. For each sample, 100 measurements were taken to obtain an average fibre diameter. A Varian UV–Vis spectrophotometer was used to measure the LVF concentration at 290 nm. 2.4. In vitro drug release The electrospun mats were cut into 50 mm squares, weighed and then placed in 5 mL of phosphate buffer saline solution (PBS, 100 mM, pH 7.4 and pH 5.0) and shaken continuously for 3 days at 150 rpm at 37 °C. At each measurement time point 1 mL of release medium was removed and replenished with 1 mL of fresh buffer. The amount of LVF released at each time point was determined by UV spectrophotometer.
Fig. 1. Preparation of CS-LVF.
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Fig. 2. Schematic diagram of preparation of CS-LVF and PVA/CS-LVF nanofibres.
The UV absorbance of LVF in buffer solutions was measured at 290 nm and converted to the LVF concentration according to a calibration curve. LVF release profiles were calculated from the LVF concentration at each time point. 3. Results and discussion In this study we report a novel drug delivery system which improves the release profile from highly soluble PVA nanofibres. The system is based on conjugating LVF to CS, combining them with PVA and then fabricating composite nanofibres by electrospinning. Again LVF was used as the model drug for its functional carboxylic acid group and ease of analytical detection by absorption at its peak at 290 nm. 3.1. FTIR FTIR spectra of LVF, CS and CS-LVF were obtained and shown in Fig. 3. Characteristic FTIR transmission peaks of LVF can be observed at
Fig. 3. FTIR spectra of LVF, CS and CS-LVF.
1294 cm−1 (assigned to stretching of amines), 1620 cm−1 for aromatic C\\C, 1720 cm−1 due to carboxyl C_O and a broad peak at 3265 cm−1 representing O\\H of carboxylic acid. The peak assignment of CS shows characteristic bands at 1643 cm− 1 due to NH2 scissoring vibration, 1590 cm−1 for carbonyl asymmetric stretching, and at 1028 cm−1 due to the C\\O of the pyranose ring [36,37]. The conjugated CS-LVF spectrum shows a peak at 1652 cm− 1 due to a secondary amide band (C\\O stretch of acetyl group); and the peak at 1575 cm−1 most likely due to N\\H stretch of amide bond, confirming the formation of amide bond [35,37–39]. 3.2. DSC DSC analysis was conducted to determine the interaction of CS with LVF. Thermograms of LVF, CS, and CS-LVF are presented in Fig. 4. The DSC thermogram of pure CS indicated one exothermic transition at 309.6 °C along with one endothermic transition at 101.9 °C. The exothermic transition is indicative of CS degradation at 309.6 °C and the occurrence of an endothermic transition at 101.9 °C could be attributed to the moisture present in samples or the presence of acetic acid moieties [40–42]. The DSC thermogram of LVF showed endothermic transitions at 94.2 °C and 237.2 °C due to the decomposition of LVF. The CS-LVF conjugate showed a new sharp endothermic transition at 250.5 °C. This endothermic transition at 250.5 °C may have arisen due to the carboxylic
Fig. 4. DSC thermograms of LVF, CS and CS-LVF.
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ascribed to dehydration of the sugar rings, depolymerization and decomposition of the acetylated and deacetylated units of the polymer [44]. The TGA spectra of the conjugated CS-LVF demonstrated slow weight loss at about 90 °C due to evaporation of water and moisture content in the polysaccharide. The second decomposition stage starts at about 195 °C which is lower than that of CS. The fast process of weight loss appears from 195 to 310 °C. This is attributed to the thermal destabilization of the CS chains due to the conjugation of LVF in a polymer chain and related dissociation of the strong intra-molecular bonding [43,45]. 3.4. 1H NMR
Fig. 5. TGA spectra of LVF, CS and conjugated CS-LVF.
linkage between the –COO− moieties of LVF and –NH3+ moieties of CS and formation of an amide bond between LVF and CS [41].
The formation of an amide linkage between CS and LVF was confirmed by 1H NMR spectroscopy. As shown in Fig. 6, CS-LVF gave rise to a series of peaks at 4.60, 3.11, 3.90 and 3.72 ppm corresponding to 1 H linked to the carbon atoms in the sugar ring (Ha, Hb, Hc, Hd, respectively) [39,46]. A new signal appeared at δ = 2.5–2.7 ppm corresponding to Ha due to the formation of an amide bond in between LVF and CS chains [46–48]. 3.5. Morphology of the nanofibres
3.3. TGA The thermogram of CS exhibits two distinct stages of weight loss (Fig. 5). One is in the range of 30–107 °C due to the loss of adsorbed and bound water and decomposition of low molecular weight polymers. The second stage starts at 240 °C and continues up to 370 °C due to the degradation of CS biopolymer [43]. This second stage is
The morphology and fibre diameters of the electrospun mats with diameter sizes distributions are shown in Fig. 7A–C. PVA electrospun nanofibres have a uniform structure and fibre diameters are similar to the PVA/CS composite. The addition of CS in the electrospinning solution greatly influenced the morphology of the electrospun nanofibres [49]. The fibre diameter distribution of the PVA/LVF nanofibres followed
Fig. 6. 1H NMR spectra of LVF, CS and CS-LVF.
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Fig. 7. SEM images of electrospun PVA/LVF (A), PVA-CS/LVF (B) and PVA/CS-LVF (C); and their relative diameter distribution.
a normal distribution. Incorporation of CS reduced the fibre diameter by approximately 50%. The average fibre diameter decreased from 200 to 300 nm to 100–200 nm. This reduction in diameter is due to the presence of CS a cationic polysaccharide with amino groups, which are
ionizable under acidic conditions. It has been reported that the addition of cationic and anionic polyelectrolytes affects the conductivity of polymer solution thus resulting in a thinner fibre [50,51]. 3.6. Cumulative release of LVF
Fig. 8. The cumulative release of LVF from: PVA/LVF, PVA-CS/LVF and PVA/conjugated CSLVF electrospun mats (pH 5.0 and pH 7.4).
