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Synthesis and Characterization of Poly(2hydroxyethylmethacrylate) Contact Lenses Containing Chitosan Nanoparticles as an Ocular Delivery System for Dexamethasone Sodium Phosphate Gautam Behl, Javed Iqbal, Niall J. O’Reilly, Peter McLoughlin & Laurence Fitzhenry Pharmaceutical Research An Official Journal of the American Association of Pharmaceutical Scientists ISSN 0724-8741 Pharm Res DOI 10.1007/s11095-016-1903-7

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Author's personal copy Pharm Res DOI 10.1007/s11095-016-1903-7

RESEARCH PAPER

S y n t h e s i s a n d C h a r a c t e r i z a t i o n of Poly(2-hydroxyethylmethacrylate) Contact Lenses Containing Chitosan Nanoparticles as an Ocular Delivery System for Dexamethasone Sodium Phosphate Gautam Behl 1 & Javed Iqbal 1 & Niall J. O’Reilly 1 & Peter McLoughlin 1 & Laurence Fitzhenry 1 Received: 21 September 2015 / Accepted: 4 March 2016 # Springer Science+Business Media New York 2016

ABSTRACT Purpose Dexamethasone sodium phosphate (DXP) is an anti-inflammatory drug commonly used to treat acute and chronic ocular diseases. It is routinely delivered using eyedrops, where typically only 5% of the drug penetrates the corneal epithelium. The bioavailability of such ophthalmic drugs can be enhanced significantly using contact lenses incorporating drug-loaded nanoparticles (NPs). Methods The mechanism of release from chitosan NPs (CSNPs), synthesized by ionic gelation, was studied in vitro. The DXP loaded CS-NPs were subsequently entrapped in contact lenses and the optical and drug-release properties were assessed. Results DXP release from CS-NPs followed diffusion and swelling controlled mechanisms, with an additional proposed impact from the electrostatic interaction between the drug and the CS-NPs. The release rate was found to increase with an increase in drug loading from 20 to 50 wt%. However, an inverse effect was observed when initial loading increased to 100 wt%. NP-laden lenses were optically clear (95–98% transmittance relative to the neat contact lens) and demonstrated sustained DXP release, with approximately 55.73% released in 22 days. Conclusions The release profile indicated that drug levels were within the therapeutic requirement for antiinflammatory use. These results suggest that these materials might be a promising candidate for the delivery of DXP and other important ophthalmic therapeutics. Gautam Behl and Javed Iqbal contributed equally to this manuscript. * Laurence Fitzhenry [email protected]

1

Pharmaceutical and Molecular Biotechnology Research Centre, Department of Science, School of Science and Computing, Cork Road, Waterford Ireland

KEY WORDS chitosan . contact lens . dexamethasone sodium phosphate . eye drops . nanoparticles . transmittance

ABBREVIATIONS CS-NP DXP EE EGDMA HEMA LC NP PEG PGT pHEMA PLGA TPO TPP

Chitosan nanoparticle Dexamethasone sodium phosphate Encapsulation efficiency Ethylene glycol dimethacrylate 2-hydroxyethlymethacrylate Loading capacity Nanoparticle Polyethylene glycol Propoxylated glyceryl triacylate Poly(2-hydroxyethylmethacrylate) Poly(lactic-co-glycolic acid) Trimethylbenzoyl-diphenyl-phosphineoxide Tripolyphosphate

INTRODUCTION The potential to use contact lenses as a vehicle for targeted and controlled delivery of ophthalmic drugs has long been discussed, and, as has been documented elsewhere [1], the first reference to contact lenses as drug delivery platforms dates back to 1965 [2]. Reasons for the interest in such a delivery device stem from issues with traditional modes of delivery such as eye drops, where as little as 1% of the drug is bioavailable, or implantable plugs or inserts, where surgery is needed. The appeal of using contact lenses as drug depots that could control the release of drugs can readily be seen and is further bolstered by the potential to improve bioavailability. Such enhancement of bioavailability has been demonstrated by the oft cited work of Hilman, using pilocarpine-loaded lenses [3]. It has also been more recently predicted by Kim and Chauhan, where they used a drug transport model to

Author's personal copy Behl et al.

