Controlled Release of Ciprofloxacin from Core ... - ACS Publications

Article pubs.acs.org/molecularpharmaceutics

Controlled Release of Ciprofloxacin from Core−Shell Nanofibers with Monolithic or Blended Core Špela Zupančič,†,‡ Sumit Sinha-Ray,† Suman Sinha-Ray,†,§,∥ Julijana Kristl,‡ and Alexander L. Yarin*,†,⊥ †

Department of Mechanical and Industrial Engineering, University of Illinois at Chicago, Chicago, Illinois 60607-7022, United States Faculty of Pharmacy, University of Ljubljana, Aškerčeva 7, 1000 Ljubljana, Slovenia § Corporate Innovation Center, United States Gypsum, 700 US 45N, Libertyville, Illinois 60048, United States ∥ Department of Materials Science and Engineering, Indian Institute of Technology, Indore, Madhya Pradesh 452017, India ⊥ College of Engineering, Korea University, Seoul, South Korea ‡

S Supporting Information *

ABSTRACT: Sustained controlled drug release is one of the prominent contributions for more successful treatment outcomes in the case of several diseases. However, the incorporation of hydrophilic drugs into nanofibers, a promising novel delivery system, and achieving a long-term sustained release still pose a challenging task. In this work we demonstrated a robust method of avoiding burst release of drugs and achieving a sustained drug release from 2 to 4 weeks using core−shell nanofibers with poly(methyl methacrylate) (PMMA) shell and monolithic poly(vinyl alcohol) (PVA) core or a novel type of core−shell nanofibers with blended (PVA and PMMA) core loaded with ciprofloxacin hydrochloride (CIP). It is also shown that, for core−shell nanofibers with monolithic core, drug release can be manipulated by varying flow rate of the core PVA solution, whereas for core−shell nanofibers with blended core, drug release can be manipulated by varying the ratios between PMMA and PVA in the core. During coaxial electrospinning, when the solvent from the core evaporates in concert with the solvent from the shell, the interconnected pores spanning the core and the shell are formed. The release process is found to be desorption-limited and agrees with the two-stage desorption model. Ciprofloxacin-loaded nanofiber mats developed in the present work could be potentially used as local drug delivery systems for treatment of several medical conditions, including periodontal disease and skin, bone, and joint infections. KEYWORDS: sustained drug release, ciprofloxacin hydrochloride, coaxial electrospinning, core−shell nanofibers, antibiotic, phase separation

1. INTRODUCTION Polymer nanofibers have been in focus due to their special internal architecture. Their remarkable features, including nanoscale diameter, high surface area-to-volume ratio, and porous structure, enable an improved mechanical performance and flexibility compared to the other forms of the same material.1−3 Polymer nanofibers are applied in filtration,4 packaging,5 and electronics,6 and revealed an immense potential in biomedical applications.2,7−9 The resemblance of polymer nanofibers to the extracellular matrix revealed them as an excellent support for cell growth, which makes them a powerful tool for tissue engineering.10 Apart from cell proliferation and adhesion, nanofibers can also be useful for simultaneous release of different therapeutic agents, such as small-molecule drugs, peptides, proteins, and pDNA.7,8,11−13 As a result, different active ingredients and polymers, both synthetic and natural, were explored as potential drug delivery systems useful for treatment of various diseases, such as chronic wounds,8 periodontal disease,14,15 and cancer.9 The rate of drug release and quantity of the released drug are crucial for © XXXX American Chemical Society

successful treatment of specific diseases. Therefore, development of nanofibers with controlled drug release features poses a challenging task. Electrospinning is one of the common techniques used to form nanofibers from polymer solution containing compatible drugs. Over the years monolithic and blend-made nanofibers loaded with drugs were investigated and revealed as potentially promising drug delivery systems. Release from blended nanofibers composed of hydrophobic and hydrophilic polymers can be tuned over 18 days with proper selection of polymers and their blend ratios.16 However, several obstacles had been encountered.17,18 For example, nonuniform distributions of drug molecules in nanofibers and even drug crystallization at the surface were reported.19−21 Accordingly, drug enrichment at the nanofiber surface resulted in undesirable burst release in Received: January 14, 2016 Revised: February 25, 2016 Accepted: March 7, 2016

A

DOI: 10.1021/acs.molpharmaceut.6b00039 Mol. Pharmaceutics XXXX, XXX, XXX−XXX

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Molecular Pharmaceutics

Figure 1. Core−shell nanofibers with monolithic or blended core. In the case of core−shell nanofibers with monolithic core the ratio between the flow rates of polymer solutions used to form the core and shell was varied. By contrast, different compositions of the core solutions were used for preparation of the core−shell nanofibers with blended core, while the core-to-shell flow rate ratio was fixed as 1:4. It should be emphasized that the blended core consists of PMMA droplets dispersed in the continuous PVA phase.

multiple cases.21,22 A sustained drug release can be achieved using nanofibers composed of hydrophobic polymers, such as poly(lactic-co-glycolic acid), polylactic acid, and polycaprolactone.22 These polymers can only be dissolved in organic solvents, where some hydrophilic drugs are insoluble or sensitive proteins denature.23 In addition, hydrophobic polymers often lack an interaction with cells or other tissue compartments.17 In 2003, a novel class of nanofibers with core−shell structure was introduced when the coaxial electrospinning process was proposed.24 Core−shell nanofibers hold great potential of improving loading of hydrophilic drugs and proteins and could reveal a sustained release without any burst effect. The core− shell nanofibers can be formed not only by coaxial electrospinning of two separate polymer solutions using coaxial nozzles24 but also by electrospinning of specially prepared emulsions or phase-separated polymer solutions though a single nozzle.12,25 Electrospun nanofibers are affected by several parameters, such as the solution composition, processing, and ambient conditions.17 In coaxial electrospinning, in addition, the interactions between the core and shell polymer solutions and their flow rates can significantly affect the nanofiber properties.18 It should be emphasized that coaxial electrospinning or emulsion electrospinning do not necessarily result in core−shell nanofibers.26 It was revealed that an appropriate combination of the material and governing parameters can result in successful incorporation of several active substances into nanofibers, including hydrophilic drugs and proteins.7,12 Furthermore, in different studies fast27 or sustained drug release,7 drug release without any burst effect7,28 or a two-stage drug release29,30 was observed. Long-term hydrophilic drug release from nanofibers is a difficult task. It is desired for treatment of various diseases, including treatment of infections with antibiotics, wound healing with growth factors, local delivery of chemotherapeutics, and bone regeneration.31 Sustained release is a powerful method to decrease fluctuation of drug in blood levels, reduce adverse effects, lessen frequency in dosing, and thus enhance patient convenience and compliance.32 However, a direct comparison of release results from nanofibers of different studies is quite challenging due to the lack of release protocol standardization and incorporation of different drugs which possess special characteristics. To the best of our knowledge, the effect of different polymer solutions and process parameters in the core−shell nanofiber morphology on sustained drug release has not been explored in full detail, albeit such study would be beneficial for tuning of the long-term drug release.

