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Long-term Sustained Ciprofloxacin Release from PMMA and Hydrophilic Polymer Blended Nanofibers Spela Zupancic, Sumit Sinha Ray, Suman Sinha-Ray, Julijana Kristl, and Alexander L. Yarin Mol. Pharmaceutics, Just Accepted Manuscript • DOI: 10.1021/acs.molpharmaceut.5b00804 • Publication Date (Web): 04 Dec 2015 Downloaded from http://pubs.acs.org on December 11, 2015

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Molecular Pharmaceutics

Long-term Sustained Ciprofloxacin Release from PMMA and Hydrophilic Polymer Blended Nanofibers Špela Zupančič1,2, Sumit Sinha-Ray1, Suman Sinha-Ray1,3,4, Julijana Kristl2*, Alexander L. Yarin1, 5*

1

Department of Mechanical and Industrial Engineering, University of Illinois at Chicago, Chicago, Illinois 60607-7022, USA 2 3

4

Faculty of Pharmacy, University of Ljubljana, Aškerčeva cesta 7, 1000 Ljubljana, Slovenia

Corporate Innovation Center, United States Gypsum, 700 US 45N, Libertyville, IL-60048, USA

Department of Materials Science and Engineering, Indian Institute of Technology, Indore, Madhya Pradesh 452017, India 5

College of Engineering, Korea University, Seoul, S. Korea

* To whom correspondence should be addressed. Department of Mechanical and Industrial Engineering, University of Illinois at Chicago, Chicago, Illinois 60607-7022, USA E-mail: [email protected]. Phone: (312) 996-3472. Fax: (312) 413-0447. Faculty of Pharmacy, University of Ljubljana, Aškerčeva cesta 7, 1000 Ljubljana, Slovenia E-mail: [email protected]. Phone: +386-1476-9521. Fax: +386-1425-8031.

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Table of contents


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Abstract Nanofibers represent an attractive novel drug delivery system for prolonged release. However, sustained release of hydrophilic drugs, like ciprofloxacin hydrochloride (CIP), from polymeric nanofibers is not an easy task. The present study investigates the effect of different hydrophobic polymers (PCL and PMMA) alone in monolithic nanofibers, or with hydrophilic polymers (PVA, PEO, and chitosan) in blended nanofibers aiming to achieve sustained CIP release. CIP release from PCL nanofibers was 46% and from PMMA just 1.5% over 40 day period. Thus, PMMA holds great promise for modification of CIP release from blended nanofibers. PMMA blends with 10% PEO, PVA or chitosan were used to electrospin nanofibers from solution in the mixture of acetic and formic acid. These nanofibers exhibited different drug-release profiles: PEO containing nanofiber mats demonstrated high burst effect, chitosan containing mats revealed very slow gradual release, and PVA containing mats yielded smaller burst effect with favorable sustained release. We have also shown that gradual sustain release of antibiotic like CIP can be additionally tuned over 18 days with various blend ratios of PMMA with PVA or chitosan reaching almost 100%. Mathematical model in agreement with the experimental observation revealed that the sustained CIP release from the blended nanofiber corresponded to the two-stage desorption process.

Key words: nanofibers, sustained release, hydrophilic drug, drug release mechanism, ciprofloxacin hydrochloride, periodontal disease, desorption



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1. Introduction For decades controlled-drug release systems have been an important topic in the field of drug delivery due to their potential to improve drug therapy. There are several motivations for developing controlled-release systems. Indeed, they can enable superior control of drug release over time, help drug in crossing physiological barriers, decrease premature drug release, target the desired site of action while minimizing drug exposure elsewhere in the body, and increase patient compliance1. The sustained drug release is crucial for successful treatment of several diseases, as for example, periodontal disease, which is associated with chronic inflammation of periodontal tissue caused by pathogen bacteria. Treatment of periodontal disease with systemic antibiotic for 14 days leads to greater clinical and microbiological improvements compared to shorter treatments. Local administration of antibiotic to the inflamed body part is an important step toward more successful treatment compared to systemic administration, which can be associated with adverse side effects, uncertain patient compliance and lower concentration of the drug at local sites resulting in development of bacterial resistance2, 3. Therefore, a well-designed drug delivery system for local antibiotic administration is desired to maintain drug during longer period of time above the antibiotic minimum inhibitory concentration to be able to inhibit growth of the pathogen bacteria. In addition, the advantages of such local delivery systems are the ease of use, consistent retention at the placement site, fewer side effects, lower dosage requirements, bypassing the first-pass metabolism, reduction in gastrointestinal side effects, and decrease in dosing frequency3. Starting from the classical local drug delivery systems, such as powders, gels, inserts, and films4, new drug delivery systems, especially on the nano-scale level, have emerged and potentially can significantly contribute to better treatment outcome3. Nanofibers, fibers with diameters in the range from submicron to micron, present one of the newest and very promising nanomaterials for a number of applications. Several studies confirmed their importance in biomedicine, as drug delivery systems, tissue engineering, or both5. Their special characteristics, including very small diameters, a high surface-tovolume ratio, and a very high porosity of nanofiber mats and individual nanofibers, contribute to increased interaction with the biosurface and consequently to retention time at local side3, 6. In addition, nanofibers with their peculiar structure resemble the extra cellular matrix and thus promote cell attachment and proliferation resulting in faster wound healing or tissue regeneration7. The most frequently used method for nanofiber preparation with loaded drug is electrospinning. Preparation of nanofibers begins from randomly dispersed polymer macromolecules and drug molecules in solution. During electrospinning, when nanofibers are formed under the action of the electric Maxwell stresses, solvents evaporate and substances precipitate and solidify and are deposited as nanofiber mats5,

8, 9

. In the dry

nanofibers polymer macromolecules and drug molecules can interact physico-chemically, for example by 4 ACS Paragon Plus Environment

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the van der Waals interactions, hydrogen bonding, electrostatic interactions and physical entanglements whereas covalent bonds are seldom. The drug release efficiency upon contact with biological fluid is affected by both the chemical structure of the carrier and distribution of drug molecules in nanofibers5, 6. In the literature different drug release mechanisms were discussed, such as diffusion, dissolution, swelling, erosion or degradation of the matrix, and desorption; some of them can also overlap1,

10, 11

.