demonstrate the increased bioavailability compared to topically administered drugs [4], and again demonstrated by Tieppo et al., where ketotifen fumarate release was extended in vivo for 26 h using molecularly imprinted contact lenses [5]. Despite the obvious benefits of contact lens use, there can also be a number of drawbacks. For example, while the tolerance for lens wear is generally high, the extended wear of a hydrogel lens has been cited to lead to not only intolerance, but also to side effects such as corneal vascularisation and corneal ulceration [6]. Coupled with this, for patients with tear stability issues, contact lens wear may not be suitable and in a recent literature study by van Tilborg et al., it was stated that general practitioners ranked contact lens as second to tear deficiency for causing symptoms of dry eye [7]. With the advent of silicon-hydrogel lenses, however, many of the side effects of extended wear of contact lenses can be more readily avoided, and, as stated by Wolffsohn et al., therapeutic lenses may be useful in the treatment of some of the conditions listed [8]. In relation to their use as drug delivery vehicles, simple drug-soaking methods, which have been applied in an effort to incorporate drugs into contact lenses, have traditionally proven unsuitable, in particular for highly hydrophilic drugs such as dexamethasone sodium phosphate [9]. In an effort to develop contact lenses capable of such use, a range of technologies have been applied, including molecular imprinting, an approach pioneered by Byrne et al. [5,10]. This technology creates specific binding sites within the polymeric network that allow for increased drug loading and the release of compounds such as hyaluronic acid [10], diclofenac sodium [11] and, as stated, ketotifen fumarate [5] indicates the potential of this promising technology. Chauhan et al. have studied and developed a number of technologies that have shown considerable potential, including the use of a hydrophobic barrier, Vitamin E, which is incorporated into the lenses [12]. This approach uses the hydrophobic vitamin to create a more tortuous path for, in particular, hydrophilic drugs to diffuse through as they migrate from the lens to the eye, extending the delivery period for up to as much as a factor of 400 for timolol [12] and 16 for the hydrophobic drug, dexamethasone [13]. The same group have also used nanoparticles prepared from both PGT [14,15] and silica-stabilised emulsions [16] for their incorporation into lenses. Ciolino et al. have recently achieved the release of the hydrophobic lantanoprost from contact lenses for up to 1 month, by encapsulating a drug-polymer film in a contact lens [17]. This drug is widely used for the treatment of glaucoma and as such, its extended release represents a clear alternative for the management of this condition. It could be said that one of the many obstacles to the commercialisation of contact lenses for ocular drug delivery has been the difference observed in the in vivo release behaviour compared to in vitro experiments. In vivo experiments using contact lenses for ocular drug delivery are required, not only

to determine the release behaviour in physiological systems, but also to demonstrate the safety of such lenses. To this end, the work of Chauhan et al. with Beagle models has shown, with the exception of some Boccasional mild redness^, no evidence of ocular toxicity [18]. In vivo safety has also been demonstrated by Ciolino et al. where New Zealand white rabbits were used [17]. The authors observed no signs of discomfort in the animal models, though some of the eyes did develop corneal neovascularisation, which was possibly due to poor fitting lenses. In an earlier work, Xu et al. also used rabbits to demonstrate in vivo release behaviour and for one of their contact lenses, observed no irritation or protein deposition [19]. The authors did, however, observe some lens deposition on another of their materials, again demonstrating the need for in vivo studies. While these examples demonstrate the usefulness of animal models, there is still an understandably lower instance of human in vivo studies. A number of research groups, however, have worked on improving in vitro experiments [11,20] in an effort to ensure that only the most suitable materials progress to human trials. A recent publication by Bobba et al. did, however, use contact lens delivery for stem cells to human eyes [21] with generally positive results, citing only two (of sixteen) patients suffering minor complications from contact lens insertion and removal. It is necessary to note that the type of procedures documented for stem cell transplant would be inherently more prone to complications than the type of contact lens use proposed in the majority of works cited in this article. As stated, one of the drawbacks of using contact lenses for drug delivery has always been their suitability for use with highly hydrophilic drugs, due to the unsuitable diffusion characteristics and the high water content of contact lenses [22]. As such, the extended release of drugs like dexamethasone sodium phosphate, a corticosteroid with a water solubility of 500 mg/ml that is used for the treatment of inflammation and post-cataract surgery, has proven to be a considerable challenge. While the use of surfactants [23] and the aforementioned hydrophobic barrier have delivered significant success, this work aims to add a degree of control to the release based not just on diffusion but also interaction on the molecular level between components of the polymer and the drug molecule. As such, the negatively charged dexamethasone has been incorporated into the positively charged chitosan nanoparticles, which have then been incorporated into contact lenses. The model lenses are prepared from poly(2-hydroxyethylmethacrylate) (pHEMA) and have been characterised using a range of techniques that demonstrate their suitability for use. This work aims to demonstrate that the combination of biocompatible and biodegradable polymer particles with contact lenses can be used for the treatment of a range of ocular conditions and that the control of not just hydrophobic, but also hydrophilic drugs is achievable.