This work aims at development of a new type of biocompatible CIP-loaded core−shell nanofiber mats, which would minimize the initial burst during the first day and sustain gradual drug release for at least 14 days at the local site of infections. CIP, hydrophilic drug, was chosen due to its high activity against diverse microorganisms, which is based on the inhibition of enzymes needed for the bacterial DNA replication. It is used for the treatment of complicated and uncomplicated urinary tract infections, skin and skin-structure infections, bone and joint infections, and periodontal disease.33,34 Moreover, such local sustained CIP delivery using nanofiber mats holds great potential for better treatment with a diminished possibility for development of bacterial resistance.15 For this aim, we systematically investigate the effect of the composition and preparation parameters on the long-term sustained drug release. First, CIP-loaded core−shell nanofibers with a hydrophobic polymer, poly(methyl methacrylate) (PMMA), used as a shell and hydrophilic poly(vinyl alcohol) (PVA) used as a core were coaxially electrospun at various flow rate ratios for the core and shell, and the effect on the nanofiber morphology and drug release profile was evaluated. Note that both polymers are biocompatible and are widely used in medical practice as drug delivery systems or implants.35−37 It should be emphasized that PMMA is FDA approved as soft tissue fillers, also known as injectable implants, dermal fillers, or wrinkle fillers. This material is also used in other medical devices, such as bone cement and intraocular lenses. Therefore, its biocompatibility has been already proven. The possible alternative materials, for example, polylactic acid (PLA), poly lactic-co-glycolic acid (PLGA), and poly L-lactic acid (PLLA), are biocompatible and biodegradable but, unfortunately, are prohibitively costly, while ethyl cellulose and cellulose acetate present a special challenge for production of core−shell nanofibers due to their relatively low viscoelasticity.38 Accordingly, it was decided to explore PMMA as a biocompatible hydrophobic polymer potentially useful for sustained drug release. Second, for the first time different phase-separated blend solutions were used as a core to increase control over the drug release profile. Finally, the drug release mechanism was explored and the drug release from blended and core−shell nanofibers with monolithic and blended cores was examined in comparison.

2. MATERIALS AND METHODS 2.1. Materials. PVA (78 kDa, 88% hydrolyzed) was obtained from Polysciences Inc., U.S.A. PMMA (Mw = 996 kDa), acetic acid (99%), formic acid (95%), and potassium B


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Table 1. Solution Compositions and Electrospinning Parameters Used for Preparation of the Core−Shell Nanofibers and the Resulting Calculated Nanofiber Composition Considering Flow Rate mass of polymer (g) in 10 g of solution core

flow rate (mL/h)

shell

sample name csPVA csPVA csPVA csPVA

1:3 1:4 1:5 30%

csPVA 70% csPVA 90%

PVA PVA PVA PMMA PVA PMMA PVA PMMA PVA

0.60 0.60 0.60 0.42 0.18 0.18 0.42 0.06 0.54

nanofiber composition (wt %)

core

shell

voltage (kV)

CIP

PVA

PMMA

PMMA PMMA PMMA PMMA

0.8 0.8 0.8 0.8

0.40 0.30 0.24 0.30

1.2 1.2 1.2 1.2

11 12 12 11

1.0 0.8 0.6 0.8

19.8 15.7 13.0 4.7

79.2 83.5 86.4 94.5

PMMA

0.8

0.30

1.2

12

0.8

11.0

88.2

PMMA

0.8

0.30

1.2

12

0.8

14.1

85.1

blended core the flow rate ratio was constant (1:4) but the blend composition in the core was varied. The details on the electrospinning parameters employed for preparation of different core−shell nanofibers are presented in Table 1. 2.3. Electrospinning of Monolithic and Blended Nanofibers. For the sake of comparison, the monolithic and blended nanofibers were prepared using the same solvent mixture as for the core−shell nanofibers. The 8 wt % PMMA and 8 wt % PVA solutions were used for electrospinning of monolithic nanofibers. Solutions for the blended nanofibers were composed of different percentages of PVA in nanofibers (30, 70, and 90%) and prepared by dissolving 0.6 g of PMMA and 0.257 g of PVA, 0.24 g of PMMA and 0.56 g of PVA, and 0.08 g of PMMA and 0.72 g of PVA with the addition of 1% CIP (compared to the mass of all solid compounds added to solution) in 10 g of solution, respectively. Electrospinning was used to produce monolithic PMMA or PVA nanofibers and blended nanofibers from the above-mentioned solutions using single nozzle. The voltage was in the 10−12 kV range, the distance from the needle to the collector was 15 cm, and the flow rate was 1.2 mL/h. In all the cases, a 22G needle was used. Monolithic nanofibers are denoted as nf PVA or nf PMMA, whereas nanofibers with incorporated 70% of PMMA and 30% of PVA are denoted as PMMA:PVA (70:30). 2.4. Characterization of Polymer Solutions and Nanofibers. To examine the presence of two polymer phases in the PMMA:PVA blends and CIP partition between both phases, a single drop was located between two glass slides and observed using a fluorescent microscope under visible light, as well as fluorescence at the excitation wavelength of 360 nm and the emission wavelength of 447 nm (Evos FL, Life Technologies Corporation, U.S.A.). The same fluorescent microscope was also used for the observations of core−shell nanofibers. All scanning electron microscopy images were obtained using a high resolution field emission microscope JEOL JSM-6320F (Research Resource Center, UIC) after sputter coating. It should be emphasized that coating thickness was set in the Cressington sputter coater by preset calibration. Pt/Pd coating was varied, being either 7 or 8 nm (based on the sample thickness) for proper visualization to minimize the effects related to the buildup of static charges. Nanofiber cross sections were observed before and after CIP release study. Fiber diameters and shell and core thicknesses were measured from SEM images using the image analysis software (ImageJ, National Institutes of Health, U.S.A.). 2.5. Fourier Transform Infrared Spectroscopy (FTIR). The FTIR spectra of pure substances CIP, PMMA, and PVA