Different models were used to rationalize the drug-release results, such as the simplified Higuchi model with burst effect12, the Korsmeyer–Peppas model12, the desorption theory of Yarin et al. (eq. 2)10, and the two-stage desorption theory (eq. 4)13. Since the adsorption is the basic loading mechanism of drugs on polymer chains, desorption is the consequence of drug contact with water. Thus, the fundamental studies into the details of adsorption and desorption in nanofibers are critically important for their rational design as drug delivery systems capable of controlled delivery and release, while minimizing potential side effects. Theoretically and experimentally, it was shown that drug desorption is the rate-limiting mechanism of the drug release from nondegradable or slowly degradable hydrophobic polymeric nanofibers10, 13. In the applications the drug release rate can be tuned by selection of polymers (hydrophilic or hydrophobic, with different molecular weights and concentrations in solution), the type of nanofibers (monolithic, blended, or core-shell nanofibers), the drug loading and distribution, solubility and compatibility of drugs in the drug–polymer–solvent system, and nanofiber morphology5. Because the drug release rate from hydrophilic nanofibers, a-polymer-matrix type system, is fast, usually all drug is released in maximum one day14,

15

. The hydrophobic polymers, such as poly(glycolic acid) (PGA),

poly(lactic acid) (PLA), poly(lactic-co-glycolic acid) (PLGA), and polycaprolactone (PCL) have potential for longer release16. By comparison, poly(methylmethacrylate) (PMMA), a biocompatible and inert polymer with high spinnability, is also a hydrophobic polymer, which was up to date only infrequently considered as a potential candidate for drug retention and release from nanofibers10. However, several studies reported sustained drug release over one week or more from hydrophobic polymers but usually most of the drug was released in the first few hours with a low-level gradual release during the following days10, 17-19. Accordingly, that kind of release is insufficient for obtaining required drug concentration at the local side. Additionally, when a drug is released from non-biodegradable or slow-degradable nanofibers only the drug at the nanofiber surface (including nanopores) can be released, since the drug embedded in the polymer structure cannot be desorbed10. Long-term release of hydrophilic drugs is also challenging due to their high aqueous solubility, with opposite solubility properties in organic solvents compared to hydrophobic polymers, poor partitioning and incompatibility with hydrophobic polymers, which results in inhomogeneous incorporation, mostly in the form crystals on the surface, with a consequent burst release16. 5 ACS Paragon Plus Environment


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Therefore, the aim of the present study is to design a proper nanofiber composition required for sustained release of hydrophilic drugs, such as CIP, for at least 14 days. This study was designed to systematically investigate the possibility for production of hydrophobic nanofibers for incorporation of hydrophilic drugs, to study drug release and determine the release mechanism, as well as to optimize nanofibers. Firstly, monolithic nanofibers from two biocompatible hydrophobic polymers, PCL and PMMA with incorporated CIP, were produced and the release profiles of CIP determined. Secondly, an appropriate hydrophilic polymer, either chitosan, or poly(ethylene oxide) (PEO), or poly(vynil alcohol) (PVA) were blended with PMMA to improve the long-term sustained drug release. The release was further tuned with varying the amount of hydrophilic polymer. We show that it is possible to produce a drug delivery system as a depot for hydrophilic drug using blends of hydrophobic and hydrophilic polymers, since the latter play a pore-forming role and help to maintain a pre-determined level of drug concentration at a desired site of action.

2. Experimental Section 2.1. Materials Acetic acid (99%), formic acid (95%), PMMA (Mw = 996 kDa), PCL (Mw = 80 kDa), PEO (Mw = 8,000 kDa), chitosan (Mw = 50-190 kDa), and potassium phosphate monobasic were obtained from Sigma Aldrich, USA. PVA (Mw = 78 kDa, 88% DH) was obtained from Polysciences Inc., USA. NaOH and acetonitrile were purchased from Fisher Scientific, USA. CIP was obtained from Alfa Aesar, USA. All the chemicals were used as received without any further purification or modification. 2.2. Solution preparation A solvent mixture of acetic and formic acid with ratio 3:1 (w/w) was used for preparations of all polymeric solutions unless otherwise stated. Two types of monolithic nanofibers were prepared from 8 wt% PMMA and 15 wt% PCL solutions. Solutions for preparation of nanofibers from blends of PMMA and hydrophilic polymer had total polymer concentrations in the 8.2 – 8.9 wt% range to avoid too elastic solutions. In brief, compositions of 90 wt% of PMMA and 10 wt% of hydrophilic polymer (PMMA:polymer (90:10, with the polymer being PEO, chitosan, or PVA) were prepared by dissolving 88.9 mg of the polymer in 10.0 g of 8 wt% PMMA solution. Solution containing 95 wt% PMMA and 5 wt% PVA (PMMA:PVA (95:5)) was prepared by dissolving 42.1 mg of PVA in 10.0 g of 8 wt% PMMA solution and solution with 15 wt% PVA and 85 wt% PMMA (PMMA:PVA (85:15)) was prepared by dissolving 123.5 mg of PVA in 10.0 g of 7 wt% PMMA solution. Finally, nanofiber blend with 70 wt% PMMA and 30 wt% chitosan (PMMA:chitosan (70:30)) was prepared from 10 g of 6 wt% PMMA 6 ACS Paragon Plus Environment