Author's personal copy CSNP-Laden Contact Lenses for Extended DXP Release

MATERIALS AND METHODS

Preparation of Nanoparticle-Laden pHEMA Contact Lens

Materials

The particle loaded pHEMA hydrogel lenses were prepared by free radical polymerization of the monomer by photoinitiation [25]. Briefly, the chitosan nanoparticles were mixed with HEMA monomer (20 mmoles) and subjected to ultrasonication in an ice bath to form a fine suspension followed by an addition of 15 μl of EGDMA and 0.02 mmoles of TPO. The final suspension was purged with nitrogen and injected into lens moulds and the reaction was allowed to proceed beneath a mercury lamp (Phillips Actinic BL PL-S 9W/10/2P 1CT) with a peak wavelength of 380 nm under an inert atmosphere for 2 h. The neat lens was prepared following a similar procedure without the addition of nanoparticles. Following the procedure, a lens of about 10 mm size was obtained with a centre thickness of 50 μm as measured by an Electronic Thickness Gauge (Model ET-3).

Chitosan low molecular weight (20–300 cP, degree of deacetylation 75–85%), sodium tripolyphosphate (technical grade, 85%), dexamethasone 21-phosphate disodium salt (≥98%), 2-hydroxyethylmethacrylate (HEMA) (≥99%, contains ≤50 ppm monomethyl ether hydroquinine as inhibitor), ethylene glycol dimethacrylate (EGDMA) (98%, containing 90–110 ppm monomethyl ether hydroquinone as inhibitor), acetic acid (glacial, ≥99.85%) and phosphate buffered saline tablets were purchased from Sigma-Aldrich Ireland Limited. 2,4,6-Trimethylbenzoyl-diphenyl-phosphineoxide (TPO) was generously donated by Bausch & Lomb Ireland. All other chemical and reagents used were of the highest purity grade commercially available.

Optical Transmittance Measurement Methods Synthesis of the Chitosan Nanoparticles Chitosan nanoparticles were synthesized by a cross-linking reaction mediated by sodium tripolyphosphate (TPP) according to the procedure reported by Calvo et al., with slight modifications [24]. Briefly, 50 mg of chitosan was dissolved in 50 ml of 1% (v/v) acetic acid solution and allowed to stir for approximately 30 min until a clear solution was obtained. A solution containing 10 mg of TPP dissolved in 5 ml DI H2O was added drop wise to the chitosan solution over a period of 30 min. The reaction mixture was stirred overnight at room temperature, followed by centrifugation at an acceleration of 24,000×g, for 30 min at 10°C using a Sigma 2–16 k centrifuge (Sigma Laboratory Centrifuges; Germany). The supernatant was discarded and the chitosan nanoparticle pellet was resuspended in 20 ml of DI H2O with mild sonication using a probe sonicator (Sonic Ultra Cell, VCX130PB, Sonics & Materials Inc; CT, USA). Finally, the nanoparticles were lyophilized by drying the frozen aqueous nanoparticle solution at −70°C and 0.01 mbar (VirTis benchtop 6 K freeze dryer; NY, USA) and stored at 4°C until further use. Dexamethasone sodium phosphate (DXP) loaded nanoparticles were synthesized following a similar procedure except the required amount of DXP was dissolved in an acetic acid solution of chitosan. The encapsulation efficiency (EE) and loading capacity (LC) were calculated according to the following formulae: EE ð%Þ ¼