phosphate monobasic were obtained from Sigma-Aldrich, U.S.A. Sodium hydroxide (NaOH) and acetonitrile were purchased from Fisher Scientific, U.S.A. CIP was delivered by Alfa Aesar, U.S.A. All the chemicals were used as received without any further purification or change. 2.2. Formation of Core−Shell Nanofibers with Monolithic and Blended Core. Coaxial electrospinning was employed to produce core−shell nanofibers with encapsulated CIP in either monolithic or blended core (Figure 1). Shell was in all cases formed from PMMA using the 8 wt % PMMA solution in acetic and formic acid in the mass ratio 3:1. The nanofiber core was formed from the 6 wt % solution of PVA in the same solvent mixture in the case of monolithic core. Blends of both PMMA and PVA in different ratios (70:30, 30:70, and 10:90) dissolved in the same solvent mixture as that for the shell were used to form the fiber core. In all the core solutions 5 wt % of CIP (compared to the mass of the core polymers together) was dissolved. The details on the mass percentage in the core and shell solutions are presented in Table 1. Six different types of core−shell nanofibers are denoted using the following nomenclature. Core−shell nanofibers with monolithic PVA core prepared with variation of flow rate of PVA solution and constant flow of PMMA solution to obtain the core-to-shell flow rate ratios of 0.4 mL/h:1.2 mL/h (1:3), 0.3 mL/h:1.2 mL/h (1:4), and 0.24 mL/h:1.2 mL/h (1:5) are denoted as csPVA 1:3, csPVA 1:4, and csPVA 1:5, respectively. Blended core−shell nanofibers with blended core composed of 30% of PVA and 70% of PMMA, 70% of PVA and 30% of PMMA, and 90% of PVA and 10% of PMMA are denoted as csPVA 30%, csPVA 70%, and csPVA 90%, respectively. During coaxial electrospinning, a custom-made coaxial needle setup was utilized for generating the core−shell polymer jets.22,39 The inner needle (25G) had an inner diameter of 0.260 mm and an outer diameter of 0.514 mm, whereas the outer needle (18G) had an inner diameter of 0.838 mm, which gives 0.324 mm distance for the flow of polymeric shell solution. A voltage of 11−12 kV was sustained by a dc high voltage power supply (Gamma High Voltage Research, Inc., Ormond Beach, FL). The distance between the electrode submerged in the coaxial needle and grounded rotating disk used as a counter-electrode was 15 cm. The disk nanofiber collector rotated with an angular speed of 250 rpm. The ambient humidity and temperature varied in the 47−64% and 21−23 °C range, respectively. The ratio between the polymer solution flow rates used to form the core and shell was varied in the case of core−shell nanofibers with a monolithic core. On the other hand, in the case of the core−shell nanofibers with C


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Figure 2. Schematic of nanopores inside polymer matrix of (a) a monolithic and (b) a core−shell nanofiber. Here 2b denotes the nanopore diameter and cw denotes the drug concentration over the nanofiber wall determined by desorption. In panel (a) L denotes the nanopore length, and in panel (b) L denotes the length of the core section of the nanopore, whereas S denotes the length of the shell section of the nanopore.

(OriginLab corporation, USA). In addition, the experimental data were matched with the desorption-limited model using OriginPro 2015 employing the Orthogonal Distance Regression to iteratively adjust the parameters. Each function was evaluated using correlation coefficient (R2).

and monolithic nanofibers from PMMA and PVA with or without drug were characterized with FTIR spectrometer with an attenuated total reflectance accessory (Nexus, Thermo Nicolet, Madison, WI, USA). Spectra in the 600 to 3600 cm−1 range with the resolution of 8 cm−1 were measured. Each recorded spectrum was an average of 16 scans. 2.6. CIP Release from Nanofibers. The release study was performed as reported by this group previously.16 In brief, the release studies for all nanofiber mat samples with dimensions 1.5 cm × 6 cm and weight of 15−20 mg were performed in glass vials containing 10 mL of 50 mM phosphate buffer with pH 7.4 maintaining sink conditions. The buffer solutions with the immersed nanofiber samples were shaken at 120 rpm on an orbital rotator (Kangjian KJ-201BD, Jiangsu, China) at room temperature. At different time moments, 0.5 mL solution samples were periodically withdrawn from a vial and replaced by the same volume of fresh buffer. The solution samples were diluted with 0.5 mL of acetonitrile to precipitate PVA (emerging from a partial dissolution of the PVA from nanofibers) because the CIP fluorescence intensity decreases in the presence of dissolved PVA. The concentrations of the released CIP in the samples of the buffer solution were measured using spectrophotometer Tecan Infinite M200 Pro (Tecan Group Ltd., Männedorf, Switzerland) at an excitation wavelength of 280 nm and an emission wavelength of 450 nm. Small traces of the acetic and formic acid salts formed with CIP (formed in the solutions used for coaxial electrospinning and electrospinning) caused an increase in the fluorescence. Therefore, a standard stock solution was prepared by dissolving CIP in a mixture of the acetic and formic acids at room temperature. This standard stock solution was neutralized after 5 h by phosphate buffer with a higher pH to reach at the end pH 7.4. Then, a calibration curve was prepared by further dilution with phosphate buffer and acetonitrile [50% (v/v)]. The release results were presented in terms of the cumulative release release (%) =