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solution containing 257.1 mg of chitosan. Because chitosan has limited solubility in acetic acid, the solvent ratio of acetic and formic acid was changed to 3:2 (w/w) and to 1:1 (w/w) in case of PMMA:chitosan (90:10) and PMMA:chitosan (70:30), respectively. In all cases, 1 wt% of CIP powder (compared to the mass of all solid compounds added to solution) was added to solutions at room temperature 1 h prior electrospinning. The polymer solutions were used for electrospinning when all the components were fully dissolved. In some cases, polymer phase separations were observed. 2.3. Electrospinning The electrospinning conditions were optimized for each solution separately to obtain beadless nanofibers. The 8 wt% PMMA solution in formic and acetic acid was found to be sufficiently viscoelastic to form beadless nanofibers, with the viscoelastic relaxation time measured in the elongational experiments as described in chapter 2 of ref. 20 Consequently, the total polymer concentrations were kept in the 8.2 – 8.9 wt% range to avoid too elastic solutions. The nanofiber morphology was then optimized by changing electrospinning parameters and not by changing the total polymer concentrations. A 21G needle was used for delivering the solutions at flow rates in the 0.3 mL/h to 1.2 mL/h range and the voltage was in the 9-19.5 kV range. The inter-electrode distance was in the 10-19.5 cm range. All the electrospinning conditions along with the polymers and blend ratios are listed in Table 1. All nanofiber mats were collected on a rotating disk of 1.25 cm rim width. Table 1. The polymers with their ratios and electrospinning conditions used to produce nanofibers with 1 wt% CIP. Nanofibers PCL 15 wt% PMMA 8 wt% PMMA:PEO (90:10) PMMA:chitosan (90:10) PMMA:chitosan (70:30) PMMA:PVA (95:5) PMMA:PVA (90:10) PMMA:PVA (85:15)

Voltage (kV) 17 9 12 12 19.5 12 12 12

Distance (cm) 10 12 19.5 13 17 15 15 15

Flow rate (mL/h) 1.2 1.0 1.2 1.2 0.3 1.2 1.2 1.2

2.4. Polymer solutions and nanofiber characterization To examine the existence of different polymer phases in the polymer blend solutions a single drop was located between two glass slides to observe it under optical microscope (Olympus BX-51). Scanning electron microscopy (SEM) of nanofiber mats was conducted using high resolution field emission electron microscope JEOL JSM-6320F (RRC UIC). The average nanofiber diameter was 7 ACS Paragon Plus Environment


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calculated from measured 50 nanofiber diameters observed in SEM images using image analysis software (ImageJ, National Institutes of Health, USA). 2.5. Drug release experiments All nanofiber mats containing CIP were cut into rectangular pieces (1.5 cm × 6 cm) weighing 12–20 mg and put into a glass vial containing 10 mL of 50 mM phosphate buffer with pH 7.4 and covered with aluminum foil to prevent drug degradation caused by light. The vials with nanofibers were shaken with 120 rpm on an orbital rotator (Kangjian KJ-201BD, Jiangsu, China) at room temperature. At predetermined time intervals, 0.5 mL acceptor medium was withdrawn and replenished by a fresh buffer. The CIP fluorescence intensity in the extracted sample was measured using fluorescence spectrophotometer Tecan Infinite M200 Pro (Tecan Group Ltd., Männedorf, Switzerland) at the excitation wavelength of 280 nm and the emission wavelength of 450 nm. With the samples of nanofiber mats containing PVA, the sampling was conducted slightly differently because the CIP fluorescence intensity decreased in the presence of dissolved PVA, which was eluted from PMMA/PVA nanofibers. The extracted 0.5 mL samples were diluted with 0.5 mL of acetonitrile to precipitate PVA, separate it from the solution and measured as described before. Fluorescence increased due to formation of acetic and formic acid salt with CIP. To account for that, standard stock solution was prepared by dissolving drug in a mixture of acetic and formic acid at room temperature. To simulate the time frame of preparation and electrospinning of polymer solutions, the CIP in acids was neutralized in 5 h by phosphate buffer with a higher pH to reach at the end pH 7.4. The prepared standard solution was further diluted with phosphate buffer with pH 7.4. The calibration curve for measuring samples with PVA was prepared by further dilution with acetonitrile (50% (v/v)). The results were presented in terms of the cumulative release as a function of the release time

Re lease(%) =

Mt × 100% Md0

(1)

where Mt is the amount of CIP released by time t, and Md0 is the total amount of CIP in the nanofiber mat.

2.6. Drug release mechanism from nanofibers To study the mechanism of CIP release, the experimental release profiles from nanofibers were measured and matched with the predictions of the desorption-limited theory10, 13. A detailed exposition of the physical and mathematical aspects of the desorption-limited theory is available in ref.10. The desorption-limited theory describes the process, where the embedded drug is released from nanopores of 8 ACS Paragon Plus Environment

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the individual nanofibers or from the outer surface of these fibers in contact with water. For a single-stage desorption process the result is described by eq. 2