total DXP added−free DXP  100 total DXP added

LCð%Þ ¼

total DXP added−free DXP  100 weight of chitosan

The transmittance measurements of the lens were carried out on a Schimadzu UV-2401 PC, UV-Visible spectrophotometer. Briefly, the lens was hydrated in 0.1 M phosphate buffer saline (pH = 7.4) for 24 h and then mounted between two glass slides. The slide with the contact lens was placed in the spectrophotometer with the help of a slide holder and transmittance values were measured at wavelengths ranging from 400 to 800 nm. In Vitro Drug Release Studies In vitro drug release studies were carried out using a Float-ALyzer (Spectra-Por®, MWCO 3.5–5 kDa) in phosphate buffer saline (pH = 7.4) at 37°C. Contact lens / 5 mg of chitosan nanoparticles were suspended in 2 ml PBS and introduced into the inner tube of the dialyzer. The dialyzer was then placed in a cylinder containing 5 ml release media with a magnetic stirring bar. The stirring speed was maintained at 50 rpm. At appropriate time intervals 100 μl samples were withdrawn from the release media and subsequently supplemented with 100 μl of fresh media. The samples were analysed on an Agilent 1200 Series system (Agilent Technologies, Waldbronn, Germany) equipped with a G1312B SL binary pump, G1329B autosampler, vacuum degasser and G1316B temperature-controlled column compartment. The mobile phase consisted of 50:50 acetonitrile: phosphoric acid buffer (pH = 3). A Waters XTerra MS C18 Column, 5 μm, 4.6 × 250 mm column was used with column temperature 25°C and a flow rate of 1 ml/min. A retention time of 1.8 min was observed for DXP at 241 nm. The calibration curve (R2 = 0.994) for DXP standards was obtained in the concentration rage of 0.08–20 μg/ml. The release study was carried out in triplicate (n = 3). The release data has been

Author's personal copy Behl et al.

presented as the mean ± standard error of three independent experiments. Characterization The SEM pictures were taken on a Hitachi S-2460N scanning electron microscope (SEM) fitted with a tungsten (W) filament. The samples were mounted on a metal stub using double sided carbon tape followed by coating with an 8 nm thick layer of gold (Au) by sputter coating for 30 min. Finally the pictures were taken at an accelerating voltage of 15 kV and observed using 500 and 200 nm scale bars. The hydrodynamic size was estimated at ambient temperature using a laser diffraction system Malvern Mastersizer 2000™ (Particular Sciences, Dublin, Ireland). Deionised water was used as the dispersion medium and measurements were carried out in triplicate. For atomic force microscopy (AFM), the aqueous nanoparticle suspension was mounted on a glass slide, air dried and observed in tapping mode under ambient conditions on a Bruker Dimension Icon® AFM Instrument. The FTIR (Fourier Transform Infrared) spectra were recorded using a Varian 660-IR series instrument by the KBr disc method.

RESULTS AND DISCUSSIONS Synthesis and Characterization of Dexamethasone Loaded Chitosan Nanoparticles Chitosan nanoparticles were synthesized by crosslinking with TPP. Generally, the particle formation has been attributed to the polyelectrolyte interaction between the ammonium group of chitosan and the anionic phosphate group of TPP, resulting in electrostatic cross-linking [26]. Figure 1 shows that spherical chitosan nanoparticles were obtained, as revealed by SEM and AFM micrographs. The particle size was found to be in the range of 50 nm using both techniques. Particle size analyses carried out in water, showed the hydrodynamic diameter of the particle was found to be in the range of 105 ± 11 nm, as presented in Table I. As the SEM and AFM measurements were carried out on dried samples, the observed increase in size could be attributed to the swelling capacity of the chitosan nanoparticles in water. In fact, this behaviour has been observed earlier with chitosan [27] and polyethylene glycolchitosan [28] nanoparticles, where a higher hydrodynamic diameter was observed as compared to TEM micrographs. The drug loading experiments were carried out with 20 wt%, 50 wt% and 100 wt% of DXP with respect to chitosan. Both loading capacity and encapsulation efficiency were increased from 4.19 to 43.06% and 20.97 to 43.06% respectively, when initial drug loading was increased from 20 to 100 wt% (Table I). Though a decrease in encapsulation