Mt × 100 M∞

3. THEORETICAL BACKGROUND In the previous work by the present group,16,40−42 it was shown that release from monolithic nanofibers is desorption-limited instead of being diffusion-limited. In refs 16, 40, and 41, it was shown that drug release from nanofiber mats is a surface phenomenon, where drug or any other model compound desorbs from the outer and/or nanopore surfaces of nanofibers (cf. Figure 2a). In the present work the core−shell nanofibers were formed from the core and shell materials dissolved in the same solvent. As a result, during coaxial electrospinning, when the solvent from the core evaporates in concert with the solvent from the shell,43 the interconnected pores spanning the core and the shell are formed. However, unlike monolithic nanofiber, only the core part will be loaded with drug, whereas the shell part will be devoid of drug (Figure 2b). Following ref 41, it can be shown that the entire nanopore (with length of about 1 μm and diameter of about 10 nm) will be filled with the buffer solution almost instantaneously after the fibers are submerged into a buffer medium in a vial. Then, a two-stage release process will begin. First, the drug-loaded core part of the nanopore will release drug by desorption from the surface into the buffer solution on the time scale τr1 ∼ L2/Deff ∼ 1−100 h, where Deff is the effective diffusion coefficient for the desorption-limited release calculated in ref 41. In addition to that, diffusion occurs on the time scale τ ∼ (L + S )2/D ∼ 10 s, where (L + S ) is the entire pore length (∼1 μm) and D is the diffusion coefficient of drug in the buffer solution (D ∼ 10−5 cm2/s). Such a huge disparity between the values of τr1 and τ implies that τ is negligibly small compared to τr1, and as soon as the drug desorbs from the core, it is released to the buffer solution bath almost instantaneously, similarly to the release from monolithic nanofibers.41 It should be emphasized that the effective diffusion coefficient is Deff = Dcw0b/ρsd0, where an initial bulk concentration of drug in the buffer solution near the nanofiber surface is denoted cw0, the initial drug surface density is ρsd0, and b is the nanopore cross-sectional radius. The value of cw0 is determined by the limiting process of drug desorption from the nanofiber surface, which results in the values Deff ≪ D (cf. ref 41 and the results in Tables 2 and 3 below). Since in the present case of the core−shell nanofibers (with blended core or monolithic cores) the core contains water-

(1)

as a function of release time, where Mt is the amount of CIP released at time t and M∞ is the total amount of CIP in the nanofiber mat. All the release experiments were conducted at least thrice. 2.7. Statistic Analysis. The data were expressed as mean ± standard deviation (SD). The results were statistically analyzed using one way analysis of variance (ANOVA) followed by Tukey’s post hoc test using software OriginPro 2015 D


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potential interaction between CIP and the host polymers could create any chemical bonds. To elucidate that, the FTIR spectrum of pristine CIP is shown in Figure 3a; in Figure 3b are

soluble PVA, the latter will leach out in parallel with the drug release process and create more pores. The PVA leaching and the accompanying drug release from the surfaces of the newly formed pores are also desorption-limited with a characteristic time scale τr2,16,40 which determines the following two-stage drug release dependence on time, ⎡ ⎡ ⎛ π 2 t ⎞⎤ ⎛ π 2 t ⎞⎤ Mt = α1⎢1 − exp⎜ − ⎟⎥ ⎟⎥ + α2⎢1 − exp⎜ − Md0 ⎝ 8 τr2 ⎠⎦⎥ ⎝ 8 τr1 ⎠⎦⎥ ⎣⎢ ⎣⎢ (2)

where α1 and τr1 denote the nanoporosity factor and the characteristic time of drug release from the pre-existing pores, and α2 and τr2 denote the nanoporsity factor and the characteristic time of drug release driven by PVA leaching.16,40 According to ref 41 the nanoporosity factor for the first stage of the process, α1 = Msd0/(Msd0 + Mbd0) < 1, with Msd0 and Mbd0 being the initial amount of drug at the nanofiber surfaces (in particular, at the nanopore surfaces) and the initial amount of drug embedded in the fiber bulk, respectively; the total initial amount of drug in the fiber is, thus, Md0 = Msd0 + Mbd0. For the second stage of the process α2 = M′sd0/(Msd0 + Mbd0) < 1, with M′sd0 being the amount of drug at the nanofiber surfaces (in particular, at the nanopore surfaces) exposed due to PVA leaching. Similarly,16,40 τri = L2/[Dcw0ib/ρsd0i], where i = 1 and 2 mark the pre-existing and the newly created pores, respectively, and cw0i and ρsd0i denote the corresponding initial drug concentration over the nanopore surface in the buffer solution determined by desorption and the initial surface density of the drug at the nanopore surface. Accordingly, cw0i = k0i exp(−Ei/ RT)ρsd0i/ρsp, where k0i are the pre-exponential factors, Ei are the desorption enthalpies of CIP, R is the universal gas constant, T is the temperature, and ρsp is the surface densities of the polymer matrix and the leachable polymer. In all the estimates below the pre-exponential factors k01 and k02 are taken as 10−3 g/cm3, and the temperature T = 300 K, the pore cross-sectional radius b is of the order of 10 nm, and the diffusion coefficient of CIP in the buffer solution D is of the order of 10−5 cm2/s.41 Pores are tortuous,44 and one can safely assume that the pore length is of the order of the nanofiber diameter, which was in the range of 500−1000 nm. The individual polymer nanofibers are composed of polymer lamellae connected by amorphous regions, with the interstitial gap between the lamellae being about the effective diameter of an individual polymer lamella. In refs 44 and 45 it was revealed that the polymer lamellar diameter is ∼10 nm. Therefore, the values of b = 10 nm and L = 1000 nm represent plausible orderof-magnitude estimates. Note that considering a tortuous pore as a random walk of step about 10b, one can find that the number of zigzags the nanopore makes while criss-crossing a nanofiber until it reaches the surface would be N = (L/10b)2 = 100, and the tortuous path of such a nanopore would be mostly along the fiber, which is expected, since electrospun nanofibers undergo strong uniaxial elongation.

Figure 3. FTIR spectra of CIP powder (a), PMMA powder and nanofibers with and without 1% CIP (b), and PVA granules and nanofibers with and without 1% CIP (c).

FTIR spectra of PMMA powder, PMMA nanofiber, and PMMA nanofiber loaded with 1% CIP, whereas Figure 3c shows the spectra of PVA powder, PVA nanofiber, and PVA nanofiber loaded with 1% CIP. A detailed interpretation of the FTIR spectra is facilitated by ref 46. Figure 3b shows that the addition of CIP does not distort the FTIR spectrum of CIPloaded PMMA nanofibers as compared to the FTIR spectrum of PMMA nanofibers without CIP. This reveals that there is no observed interaction between CIP and PMMA. Figure 3c shows that PVA has a broad peak between 3200 and 3550 cm−1 owing to hydroxyl group.46 The addition of CIP shifted the peak related to hydroxyl stretching from 3393 to 3399 cm−1. This can be attributed to the hydrogen bonding formed between PVA and CIP as was reported in ref 46, which reveals that there are weak interactions between PVA and CIP. The latter implies that the CIP release from all the core−shell nanofibers formed in the present work is expected to follow the desorption-limited mechanism described by eq 2. 4.2. PVA−PMMA Core−Shell Nanofibers with Monolithic Core. 4.2.1. Morphology of PVA−PMMA Core−Shell Nanofibers with Monolithic Core. In the present study, a drug reservoir system in the form of the core−shell nanofibers was