  π2 t   Mt = α 1 − exp  −  M d0  8 τr   

(2)

where Mt is the amount of CIP released by time t, Md0 is the total initial amount of CIP in the nanofibers, α is the nanoporosity factor, and τr is the characteristic time of the release process10. The total initial amount of CIP admixture, Md0, in the nanofibers is the sum of the CIP admixture Msd0 exposed at the surface (including the surfaces of nanopores open to the outer surface) and the embedded CIP Mbd0 (Md0=Msd0+Mbd0). The nanoporosity factor can be calculated as α=Msd0/(Msd0+Mbd0). The characteristic time τr obtained from matching the experimentally measured release profile can be used to calculate the effective diffusion coefficient, Deff, by equation Deff= L2/τr, where L presents the pore length, which is of the order of 1000 nm. The desorption factor, k(T), is then found as

k(T) =

ρ p 2aD eff

(3)

Db

with the molecular size, 2a, being of the order of 10-10 m, the pore cross-sectional radius, b, of the order of 10 nm, and the diffusion coefficient of CIP in the buffer solution, D, of the order of 10-5 cm2/s, ρp being the PMMA or PCL density, and T being temperature10. It should be emphasized that parameter values used are the plausible order-of-magnitude values. From the Clapeyron- like (or BET-like) desorption law, the desorption enthalpy (the activation energy) E is equal to

 k(T)  E = −RT ln    k0 

(4)

with the value of the pre-exponential k0 taken as 10-3 g/cm3 and temperature, T=300 K10. The characteristic time τr characterizes the duration of the release process. It depends on the polymer density, as well as the kinetic parameters of desorption, i.e. on the pre-exponential k0 and the activation energy E of the desorption process. Accordingly, τr manifests only the chemical nature of the polymer-drug interactions responsible for the sorption-desorption processes in the presence of water and is unaffected by polymer concentration and molecular weight10, 21. The two-stage desorption process is described by the following equation13 9 ACS Paragon Plus Environment


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   π2 t    π2 t   Mt = α1 1 − exp  − + α − 1 exp  −  2  τ Md 0 8    8 τr 2   r1   

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(5)

where α1 and τr1 correspond to drug release from the pre-existing pores, and α2 and τr2 correspond to the release of a leachable component. The experimental data discussed in the following section were matched with the desorptionlimited models using the software OriginPro 2015 (OriginLab corporation, USA) employing the Orthogonal Distance Regression to iteratively adjust the parameters. Each function was evaluated using correlation coefficient (R2).

3. Results and Discussion 3.1. CIP-loaded monolithic PMMA and PCL nanofibers The incorporation of drug into nanofibers to prolong release with minimal burst effect, as well as release over one week or more is often highly desirable16. Since hydrophilic polymers release the drug too rapidly (on the scale of one hour to one day) it becomes imperative to use hydrophobic polymers to slow down and retard the release rate. In the present study, we prepare two different types of monolithic hydrophobic nanofibers from PCL or PMMA with incorporated CIP and measured the resulting release. The first step for successful production of nanofibers with incorporated drug in hydrophobic polymers is the selection of an appropriate solvent, which dissolves both the drug and polymer. To obtain 1 wt% drug loading in nanofibers, the drug solubility in a selected solvent must exceed 1 mg/mL. CIP is soluble in water and acids22, 23, whereas PCL and PMMA - in organic solvents, thus the only possible solvents meeting the requirements were formic and acetic acids. Even though acetic acid is more biocompatible compared to formic acid and consequently preferential, acetic acid alone does not allow formation of beadless PCL nanofibers due to its low dielectric constant (ε=6.2)24. Therefore, a minimum amount of formic acid (ε=58)24 was used to increase the dielectric constant of the solvent mixture, which resulted in better spinnability of polymer solutions and lower bead formation on nanofibers due to the capillary instability. The PMMA nanofibers were smooth with the shape of a bow tie in cross-section and the average diameter in the longer dimension measured as 855±140 nm (Figure 1). The bow-tie structure can be explained by the buckling phenomenon resulting from the excessive solvent evaporation. In this case the nanofiber skin gets solidified first. The remaining solvent in the nanofiber core evaporates and diffuses out through the solidified skin resulting in a decreased pressure in the core. As a result, the nanofibers collapse under the action of the atmospheric pressure from the outside. Depending on the pressure differential applied to the nanofiber core, different cross-sectional configurations of the collapsed 10 ACS Paragon Plus Environment

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nanofibers could arise: the elliptical ones (ribbons) at relatively low pressure differentials, or a bow-tie structure at higher pressure differentials25-28. By comparison, PCL nanofibers revealed circular crosssections, rough and uneven surface with the average diameter of 245±84 nm (Figure 1). The drug was successfully incorporated in both types of nanofibers and no drug crystals were observed on the surface.

Figure 1. SEM images of dry PMMA (the left-hand side column) and PCL (the right-hand side column) monolithic nanofibers with 1 wt% CIP incorporated. Panel (a) shows nanofiber mats, panel (b) the zoomed-in view of nanofibers, and panel (c) shows the nanofiber cross-sections.

3.2. The effect of PMMA and PCL nanofibers on CIP release The CIP release profiles from PMMA and PCL nanofiber mats are shown in Figure 2. Although CIP was successfully incorporated into PMMA nanofibers, as it can be concluded from the lack of drug crystals in the SEM images (Figure 1), only 1.5% of CIP was released in 40 days from PMMA, albeit without an initial burst effect. On the other hand, PCL nanofibers released 30% of CIP in the first 10 h, whereas the next 16% were released during the following 39 days. There are several factors affecting drug release from these polymers, such as the chemical nature of the polymer, the polymer concentration (in the solution used for electrospinning) and the molecular weight affecting nanofiber porosity, as well as the drug interaction with the polymer. The higher drug release from PCL nanofibers compared to the minimal release from PMMA nanofibers are in line with the previously reported results 10. Both polymers 11 ACS Paragon Plus Environment


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are sufficiently and approximately identically hydrophobic (as is seen from their contact angles: PCL = 82.5° or 90° and PMMA = 77°)29, 30, 31. Several studies observed burst drug release from PCL nanofibers, which was a consequence of forced drug migration to the nanofiber surface during electrospinning process18,

19, 32

. The glass transition temperature Tg of the polymers affects drug incorporation in the

nanofibers and drug release from them. Polymers with lower Tg, such as PCL (Tg = -60 °C), which were in the rubbery state during release study had exhibited a stronger burst effect compared to polymers with higher Tg, e.g. PMMA with Tg =105 °C or PLLA with Tg = 60–65°C17, 18, 33, 34. PMMA can successfully incorporate the drug in the matrix and its glassy state minimizes water penetration into nanopores in the individual nanofibers and inside nanofiber mats, as our results in Figure 2 show.