efficiency has been reported earlier with increase in initial drug loading content, this phenomenon was not observed in the present study [29,30]. In one of the reports involving encapsulation of an isomer of DXP, betamethasone sodium phosphate, the decrease in encapsulation efficiency was attributed to the increased potential of drug for diffusion into the external solution [30]. However, it is worth noting that the authors used very high concentrations of betamethasone sodium phosphate (0.2% w/v to 1% w/v solution) for encapsulation studies, whereas in this work the highest concentration of drug used was not more than 0.1%. The hydrodynamic diameter was found to increase with increase in drug loading from 20 to 50 wt%, whereas a further increase to 100 wt% resulted in a decrease in size to 132 ± 4 nm (Table I). Similar results have been observed earlier where the increase in size was attributed to the increasing drug content. Conversely, the decrease in size above a certain concentration of drug was attributed to enhanced electrostatic cross-linking between a negatively charged drug like DXP, and the cationic ammonium group of chitosan [30,31]. The FTIR spectra of chitosan, chitosan-TPP nanoparticles, DXP and DXP loaded chitosan nanoparticles are shown in Fig. 2. The characteristic peaks of chitosan at 3435 cm−1 and 1077 cm −1, corresponding to –OH and –C-O-C– stretching vibrations, were observed at 3404 cm−1 and 1082 cm−1 in chitosan nanoparticles. The 3404 cm−1 peak became broad, indicating enhanced hydrogen bonding [29]. The –NH2 bending vibration peak, observed in chitosan at 1642 cm−1, shifted to 1542 cm−1 in the case of chitosan-TPP nanoparticles and a distinct new peak appeared at 1636 cm−1. These results were in agreement with earlier observations with chitosan films modified by phosphate ions, and may be attributed to the electrostatic cross-linking between the phosphate group of TPP and the ammonium group of chitosan [29,32]. In the case of the FTIR spectra of DXP, peaks at 1716 cm−1, 1666 cm−1 and 1622 cm−1, corresponding to –C=O stretching, -COO − asymmetric stretching and -C=C- stretching vibrations, respectively, were observed in DXP loaded chitosan nanoparticles at 1716 cm−1, 1661 cm−1 and 1623 cm−1. All other peaks observed in chitosan-TPP nanoparticles were present in DXP loaded chitosan nanoparticles at their respective places, indicating that DXP had been loaded into chitosan nanoparticles [33]. In Vitro Release Studies of DXP from Chitosan Nanoparticles Prior to entrapment of the nanoparticles in the pHEMA lens, the release profiles of the DXP loaded chitosan nanoparticles at different loading capacities were studied in PBS (pH = 7.4) at 37°C (Fig. 3). Initially, the DXP release increased rapidly with time. Further increase in time resulted in a gradual increase in the amount of DXP released, with release ultimately

Author's personal copy CSNP-Laden Contact Lenses for Extended DXP Release Fig. 1 SEM (a) and AFM (b) micrographs of the chitosan nanoparticles.

assuming a plateau at the longest time studied, i.e. 35 days. This may be attributed to the higher initial concentration of the DXP inside the particles maintaining a higher diffusion gradient, which gradually decreased with time [34]. The release rate was found to increase with an increase in initial drug loading from 20 to 50 wt%, corresponding to 4.19 and 10.65% loading capacities, respectively. After 20 days, about 60% release was observed in nanoparticles with 4.19% loading capacity and nearly 90% release was observed from particles with 10.65% loading capacity (Table II). However this increasing trend in release rate was not observed in nanoparticles with higher initial drug loading of 100 wt%, corresponding to 43.06% loading capacity, where nearly 65% release was observed in 20 days. The increasing trend of release rate with increase in loading capacity is a common phenomenon [35,36] and could be due to the increasing concentration gradient of drug between the nanoparticles and the release medium [34]. As well as this, release rate has also been reported Table I Drug Loading Data and Hydrodynamic Diameter of the Chitosan Nanoparticles Initial DXP loading w.r.t. chitosan (%)

Loading capacity (%)

Encapsulation efficiency (%)

Hydrodynamic diameter (nm)

0 20 50 100

– 4.19 10.65 43.06

– 20.97 21.31 43.06

105 ± 11 131 ± 6 170 ± 8 132 ± 4

to depend on cross-linker concentration [31], polymer molecular weight and particle size [30,37]. Especially where charged drugs or polymers are involved, electrostatic interaction between the drug and polymer govern the rate of release [37,38]. Positively charged drugs have been shown to interact with the negatively charged carboxyl group of poly(lactic-coglycolic acid) (PLGA) resulting in decreased release rates compared to their counterparts [38]. In another instance, a decrease in release rate was observed with increase in drug content, similar to what was observed in this study. In that instance, the author studied the release profiles of betamethasone sodium phosphate, an isomer of DXP, from chitosan alginate nanoparticles [30]. There it was observed that a higher drug content resulted in reduced hydrodynamic diameter, as was observed in our case. This could be attributed to the enhanced electrostatic cross-linking due to an increased amount of drug, resulting in a lower swelling capacity and thereby demonstrating a decrease in the rate of release. In order to elucidate the mechanism of DXP release from the chitosan nanoparticles, various mathematical models were applied. The applicability of the Korsmeyer–Peppas [39] equation was tested up to 60% of the initial drug release. The zero-order, first-order [40], Higuchi [41], and Hixson– Crowell [42] equations were tested for the initial burst phase (i.e. the first 3 days) where the release was observed to be 23, 84 and 53% for particles with DXP loadings of 4.19, 10.65 and 43.06%, respectively. These models were also applied to the sustained release phase observed from the fourth day to the end of drug release (i.e. 35 days). The correlation coefficient (R2) values were calculated for the linear curves obtained by