4. RESULTS AND DISCUSSION 4.1. Interaction of “Host” Polymers and CIP. In the present work, in the case of the core−shell fibers with monolithic PVA core, the latter was used as a “host” polymer for CIP. Similarly, in the case of the core−shell fibers with blended core, PVA and PMMA were used as “host” polymers for CIP. It was extremely important to understand whether a E


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Figure 4. SEM images of core−shell nanofibers with monolithic core for different core flow rates: (a) csPVA 1:3, (b) csPVA 1:4, and (c) csPVA 1:5. Image set (1) shows the overall views of the dry samples. The image sets (2) and (3) show the cross sections of the core−shell nanofibers before and after CIP release, respectively. The arrows point at the longitudinal holes throughout the nanofibers.

nanofibers were obsereved in ref 49 and collapsed core−shell nanofiber were reported in ref 47. The formation of the sixshaped core−shell nanofibers observed in the present work could be related to a faster solidification rate of the shell compared to the core, as was shown for hollow core−shell nanofibers.43 The present results also imply that as the solvent evaporates through the cross section, the polymer in the core could be pulled toward the shell resulting in a hole in the middle (panels a2−c2 in Figure 2). Change of the cross-section shape during solvent evaporation could also result in nanofiber touching each other and a partial merging. The DSC thermogram (cf. Figure S1) reveals that the internal nanofiber structure could relax due to merging, which diminished the degree of crystallinity compared to the parent polymers. It should be emphasized that a diminishing degree of crystallinity can be observed from the DSC thermogram in Figure S1. It shows that while pristine PVA polymer used in this work possesses well-defined melting and glass transition temperatures, PVA nanofibers do not reveal any clear melting peak. Note that while the glass transition elucidates the amorphous nature of a polymer, the presence of a melting point signifies polymer crystallinity.50 In the DSC thermogram the absence of a clear melting point for PVA nanofibers corresponds to the loss of crystallinity. However, PMMA in the shell and PVA in the core did not completely intertwine, which was observed by two separated characteristic glass transition temperatures of PVA and PMMA (cf. Figure S1). Different structures of nanofibers formed with different flow rates determined changes in the average nanofiber diameter. Such nanofibers as csPVA 1:5 had the average dimeter of 1010 ± 165 nm compared to both core−shell nanofibers csPVA 1:3 and csPVA 1:4, which had significantly smaller (p < 0.05) average diameter, 705 ± 140 nm and 670 ± 140 nm, respectively. The cross sections of all nanofibers were compared before and after CIP release study. Water-soluble PVA was partially dissolved in the surrounding buffer resulting in the increase in the middle holes from 110 ± 40 nm to 355 ± 85 nm in the case

designed. CIP was incorporated in the hydrophilic PVA matrix in the core, and PMMA was chosen as a hydrophobic shell serving as a regulating barrier for sustained drug release. It is important to elucidate the effect of core and shell sizes on the nanofiber morphology and consequently on the CIP release. The shell and core thickness can be effectively controlled by varying the flow rate ratios of both the constituents. Therefore, in the present case, the flow rate used to form the shell was constant at 1.2 mL/h to obtain stable coaxial electrospinning with no solution dripping and a stable jetting Taylor cone at the spinneret orifice exit with continuous jet ejection. On the other hand, the flow rate used to form the core was varied to achieve the core−shell flow ratios of 1:3, 1:4, and 1:5 (Table 1). A strong effect of the flow rate ratio on the nanofiber morphology is clearly seen in Figure 4. Even though the SEM images with the overall views of the nanofiber show relatively smooth and uniform nanofibers (Figure 4a1−c1), the cross-sectional images of the nanofibers before any release study (Figure 4a2−c2 and Figure 4a3−c3, respectively) show that the cross sections reveal unique ribbonlike (Figure 4b and 4c) and six-like structures (Figures 4a and 4b). In many cases, the fibers also have revealed a throughout hole in the middle. It can be also seen that sometimes the nanofibers have merged together. These observations can be interpreted as follows. During coaxial electrospinning the nanofiber skin gets solidified first on the surface due to the excessive solvent evaporation. However, the core still contains solvent, which is exactly the same as the solvent in the shell. The solvent from the core evaporates through the nanopores already opened in the shell, which decreases pressure in the core. The nanofibers collapse under the action of the higher outside atmospheric pressure if the shell is too thin.47,48 This results in the ribbon-like nanofibers or, due to the buckling of the shrinking ribbon, even in the six-like nanofibers. Nanofiber collapses have already been reported in the literature. For example, monolithic PMMA nanofibers may also collapse and acquire a bow-tie shape.16 In addition, the monolithic ribbon F


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the first 30 min, whereas at a higher ratio of 1:4, the drug was released in a two-stage process, i.e., 62% of CIP was released as the burst effect in 1 day and in the next 31 days 15% of CIP was released in addition. The optimal ratio between the core and shell was 1:5, where a sustained CIP release was obtained for 28 days without any burst effect. Figure 6 shows, that irrespective of the system, the overall drug release saturated at 70−95%. If the release were diffusiondriven, the entire drug embedded in the polymer matrix would have been released. This release saturation below 100% points at the desorption-limited mechanism of drug release. Accordingly, eq 2 was fitted to the data, as shown in Figure 6. From Figure 6a,b, it can be seen that the desorption-limited theory, eq 2, describes the CIP release phenomenon very well. The corresponding parameters involved in eq 2 are listed in Table 2. Figure 6 shows that, with a higher amount of PVA in the core (csPVA 1:3), a lower cumulative drug release was observed and also the release process saturated faster. This effect can be explained as follows. Core−shell nanofibers csPVA 1:3 were composed of 20% of PVA and 80% of PMMA, whereas in csPVA 1:5 13% of PVA and 87% of PMMA. In the first case, the core flow rate was higher and in some sections the shell did not cover the core resulting in the core−shell nanofiber imperfections observed as holes after nanofiber immersion in water (Figure 7). In addition, due to the same solvent system for the core and shell there was also an increased possibility that PVA diffused to the surface and acted as in situ prepared porogen. Both reasons combined with the hollow core of csPVA 1:3 and cracks exposed some PVA to the nanofiber surface (Figure 7). Accordingly, these nanofibers released drug within a few hours as was also observed in blended PVA and PMMA nanofibers with 15% of PVA.16 PVA on the surface increased wettability of the nanofiber surface, CIP desorbed from PVA and simultaneously, PVA started leaching out, which further enhanced water imbibition in the core. As soon as water entered in previously encapsulated area, CIP desorption began from the newly formed space until it reached saturation. Some CIP diffused in PMMA during solvent evaporation and could not be desorbed resulting in 70% total release. In addition, the values of E1 and E2 were comparable between csPVA 1:3 (E1 = ∼31 kJ/mol and E2 = ∼44 kJ/mol, Table 2) and PMMA:PVA (85:15) (E1 = ∼32 kJ/mol and E2 = ∼44 kJ/mol).16 On the contrary, the csPVA 1:5 nanofibers with lesser amount of imperfections (observed in the SEM images) resulted in a fully