Figure 2. Release profiles of CIP from PCL and PMMA monolithic nanofibers. (a) The release profiles versus linear axes. (b) The release profiles in semi-logarithmic axes. Zoomed-in release profiles of CIP from PMMA nanofibers are shown in the insets. Error bars correspond to the standard deviation from the mean value. The dashed lines are the best fits of the measured release data using the single-stage desorption-limited model (eq. 2) and the solid line corresponds to the two-stage desorption-limited model, (eq. 5).

PMMA and PCL nanofibers can be considered as non-biodegradable matrix systems because PMMA practically does not degrade at all, whereas PCL degrades over a span of two years, which is too slow to have an impact on a 40-day release study10, 35. Drug release form PCL and PMMA nanofibers saturated at the level much lower than the 100% release (Figure 2), which evidently excludes drug diffusion in nanofiber matrix as a driving mechanism of the release process (diffusion can never stop below 100%10, 35). This release pattern from hydrophobic, non-biodegradable matrix system is compatible with the desorption-limited mechanism10, which consists of three stages: (1) filling the nanopores with 12 ACS Paragon Plus Environment

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water, (2) desorption of drug from the nanofiber surface (including nanopore surface) in contact with water, and (3) diffusion of drug in water from the pore. Stages 1 and 3 are very rapid10, whereas the desorption stage is the rate-controlling release mechanism. It should be emphasized that the desorptionlimited theory also explains why not the entire amount of drug can be released (because the drug embedded in the polymer bulk cannot be accessed by water, and thus cannot be desorbed)10. In addition, the initial 48-h release from polymer films was also attributed to desorption35. However, in case of drug release from nanofibers this mechanism is more expressed compared to all other drug delivery systems due to the fact that nanofibers possess extremely large surface and numerous pores. The CIP release profiles from monolithic PMMA and PCL nanofiber matched with the single-stage desorption-limited model (eq. 2) are shown in Figure 2. The model parameters corresponding to release from PMMA nanofibers are listed in Table 2. The high correlation coefficient (0.9998) shows that the CIP release from PMMA nanofibers is described accurately by the single-stage desorption-limited model. The corresponding nanoporosity coefficient α = 1.3% which arises from the level at which CIP release reached the plateau value. It is higher compared to the value of α for PMMA nanofibers, which had incorporated rhodamine B10. The result is expected, because the present nanofibers were prepared from a lower concentration polymer solution compared to Ref. 10, 8 wt% here versus 15 wt% PMMA solution. The activation energy E is smaller in the present case of CIP interaction with PMMA compared to rhodamine B since both compounds interact differently with PMMA10.

Table 2. Parameters of the single-stage desorption-limited theory for CIP release from PMMA nanofibers.

τr α Deff ρp k(T) E 2 3 3 (%) (h) (cm /s) (g/cm ) (g/cm ) (kJ/mol) -14 -11 0.9997 1.3±0.1 101.7±22.9 2.7×10 1.18 3.22×10 43.0 R2

PMMA

The release profile of CIP from PCL nanofibers is composed of an initial burst effect followed by gradual release over 20 days (Figure 2) which differ from the other studies dealing with rhodamine B release from PCL nanofibers10. During the solution preparation and electrospinning some of PCL macromolecules hydrolyzed in the presence of organic acids36 resulting in a mixture of PCL macromolecules of different chain length. The hydrolysis of ester bonds in PCL chains increased the amount of hydrophilic groups (carboxylic and hydroxyl group) in the nanofibers. The -COOH group was ionized in the buffer solution with pH 7.4 and together with the -OH group decreased the hydrophobic nature of PCL macromolecules. More hydrophilic parts swelled and allowed water to penetrate into 13 ACS Paragon Plus Environment


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nanofibers through the newly formed nanopores. Therefore, the polymer (PCL) properties changed in situ. The two-stage release was also confirmed by matching with the single-stage and two-stage desorption models. The two-stage desorption model fits the data better than the single-stage desorption model (R2Twostage desoprtion

= 0.9998 versus R2single-stage desoprtion = 0.9958) (Figure 2). The corresponding parameters of the

two-stage model are listed in Table 3. The initial desorption was fast, within 2h, which was significantly faster compared to the release of rhodamine B or albumin10, 21, which implies weaker interactions between CIP and PCL than with the former compounds. By contrast, newly formed -COOH and -OH groups could contribute to stronger ionic and hydrogen bonds between polymer and CIP resulting in stronger CIP bonding to shorter PCL chains. However, PCL nanofibers with shorter PCL chains revealed better and much longer gradual release of CIP compared to the other studies