Author's personal copy Behl et al. Fig. 2 Comparative FTIR spectra of chitosan (a), chitosan nanoparticles (b), DXP (c) and DXP loaded chitosan nanoparticles (d).

regression analysis of their respective plots [34]. The respective R2, diffusion exponent (n), and rate parameter (k) values are listed in Tables II and III. The diffusion exponent value ‘n’ obtained from the Korsmeyer-Peppas equation at three DXP loading capacities, i.e. 4.19, 10.65 and 43.06% were 0.88, 0.87 and 0.88, respectively, indicating that DXP release from the nanoparticles occurred through a non-Fickian or anomalous diffusion, i.e. the superposition of both the diffusion-controlled and swelling-controlled release phenomena. These results are in agreement with earlier reports on chitosan nanoparticles for the release of sumatriptan succinate [43] and zidovudine [44], where non-Fickian release was observed. Further, the value of the diffusion exponent (n~0.8) gives an indication that the release of DXP from chitosan nanoparticles did not primarily follow a zero-order release mechanism (R2~0.83–0.98 and 0.69–0.79 for initial burst phase and sustained release phase); rather a first-order release was followed, as confirmed by a higher R2 value for the first-order equation, i.e. 0.99, 0.98, 0.99 and 0.83, 0.88, 0.89 at 4.19, 10.65 and 43.06% drug loading capacities, respectively, for initial burst phase and sustained release phase (Tables II and III). This indicates that

the release of DXP is dependent on matrix drug load, which is in agreement with the observation of a higher initial release followed by achievement of the steady state. The Hixson– Crowell equation gives the information about drug release related to a change in surface area and diameter of the particle and the Higuchi equation describes the system where the rate of drug release is primarily controlled by a diffusion mechanism [45]. A linear relationship was also obtained in the case of the Higuchi and Hixson–Crowell equations and hence, drug release kinetics of this type may be governed by a mechanism involving erosion/diffusion, a change in surface area and diameter of the nanoparticles, as well as a change in diffusion path length during the release process. The applicability of all these equations further reveal that more than one mechanism is involved for the release of DXP from chitosan nanoparticles, as also indicated by the diffusion exponent value (n ~ 0.8) from the Korsmeyer–Peppas equation [34]. Similar release behaviour has been observed previously by other groups on different systems, where the Korsmeyer– Peppas equation has been applied along with other mathematical models to confirm the anomalous release mechanism from the value of the diffusion exponent ‘n’ [42,45–48]. Synthesis and Characterization of Nanoparticle-Laden pHEMA Contact Lens

Fig. 3 In vitro release profiles of the DXP loaded chitosan nanoparticles at 4.19, 10.65 and 43.06% loading capacities. Results are means ± standard error of three (n = 3) independent experiments.

In order to prepare an effective delivery system, DXP loaded chitosan nanoparticles were entrapped in a pHEMA lens. The nanoparticles were dispersed by ultra-sonication in the HEMA monomer, followed by the addition of the EGDMA cross-linker and TPO, the photoinitiator. This composition is typically used for the preparation of pHEMA-based soft contact lenses [25]. DXP loaded chitosan nanoparticles of higher loading capacity (43.06%) were entrapped in varying amounts i.e. 0 μg, 60 μg, 80 μg, 100 μg and 200 μg and transmittance

Author's personal copy CSNP-Laden Contact Lenses for Extended DXP Release Table II

Release Kinetics Data for Initial Burst Phase (i.e. First 3 Days) Obtained from Zero Order, First Order, Higuchi and Hixson-Crowell Equations

LC / EE (%)

First 3 days (Drug release during initial burst) Zero order rate constant K0 (h−1)

First order rate constant K1 (h−1)

Higuchi rate constant KH (h-1/2)

Hixson-Crowell rate constant (K3) (h-1/3)