of csPVA 1:3. Accordingly, the wall thickness decreased from 165 ± 70 nm to 115 ± 40 nm, which suggests that the PVA deposit on the inner side of the hollow nanofibers was about 50 nm thick (panel a3 in Figure 4). Moreover, newly formed holes of the size 55 ± 18 nm in the sides of the ribbons were observed in csPVA 1:5 nanofibers after CIP release (panel c3 in Figure 4). In addition, CIP distribution in csPVA 1:5 nanofibers was studied using the fluorescence microscope (Figure 5). In all the

Figure 5. An overlay of fluorescent and bright-field images of CIPloaded core−shell nanofibers: (a) csPVA 1:3, (b), csPVA 1:4, and (c) csPVA 1:5. CIP is marked with blue color and arrows, and black lines present the edges of the core−shell nanofibers.

cases, CIP was uniformly distributed in the core of the core− shell nanofibers. Figure 5 presents an overlay of the fluorescent and bright-field images, where CIP appeared to be restricted within the core, which is distinguishable from the shell. In the case of csPVA 1:3 nanofibers the core is shifted from the fiber center (Figure 5a). The core in the csPVA 1:4 has been split, with its parts being displaced to both sides of the ribbon nanofibers, which is revealed by the fluorescence-less area seen in the middle of the nanofibers (Figure 5b). Unfortunately, the relatively low resolution of the fluorescent microscope limited the accurate measurement of the core and shell thicknesses, as can be seen in Figure 5c for the csPVA 1:5 nanofibers (with the approximately 100 nm thin shell determined using SEM). 4.2.2. CIP Release from PVA−PMMA Core−Shell Nanofibers with Monolithic Core. The in vitro CIP release from PVA/PMMA core−shell nanofibers with monolithic core was studied in the buffer solution with pH 7.4 to test their potential applications as a local drug delivery system. The different ratios between the core and the shell polymer solution flow affected CIP release, as is evident in Figure 6. Due to lower core flow rate a smaller core was seen in csPVA 1:5, which diminished the burst effect and facilitated the sustained drug release. In the case of the core-to-shell ratio 1:3, most of CIP was released during

Figure 6. CIP release profiles from core−shell nanofibers prepared with different flow-rate ratios between the core and the shell polymer solutions for (a) first 2 days and (b) total 30 days of release. The experimental data is shown by symbols. The release profiles were fitted with the predictions of the two-stage desorption model described by eq 2 (curves). G


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Table 2. Parameters of the Two-Stage Desorption Theory Determined in Figure 6 by Matching the Experimental Data for CIP Release from Different Core−Shell Nanofibers α1 (%) csPVA 1:3 csPVA 1:4 csPVA 1:5

58.1 ± 1.1 67.8 ± 1.1 58.1 ± 8.0

α2 (%) 12.3 ± 1.2 16.6 ± 2.2 36.7 ± 3.8

τr1 (h) 0.7 ± 0.6 6.1 ± 0.7 74.8 ± 10.4

τr2 (h)

Deff1 (cm2/s)

161.6 ± 53.0 388.7 ± 147.6 612.7 ± 297.7

−12

4.0 × 10 4.6 × 10−13 3.7 × 10−14

Deff2 (cm2/s) −14

1.7 × 10 7.1 × 10−15 4.5 × 10−15

E1 (kJ/mol)

E2 (kJ/mol)

30.6 36.0 42.3

44.2 46.4 47.5

Figure 7. Core−shell imperfections are seen as incomplete shell, cracks, and open ends observed on the sample csPVA 1:3 after the release study. Scale bars in all images are 500 nm.

Figure 8. SEM images of (a) csPVA 30%, (b) csPVA 70%, (c) csPVA 90%, and (d) PMMA:PVA (70:30) nanofibers. Subscripts 1 in the panel notation correspond to the fibers before CIP release, and subscripts 2 in the panel notation corresponds to the fibers after CIP release.

sustained release. These nanofibers did not have PVA exposed directly to the nanofiber surface as was observed in blended nanofibers with either 5 or 10% of PVA or csPVA 1:3, and therefore did not result in the burst effect. Nanopores in csPVA 1:5 were immediately filled with water and desorbed the assessable CIP from the core part of nanopores. Tortuous path and space shortage increased the time needed for PVA chains’ disentanglement, dissolution’ and diffusion from the core. Therefore, PVA leaching slowly allowed desorption of CIP from previously inaccessible nanofiber bulk. To conclude, it is possible to control the rate and amount of the drug released from the core−shell nanofiber mats by simply changing the flow rates used during the fiber formation. 4.3. PVA−PMMA Core−Shell Nanofibers with Blended Core. 4.3.1. Morphology of PVA−PMMA Core−Shell Nanofibers with Blended Core. PVA and PMMA blended nanofibers can release CIP for more than 14 days as was previously reported, and the release rate can be varied by percentage of PVA in nanofibers.16 Similarly, PVA−PMMA core−shell nanofibers are capable of releasing CIP for more than 14 days, whereas the initial burst can be avoided and the process sustained over almost 30 days without sacrificing the total amount of drug incorporated in the nanofibers, as was demonstrated in the previous subsection. Accordingly, it is logical to explore whether the drug release can further be controlled over a longer period of time by using blended core