10, 18, 21

. The CIP incorporated in PCL

nanofiber mat could have been an option for controlled drug release for its weaker initial burst effect and subsequent sustained release. However, it should be emphasized that only 46% of drug was released from PCL nanofibers and the rest stayed embedded even after the 40-day long release study. This renders the PCL nanofibers inefficient for sustained drug release. By comparison, PMMA nanofibers without any burst release present an inappropriate drug delivery tool for 20 days, albeit a potentially very good hydrophobic material to tune the drug release in blends with hydrophilic polymers. Table 3. Parameters of the two-stage desorption theory determined from matching the experimental data in Figure 2 for CIP release from PCL nanofibers. R2 PCL

0.9998

α1

α2

τr1

τr2

(%)

(%)

(h)

(h)

28.5±0.4

18.3±0.5

1.9±0.2

223.4±16.2

Deff,1 2

(cm /s) -12

1.5×10

E1

E2

2

(kJ/mol)

(kJ/mol)

-14

33.2

45.1

Deff,2 (cm /s) 1.2×10

3.3. CIP-loaded nanofibers from PMMA blended with either chitosan, or PEO, or PVA As was shown in the previous section, PMMA presents a perfect hydrophobic material to be blended with hydrophilic polymers to tune the release rate and reach a desirable release profile over a preferred period of treatment. In literature, to the best of our knowledge, there are no publications on nanofiber blends composed of PMMA and hydrophilic polymers used for drug release. Therefore, three hydrophilic polymers (PEO, PVA and chitosan) were studied in blends with PMMA to increase CIP release from PMMA and thus this study provides a unique drug-polymer design for sustained CIP release. Blending of 10 wt% of hydrophilic polymer in PMMA nanofibers had a strong impact on the nanofiber morphology (Figure 3). The PMMA and chitosan solution was more conductive compared to PMMA-alone solution due to polycationic nature of chitosan. The increased solution conductivity 14 ACS Paragon Plus Environment

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resulted in thinner blended nanofibers compared to PMMA-alone nanofibers with the average diameter of 555±70 nm. This is because the higher electric charge carried by the jet increased stretching electric forces and the electrically-driven bending instability causing formation of thinner nanofibers (Figure 3). The solidification of the polymer surface in this blend was slower and enabled the jet to shrink uniformly and form nanofibers with circular cross-sections. The addition of PEO to PMMA resulted in nanofibers with mostly circular cross-sections, albeit with some nanofibers with cross-sections in the shape of an ellipse or in a few merged nanofibers (Figures 3 and 4). The incorporation of high-molecular weight PEO resulted in higher elastic forces in the solution stemming from an increased macromolecular entanglement, which leads to formation of thicker fibers (1475±490 nm) compared to PVA nanofibers with lower molecular weight (958±160 nm) and PMMA nanofibers (855±140 nm). The incorporation of PVA in PMMA nanofibers did not change the nanofiber structure compared to that of PMMA nanofibers, i.e. the bow-tie cross-sections were observed once more (Figures 3 and 4).

Figure 3. SEM images of blended nanofibers composed of 90 wt% of PMMA and 10 wt% of hydrophilic polymer. PMMA blended with chitosan (the left-hand side column), PMMA blended with PEO (the middle column), and PMMA blended with PVA (the right-hand side column). In all cases the nanofibers incorporated 1 wt% CIP. Panels (a) show nanofiber mats (at the same magnification) and panels (b) show the zoomed-in view of nanofibers.

Some of the prepared PMMA blends with hydrophilic polymers formed two-phase solutions due to microphase separation when two thermodynamically incompatible linear polymers were mixed. 15 ACS Paragon Plus Environment


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Hydrophobic PMMA was separated from hydrophilic PEO and PVA (Figure 4). PMMA and PVA resulted in a uniform dispersion, whereas a non-uniform dispersion was observed in the case of PMMA/PEO blend. The stability of the phase-separated polymers was evaluated and the size of dispersed phase did not change during 24 h. The phase-separated polymers affected the inner nanofiber structure, as is seen from their cross-sections (Figure 4). Similar results were reported previously37, where PMMA and polyacrylonitrile formed core-shell fibers. The cross-sections of both PMMA/PVA and PMMA/PEO fibers were solid without any voids. After the drug release experiments, the nanofibers were dried to observe their cross-sections. In the case of PMMA:PEO (90:10), SEM images revealed some intact crosssections and cross-sections with multiple holes. This is reminiscent of the morphologies reported in37. By contrast, PMMA and PEO polymers in the same ratio were also investigated in the other studies, where miscibility of both polymers in blends was reported38,

39

. However, this conclusion differs from the

present case, since different molecular weight polymers are used here: PEO with Mw = 8 MDa and PMMA with Mw = 996 kDa versus PEO with Mw = 463 kDa and PMMA with Mw = 315 kDa used before39. Also, solvents (the present mixture of formic and acetic acids versus chloroform), and methods of preparation of polymer blends (electrospinning versus solvent evaporation) were different here in comparison with ref. 39. After immersion in water, PVA dissolved and left behind small holes in PMMA:PVA (90:10) nanofibers. These holes were too small to be resolved properly by SEM. However, the emergence of the porous cross-sections suggests intertwined polymer phases in the overall structure, even though optical microscope images clearly show their separation in solution. Note also that PMMA/chitosan solution did not show phase separation. As reported in ref. 40, PMMA and chitosan form chemical bonds between effective side groups and due to an increased intermolecular interaction an enhancement in tensile properties was also seen resulting from such blends up to a certain concentration. The latter indicates a limited phase separation between these two polymers.


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Figure 4. The optical microscope images of: (a) PMMA:PEO (90:10), and (b) PMMA:PVA (90:10). In both cases, the two-phase polymer blends (the left-hand side column), and nanofiber cross-sections before (the middle column) and after drug release in phosphate buffer for 18 days (the right-hand side column). The scale bars of nanofiber cross-section inserts present 1 μm.