Fig. 8 shows the release profiles of the LVF at two pH levels from the controls and two composite nanofibres; one with CS/LVF in the composition and the other with drug covalently attached to CS (Conjugated CS-LVF). The uniform PVA/LVF electrospun mat showed an extreme burst nature, releasing 90% of the drug in the first 8 h. This very fast release behaviour of LVF was attributed to the swelling properties and dissolution of the PVA matrix within the medium. It has been reported that as soon as the PVA matrix began to swell, LVF molecules were leached out from the matrix very quickly [52,53]. A similar phenomena has been reported by Kenawy et al. [54]; where they observed a fast burst release of ketoprofen from PVA with a release of about 85.24% of the drug content within the first 2 h. The amount of LVF released increased with decreasing pH value. As shown in Fig. 8 the addition of CS in to PVA matrix only slightly reduced the burst release to 82% after 8 h in pH 7.4. This indicated that addition of CS to the PVA matrix may have enhanced the swelling and dissolution properties of PVA [55]. The release profile of LVF from PVA-CS/LVF at
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pH 5.0 shows a faster release than those at pH 7.4 due to faster dissolution of CS in the acidic medium. In other words, incorporation of hydrophobic CS into PVA nanofibres decreased the diffusion and therefore the release from the nanofibres at pH 7.4. A recent report shows that composite nanofibres of PVA and polyvinyl acetate could reduce the burst release of ciprofloxacin to 55% after 24 h from 82% in control PVA nanofibres [20]. On the other hand, the release profile of the PVA/conjugated CS-LVF nanofibres at pH 7.4 was significantly different to the other two nanofibres. The release was 27% after 8 h and gradually increased to 59% after 72 h. At pH 5.0 it released 32% of LVF content after 8 h at pH 5 and reached to 85% after 72 h. This is due to the covalently grafted (via an amide bond) to CS. The release occurs first by CS dissociation and hydrolysis of the amide bond from composite nanofibres and then by simple diffusion. The cleavage of the amide bond depends on the pH value, the lower the pH value the faster hydrolysis of amide bond and higher drug release from the composite nanofibres containing conjugated CS [56]. These results indicate the notable stability of the amide bond at pH 7.4 which resulted in limiting the diffusion rate of LVF in the matrix. A similar drug release profile has been reported by Jiang et al. [28]. Composite nanofibres of poly(lactide-co-glycolide) (PLGA) and poly(ethylene glycol)-g-chitosan (PEG-g-CS) with ibuprofen (IBU) conjugated to CS were fabricated. They observed that the burst release of IBU from PLGA membrane dropped from 85% to 20% in PLGA-PEG-g-CS (IBU was covalently conjugated to the CS via amide bond) after 4 days. This slow release is due to the amide bond between IBU and CS. Another report shows the slow release behaviour of IBU from polyethylene glycol hydroxyethyl amide (PEG-HEA) due to the hydrolysis of amide bond between IBU and hydroxyl ethyl amide [57]. 4. Conclusion This paper describes a new composite nanofibres for achieving sustained drug release. The novel nanofibres combines covalently binding the carboxyl group of the LVF to the amine group of CS through the formation of an amide bond and confining the conjugated LVF into PVA nanofibres using electrospinning. Such an approach significantly reduced the burst release of the drug from uniform PVA nanofibres after 8 h and led to a sustained release over 72 h in phosphate buffer. The findings show that chemically binding a drug to the polymer backbone of CS and distribution of the conjugated polymer in nanofibres provides an effective means of achieving sustained drug release efficacy. The improved drug delivery, with approximately 50% of the drug was delivered over 72 h, is due to the cleavable bond between the drug and the nanocarrier. Acknowledgement This work has been supported by CSIRO Manufacturing. References [1] V.V. Ranade, J.B. Cannon, Drug Delivery Systems, CRC Press, 2011. [2] K. Jain, Drug delivery systems—an overview, in: K. Jain (Ed.), Drug Delivery Systems, Humana Press 2008, pp. 1–50. [3] D. Paolino, P. Sinha, M. Fresta, M. Ferrari, Drug delivery systems, Encyclopedia of Medical Devices and Instrumentation, 2006. [4] I.I. Slowing, J.L. Vivero-Escoto, C.-W. Wu, V.S.-Y. Lin, Mesoporous silica nanoparticles as controlled release drug delivery and gene transfection carriers, Adv. Drug Deliv. Rev. 60 (11) (2008) 1278–1288. [5] P. Karuppuswamy, J. Reddy Venugopal, B. Navaneethan, A. Luwang Laiva, S. Ramakrishna, Polycaprolactone nanofibers for the controlled release of tetracycline hydrochloride, Mater. Lett. 141 (2015) 180–186. [6] M. Naghibzadeh, M. Adabi, Evaluation of effective electrospinning parameters controlling gelatin nanofibers diameter via modelling artificial neural networks, Fibers Polym. 15 (4) (2014) 767–777. [7] C. Ribeiro, V. Sencadas, J.L.G. Ribelles, S. Lanceros-Méndez, Influence of processing conditions on polymorphism and nanofiber morphology of electroactive poly(vinylidene fluoride) electrospun membranes, Soft Mater. 8 (3) (2010) 274–287.
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