formed with PMMA and PVA blend. The SEM images of such core−shell nanofibers with blended core are shown in Figure 8. In addition, blended nanofibers from PMMA and PVA with the lowest amount of PVA [PMMA:PVA (70:30)] were chosen for comparison (Figure 8d). Figure 8a−c shows that as the PVA content in the core increases, the nanofiber begins losing its circular structure, which is related to the void opening in the fiber middle followed by collapse, as described before. Accordingly, the average nanofiber diameter significantly (p < 0.05) decreased from 825 ± 200 nm in the case of csPVA90% to 635 ± 160 nm in the case of csPVA 30%. The blend PMMA:PVA (70:30) resulted in circular nanofibers with a smooth surface and an average diameter of 575 ± 125 nm. Figure 8a−c shows that the nanofibers also merged with each other. SEM images in Figure 8 suggested that the blended polymers PVA and PMMA were immiscible, even though being dissolved in the same solvent. To verify this fact the blended solution prior to coaxial electrospinning was observed under fluorescent microscope. The mixtures of PMMA and PVA in ratio 30:70 (Figure 9a) and 70:30 (Figure 9b) resulted in phase-separated solutions, where in the first case small PMMA droplets were dispersed in the PVA continuous phase, whereas in the second case larger PVA droplets were dispersed in the PMMA continuous phase. The phase separation in the blends was a consequence of thermodynamic incompatibility of linear H


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Figure 9. Optical microscopic images of two-phase blend separation of (a) 30% of PMMA and 70% of PVA, and (b) 70% of PMMA and 30% of PVA in solution (the left-hand side column). The corresponding distribution of fluorescent CIP in both phases observed under fluorescent microscope (the right-hand side column), where the brighter color presents higher concentration of CIP.

polymers PVA and PMMA. CIP was preferentially distributed in the PVA phase, as is evident from a stronger CIP fluorescence in PVA (a brighter blue color in Figure 9) compared to that in PMMA phase (a darker blue color in Figure 9). Hydrophilic PVA with its several hydroxyl groups offered a stronger hydrogen bonding to CIP, which is a hydrophilic molecule with 2 hydrogen donors and 6 hydrogen acceptors [calculated using Advanced Chemistry Development (ACD/Laboratories) Software V11.02], compared to hydrophobic PMMA. It should be emphasized that the FTIR spectra (Figure 3) have also revealed a higher affinity of CIP to PVA than of CIP to PMMA owing to hydrogen bonding stemming from the hydroxyl group. Accordingly, a higher concentration of CIP in PVA phase compared to that in the PMMA phase was also expected in nanofibers, albeit some drug molecules could be transferred into the PMMA shell with solvent during the evaporation and solidification process. 4.3.2. CIP Release from PVA−PMMA Core−Shell Nanofibers with Blended Core. The release profiles produced by blended nanofibers and core−shell nanofibers with blended core were quite different. Blended nanofibers PMMA:PVA (70:30) released almost 90% of the drug in first 4 h due to a high content of hydrophilic polymer on the nanofiber surface (Figure 10a). In addition, the release from the blended PMMA:PVA (30:70) and PMMA:PVA (10:90) nanofibers was even faster, reaching the maximum released amount in the first 2 h (not shown). On the other hand, the incorporation of nanofiber blends as a core in the core−shell nanofibers strongly sustained CIP release (Figures 10b and 10c) compared to the blend-alone nanofibers or core−shell nanofibers with monolithic core (Figure 10). Figures 10b and 10c show that the addition of 10% of PMMA to the PVA core resulted in a minimal burst effect and sustained CIP release over 24 days. The nanofibers csPVA 70% with blended core were releasing drug without any burst effect for 14 days. By comparison, csPVA 30% nanofibers released very slowly reaching 25% release in 24 days. The choice of polymer blend in the core presents a useful tool for the drug release engineering. As was previously shown,

Figure 10. CIP release profiles from (a) monolithic PMMA:PVA (70:30) nanofibers and (b) core−shell nanofibers prepared at the core flow rate 0.3 mL/h and shell 1.2 mL/h; with monolithic core:csPVA 1:4, and with different composition ratio PVA and PMMA: csPVA 30%, csPVA 70%, and csPVA 90% for the first 3 days and (c) for the first 25 days. The experimental data is shown by symbols. The release profiles were fitted with the predictions of the two-stage desorption model described by eq 2 (curves).

an increased amount of PVA in blended nanofibers resulted in a faster release of CIP in water.16 In addition, also a higher content of PVA in core−shell nanofibers reported in section 4.2 can significantly increase the rate of the CIP release rate from the core. Incorporating PMMA, polymer with high desorption enthalpy with respect to CIP,16 is expected to impede the process. Figures 10b and 10c clearly shows that with the incorporation of a larger amount of PMMA in the core, the release rate can be tailored and the initial burst can be fully avoided. Note also that Figure 10c shows that, during even 25 days (the duration of the present experiment), csPVA 30% was still releasing CIP and did not reach its saturation. The experimental release data were fitted with the predictions of the two-stage desorption theory (eq 2) as shown in Figure 10 and Table 3. It can be seen that eq 2 can describe the release profile for csPVA 90%, whereas eq 2 could not successfully describe the release profiles csPVA 30% and csPVA 70%. The release profiles of csPVA 30% and csPVA 70% revealed sigmoid shapes uncharacteristic of the two-stage I


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Table 3. Parameters of the Two-Stage Desorption Theory Determined in Figure 10 by Matching the Experimental Data for CIP Release from Different Core−Shell Nanofibers α1 (%) PMMA:PVA (70:30) csPVA 30% csPVA 70% csPVA 90%

84.3 0.0 91.6 38.0

± ± ± ±

1.4 0.5 46.9 2.1

α2 (%) 11.1 100.0 0.0 58.3

± ± ± ±

1.2 73.1 195.8 1.9

τ1 (h) 0.4 0.3 211.6 10.6

± ± ± ±

τ2 (h)