3.4. The effect of chitosan, PEO and PVA polymers in blends with PMMA on CIP release The effect of incorporation of different hydrophilic polymers in blends with PMMA on CIP release is illustrated in Figure 5. The type of hydrophilic polymer admixture significantly influences the release profile from nanofibers. The fastest CIP release was obtained from PMMA:PEO (90:10), where the burst effect was observed reaching 69% in the first 6 h and an additional 8% release in following 17 days. By contrast, nanofibers with chitosan did not show any bust effect and a slow gradual release was observed during 20 days reaching 11%. PMMA:PVA (90:10) nanofibers revealed a two-stage release pattern, where release reached 56% in the first 13 h and additionally 38% in the following 19 days.



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Figure 5. Release profiles of CIP from blended nanofibers with 10 wt% of either PVA, or PEO, or chitosan added to PMMA. The solid lines are the best fits of the measured release data to the two-stage desorption model (eq. 5).

The basic desorption-limited drug release mechanism here is similar to the case of PMMA nanofibers described before. However, this mechanism is accompanied by additional steps due to the incorporation of hydrophilic polymers, which behave completely different compared to the hydrophobic one. At the beginning of the release process, a hydrophilic polymer increases wettability of the nanofiber surface, which enhances nanofibers solvation and drug desorption from the outer surface of PMMA/hydrophilic blended nanofibers, as well as from the first-generation nanopores. This stage may be compared with CIP release from PCL (as described above). Simultaneously, hydrophilic polymers start leaching from the surface toward the inner parts of the nanofibers acting as porogens. The leaching process increases the distances between individual PMMA chains, and forms pores and channels open to the outer surface, which further enhance water imbibition. As soon as water encounters previously encapsulated area, CIP desorption begins from the newly formed surface. Then the drug is dissolved in the buffer solution that imbibes the surrounding medium. In some time, polymers chains disentangle, relax and in case of water-soluble polymers they dissolve, forming bigger pores, which furthermore promote water imbibition to the inner parts of the blended nanofibers. Overall, hydrophilic polymers increase nanofiber wettability and act as porogens forming new pores in the previously inaccessible nanofiber bulk. In brief, the mechanism described above includes several processes: solvent diffusion and imbibition, hydrophilic polymer swelling and dissolution, drug desorption from the pre-existing and newly formed surfaces, polymer chain disentanglement, and diffusion of the desorbed drug in the buffer solution. At any stage, the slowest process becomes the rate- limiting step and ultimately controls the 18 ACS Paragon Plus Environment

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release rate. Sometimes one or more processes may proceed extremely fast and hence have no chance to control the release rate and the overall release profile. In such cases, these processes could be omitted from the drug release models, as only the rate-limiting steps determine the release rates and are reflected in the existing models35. As described, nanofiber release mechanism includes several processes and most of the models neglect at least some of them. Therefore, they would be inappropriate for description of drug release from nanofibers if they are incapable of accounting for the presence of porogens. Only the latter allow drug release of 100% to be reached. It should be emphasized that the best fitting (R2) of the experimental data in all the cases of CIP release was achieved with the two-stage desorption theory, as shown in Figure 5. The corresponding parameters are listed in Table 4. The release process consists of two stages, where in the first one, the drug is released from the pre-existing nanopores, and in the second one, the drug is released from the nanopores or microcracks formed as a result of leaching or dissolution of the fiber body or the generation of microcracks13.

Table 4. Parameters of the two-stage desorption theory determined in Figs. 5 an 6 for by matching the experimental data for CIP release from different blended nanofibers.

E2

(cm2/s)

(kJ/mol)

(kJ/mol)

1.5×10-12

1.1×10-14

32.9

45.3

192.5±29.1

5.5×10-13

1.4×10-14

35.5

44.6

8.9±2.0

403.2±71.3

3.1×10-13

6.9×10-15

36.9

46.4

25.4±5.0

2.3±7.3

141.9±39.4

1.2×10-12

2.0×10-14

33.5

43.8

0.0±1.8

11.8±1.5

67.9±0.0

274.8±67.8

4.1×10-14

1.0×10-14

42.0

45.5

5.7±1.5

78.6±1.5

7.5±6.5

287.3±20.1

3.7×10-13

9.7×10-15

36.5

45.6

α2

τr1

τr2

Deff,1

(%)

(%)

(h)

(h)

(cm2/s)

0.9996

68.5±1.3

9.1±1.4

1.8±5.5

257.0±135.2

0.9999

31.7±4.0

62.0±2.8

5.1±6.8

1.0000

66.5±1.2

38.5±2.4

0.9994

72.3±5.3

0.9968 0.9999

R PMMA:PEO (90:10) PMMA:PVA (95:5) PMMA:PVA (90:10) PMMA:PVA (85:15) PMMA:chitosan (90:10) PMMA:chitosan (70:30)

E1

α1

2

Deff,2

Figure 5 shows that the type of hydrophilic polymer can dramatically alter the CIP release pattern. PEO demonstrated the burst effect, chitosan revealed a completely gradual release without any burst effect, and PVA yielded a two-stage release with both effects present. Still in all the cases, the limiting processes are the hydrophilic polymer leaching and CIP desorption. The two-stage desorption theory outlined above accounts for both processes. The burst release from the PMMA:PEO (90:10) nanofibers is related to the high PEO solubility in water and its long macromolecular chains, which 19 ACS Paragon Plus Environment


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facilitated water penetration into PMMA nanofibers and desorption of the embedded drug almost at the same time, which resembles more a single-stage rather than a two-stage process. The entire release process was very fast with the characteristic time less than 2 h (Figure 5 and Table 4). At this stage newly developed porosity was concurrent with the drug desorption to that extent that at the next stage small additional drug release was found which corresponds to α2= 9.1 and τr2=227 h for PMMA:PEO (90:10) nanofibers in Table 4. As is seen in Figure 4a, due to the original two-phase solution, some PEO chains stay together and did not form a perfect blend with PMMA. This results in some drug being embedded in PMMA alone, with no access of water and thus no drug release from there. Therefore, the drug release from PMMA/PEO nanofibers saturated at 75% (Figure 5). PVA also promotes a burst effect because PVA on the nanofiber surface increased wettability of the nanofibers. As a result, some PVA chains were dissolved, and CIP near the surface could be desorbed at an enhanced rate resulting in release of 66.5% in the first day (which corresponds to the value of α1 for PMMA:PVA (90:10) in Table 3). However, a gradual release over 19 days was also observed from these nanofibers, corresponding to the long characteristic time τr2 = 403.2 h and α2 ≈ (100– α1)%, i.e practically to 100% release of CIP. The sustained release was the consequence of a lower PVA solubility in water compared to that of PEO, shorter PVA chains, and formation of good blend with PMMA. The advantage of good blend was expressed in the high ultimate release percentage (94%). The incorporation of chitosan into nanofibers did not promote the burst effect, which is formally expressed by α1=0. Because PMMA/chitosan blend solution did not separate and form a two-phase solution, there is a high possibility that such nanofibers consist of uniformly entangled PMMA and chitosan macromolecular chains. Better incorporation of chitosan into nanofiber structure and lower wettability reduce the initial burst release to minimum. Chitosan has a pKa in the range of 5.5 – 6.5. It is soluble at lower pH, whereas at pH 7.4 it is not ionized and accordingly insoluble in a selected medium41. However, chitosan was swollen under certain conditions in the release experiment due to its hydrophilic hydroxyl and amino groups. Nevertheless, chitosan swelling was much slower than that of PVA or PEO, which resulted in a slow release. Because the incorporation of PVA and chitosan in PMMA blended nanofibers revealed gradual sustained release, both polymers were chosen for a further investigation. 3.5. The effect of the varying ratios of PVA and chitosan in blends with PMMA on CIP release The changing ratios of hydrophobic and hydrophilic polymers dramatically alter the rate of CIP release (Figure 6). On the contrary, the pattern of CIP release stayed the same for specific polymer, the two-stage release for PVA (Figure 6a) and gradual sustained release for chitosan (Figure 6b). The CIP release profiles from blended nanofiber with the parameters of the matched two-stage desorption model


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are presented in Figure 6 and Table 4. The high correlation coefficients in all cases (R2>0.997) reveal good matching of the-two-stage desorption model with the experimental data. In the case of PVA, all nanofibers ultimately released up to 80-95% of CIP in 18 days. A lower percentage of PVA in nanofibers resulted in a lower burst effect and a following steeper gradual release. The percentage of drug released at the burst effect stage is presented in Table 4 by the value of α1 and during the gradual release stage - by the value of α2. In the case of PMMA:PVA (85:15) nanofibers, 72.3% of CIP was released in the first 3 h, whereas in the following 17 days the subsequent 25.4% was released. By contrast, nanofibers with 5 wt% PVA released only 31.7% of drug initially and later during the gradual sustained release stage the additional 62.0% was released. The blended nanofibers with 30 wt% of chitosan released during the sustained stage 72% in 18 days with minimal burst effect in the first day. The lower is the desorption energy of CIP, the faster is its release. At the first stage of the process the activation energy of the process with PVA or chitosan was about 35 kJ/mol and then increased to about 44 kJ/mol (Table 4). The increase in the activation energy corresponds to a slower release limited by dissolution of porogen. It should be emphasized that the activation energy strongly depends on the chemical interaction, and surface polarizability. To conclude, changing the ratio of PMMA to PVA or chitosan tuned drug release to obtain small burst effect for rapidly achieving the antibiotic inhibitory concentration and later a sustained CIP release over 18 day.

Figure 6. Release profiles of CIP from blended PMMA nanofibers incorporating different (a) PVA and (b) chitosan percentages. The solid lines are the best fits of the measured release data to the two-stage desorption model (eq. 5).



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4. Conclusion CIP, hydrophilic antibiotic, was successfully incorporated in PCL and PMMA monolithic nanofibers, as well as in blended PMMA/hydrophilic polymer nanofibers formed by electrospinning of solutions in acetic and formic acid. These are the suitable solvents for dissolving all hydrophobic and hydrophilic polymers as well as the drug. CIP release from PCL nanofibers achieved 46% and from PMMA – only about 1.5% over the 40-day period. Accordingly, PMMA presents a really good hydrophobic material to tune drug release in blends with hydrophilic polymers. However, the different chemical structure of polymers caused polymer phase separation in solution. That fact influenced the resulting blended nanofiber structure, as was observed with PMMA and PEO blends. In addition, either chitosan, or PEO, or PVA in blends with PMMA dramatically alter the CIP release profiles: chitosan revealed a slow gradual release, PEO demonstrated a strong burst effect up to 75% in 1 day, while PVA yielded a two-stage release with both effects present. Furthermore, the changed ratios of PMMA to PVA or chitosan tuned drug release to obtain small burst effect for rapidly achieving the antibiotic inhibitory concentration and later a sustained release over 18 day. It is shown that the sustained CIP release from blended nanofibers agrees with the two-stage desorption theory with high degree of correlation.

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. The authors gratefully acknowledge Patrick Comiskey's help at the initial stage of this work.

Disclosure The authors report no conflicts of interest in this work.


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