0.2 0.0 118.3 1.5

7.9 2098.0 2939.6 137.8

desorption theory. The lag time of 2 days in the case of csPVA 70% and of 6 days for csPVA 30% was observed with the subsequent faster release. A continuous shell did not enable burst release, furthermore PMMA in the core additionally retained desorption of both CIP and PVA due to the incorporation as a core polymer blend, which phase separated. Wettability of polymer nanofibers deserves an additional comment. For any individual polymer or polymer blend, wettability can be characterized by measuring the contact angle of a sessile water droplet on a cast film,43 and for a uniform blend should follow the Cassie−Baxter equation. On the other hand, the presence of air in pores of nanofiber mats can dramatically increase the contact angle, and thus diminish their wettability by water for a long period of time.51 Therefore, CIP release from these core−shell nanofibers can be also altered due to the effect of nanofiber architecture. Although we tried to decrease the variations between samples with maintaining the same process conditions (the distance between nozzle and grounded rotating disk, applied voltage, the speed of rotating disk) and size of nanofiber mats used for release study, the nanofiber mat architecture, e.g., the inter-nanofiber porosity, could change to some extent and affect hold-up of air parcels inside the mat and, thus, the fiber contact with the surrounding buffer solution in the release experiments. As shown in the previous study,52 the air parcels between hydrophobic nanofibers can affect drug release. This phenomenon is not accounted for in the two-stage desorption equation, which implies that all fibers are equally surrounded by the buffer solution. However, a good agreement of the predictions of the two-stage model eq 2 with the experimental data in several cases in Figures 6 and 10 implies that the effect of an unaccounted air hold-up and nanofiber mat architecture was relatively small at least in several cases in the present experiments. 4.4. Comparison of CIP Release from Blended Nanofibers and Core−Shell Nanofibers with Monolithic and Blended Core Composed of PVA and PMMA and Loaded with CIP. In this section the time-dependent release of CIP from blended nanofibers is compared with the timedependent CIP release from the core−shell nanofibers with monolithic and blended cores composed of PVA and PMMA (Table 4). Because these nanofibers were developed for sustained release, release after 1 h and 1 day expresses the burst effect, whereas the release percentage in the first and second weeks reveals the capability for the sustained release. The comparison of similar amount of PVA in nanofibers (5%, 10%, and 15%, Table 4) but different composition clearly shows the trend that core−shell nanofibers strongly decrease burst effect of blended nanofibers, whereas blend in the core of the core−shell nanofibers additionally increases the gradual release due to the double sustained release mechanism. PMMA and PVA polymers present a good combination for sustained drug release. The amount of PVA in nanofibers and composition of nanofibers (blended, core−shell with mono-

± ± ± ±

1.7 17.0 0.0 10.1

Deff1 (cm2/s) 6.9 9.3 1.3 2.6

× × × ×

−12

10 10−12 10−14 10−13

Deff2 (cm2/s) 3.5 1.3 9.4 2.0

× × × ×

−13

10 10−15 10−16 10−14

E1 (kJ/mol)

E2 (kJ/mol)

29.2 28.5 44.9 37.4

36.7 50.6 51.4 43.8

Table 4. Time-Dependent Release of CIP from the Blended Nanofibers versus the Time-Dependent CIP Release from the Core−Shell Nanofibers with Monolithic and Blended Core Composed of PVA and PMMA additional CIP release (%) % PVA in nanofibers csPVA 1:3 csPVA 1:4 csPVA 1:5 csPVA 30% csPVA 70% csPVA 90% PMMA:PVA (95:5)a PMMA:PVA (90:10)a PMMA:PVA (85:15)a PMMA:PVA (70:30) a

1h

1 day

1 week

2 weeks

Core−Shell Nanofibers with Monolithic Core 20.0 53.3 6.9 6.2 15.8 22.3 43.9 8 13.0 1.5 17.0 50 Core−Shell Nanofibers with Blended Core 6.8 0.0 0.6 5.9 11.1 1.0 4.1 44.5 14.2 7.3 38.4 37.6 Blended Nanofibers 5 6.7 37.1 26.2

total

3.1 5.6 11.5

69.5 79.8 80

11.8 31.9 8.9

18.3 81.5 92.2

17.1

87.1

10

31.2

34.8

15.9

8

89.9

15

69.3

10.5

13.2

1.5

94.5

30

78.8

16.2

0

0

95

Data obtained from the previous work of this group.16

lithic or blended core) can dramatically alter drug release. However, the most appropriate nanofibers for the disease treatment are mostly determined by the type and location of the disease. Infections are treated with antibacterials, including CIP, where drug concentration at the local site must exceed a minimum inhibitory concentration usually for 5 to 14 days to inhibit growth of pathogen bacteria, decrease systemic adverse effects and possibility of occurrence of antibiotic resistance. In addition, these kinds of nanofibers can also be used for delivery of some other drugs. For example, two-phase drug release of nonsteroidal anti-inflammatory drugs, antihistaminic drugs, or antipsychotics with initial burst release relieve the symptoms in the shortest time possible and then the drug effect would be prolonged for several more hours to optimize the therapy. On the other hand, the gradual release without burst effect is necessary for drugs with strong adverse effects, including chemotherapeutics, hormones, and growth factors.34 It should be emphasized that controlled drug release from nanofibers with a lag time holds great potential for treatment of specific diseases according to chronotherapy.

5. CONCLUSION In the present work the problem of hydrophilic drug incorporation into nanofibers and the subsequent long-term controlled release was solved with a simple, inexpensive and easily scalable method, namely, by the development of core− shell nanofibers with monolithic or novel blended core loaded with CIP. The investigation of the effect of composition and J


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preparation parameters on the drug release has revealed two possibilities to achieve and tune sustained release from nanofibers. First, the variation of flow rate ratios between monolithic core and shell during coaxial electrospinning strongly affected nanofiber morphology and CIP release. Core−shell nanofibers with lower amount of PVA in the core (csPVA 1:5) produced with core:shell flow ratio 1:5 were able to prevent burst release and achieve long sustained drug release over 4 weeks. Second, phase-separated blend (PVA:PMMA = 70:30) as a core and PMMA as a shell of the core−shell nanofibers controlled CIP release for more than 25 days with a lag time of 2 days and could be further tuned with changing the PVA and PMMA ratios in the core. Finally, the experimental drug release profiles affected by the formation of the interconnected pores spanning the core and the shell reveal good agreement with the release mechanism associated with the two-stage desorption theory. In conclusion, ciprofloxacin-loaded nanofiber mats developed in the present work could be potentially useful as local drug delivery systems for treatment of several medical conditions, such as periodontal disease and skin, bone, and joint infections.

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ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acs.molpharmaceut.6b00039. DSC and TGA details (PDF)

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Phone: (312) 996-3472. Fax: (312) 413-0447. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS The authors gratefully acknowledge the Ministry of Education, Science, and Sport of the Republic of Slovenia and the Slovenian Research Agency for financial support through the Research Program P1-0189 and Project J1- 6746. The authors would also like to thank Slovene Human Resources Development and Scholarship Fund for providing the grant.

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REFERENCES

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