*Korsmeyer–Peppas rate constant K (h-n), n

4.19 / 20.97

7.55 R2 = 0.98

0.08 R2 = 0.99

0.26 R2 = 0.98

0.12 R2 = 0.99

10.65 / 21.31

24.09 R2 = 0.83

0.58 R2 = 0.98

0.61 R2 = 0.95

0.65 R2 = 0.95

0.11 h-0.88 n = 0.88 R2 = 0.95 0.63 h-0.87 n = 0.87 R2 = 0.99

43.06 / 43.06

16.12 R2 = 0.95

0.23 R2 = 0.99

0.49 R2 = 0.99

0.31 R2 = 0.98

0.28 h-0.88 n = 0.88 R2 = 0.95

*Korsmeyer-Peppas equation values calculated for the initial 60% of drug release

of the contact lenses were measured in the wavelength range of 400–800 nm to ascertain their optical clarity [49]. The addition of the nanoparticles to pHEMA was not found to alter the optical transparency of the contact lens and, as Fig. 4b demonstrates, an average transmittance of 95–98% relative to the neat contact lens, was observed in all of the lens samples. Additionally, SEM micrographs clearly revealed that the surfaces of the neat and nanoparticle-laden pHEMA contact lenses were similar, though with nanoparticles clearly visible in the latter (Fig. 4c and d). The contact lenses were further characterized by FTIR spectroscopy, as seen in Fig. 5. The characteristic peaks of HEMA (Fig. 5a) at 3440 cm−1, 1720 cm−1, 1633 cm−1 and 1165 cm−1, corresponding to –OH, –C=O, -C=C- and –C-O-C– stretching vibrations were observed in the spectra of the pHEMA contact lens (Fig. 5b) at 3440 cm−1, 1720 cm−1 and 1159 cm−1, as was the disappearance of the peak at 1633 cm−1, indicating the consumption of -C=C- in the polymer [50]. The peaks corresponding to -COO− asymmetric stretching, -C=Cstretching, –NH2 bending and –C-O-C– stretching vibrations, in the case of DXP loaded chitosan nanoparticles were also observed in the spectra of nanoparticle-loaded pHEMA contact lenses (Fig. 5d) at 1661 cm−1, 1623 cm−1, 1537 cm−1 and 1075 cm−1. All other peaks of the contact lens were found Table III Release Kinetics Data for Sustained Release Phase

LC / EE (%)

to be present at their respective places in nanoparticle-loaded pHEMA contact lens. In Vitro Release of DXP from Nanoparticle-Laden pHEMA Contact Lens The main aim of the work was to entrap nanoparticles with maximum drug loading into the lens without compromising the optical clarity. The contact lenses with 200 μg of DXP loaded chitosan nanoparticles, with 43.06% DXP loading capacity, were chosen to carry out the release study, as they have maximum DXP loading and optimum optical clarity (95% transmittance relative to the neat contact lens). The release profile was studied in PBS at 37°C. The lenses were suspended in 2 ml PBS in the inner tube of a Float-A-Lyzer that was placed in a cylinder containing 5 ml PBS. Figure 6 shows the DXP release profile from the contact lens. A maximum of 48 μg (55.73%) of drug was released over a period of 22 days. The DXP release was found to increase continuously for 10 days and then followed a gradual increase with time, ultimately assuming a plateau at the longest time studied, i.e. 22 days. About 43 μg of DXP, corresponding to 50% release was observed in 10 days. The DXP release rate from the lens was found to be nearly three times lower when compared to chitosan nanoparticles of similar

4th day—end of therapeutic release Zero order rate constant K0 (h−1)

First order rate constant K1 (h−1)

Higuchi rate constant KH (h-1/2)

Hixson-Crowell rate constant (K3) (h-1/3)

4.19 / 20.97

0.78

0.016

0.11

0.019

10.65 / 21.31

R2 = 71 0.25

R2 = 0.83 0.019

R2 = 0.80 0.02

R2 = 0.82 0.015

43.06 / 43.06

R2 = 0.69 0.29

R2 = 0.88 0.009

R2 = 0.83 0.03

R2 = 0.88 0.009

R2 = 0.79

R2 = 0.89

R2 = 0.84

R2 = 0.89

Author's personal copy Behl et al. Fig. 4 (a) Pictures of pHEMA lens and nanoparticles entrapped pHEMA lens. (b) Optical transmittance of the pHEMA lenses with varying amount of DXP loaded chitosan nanoparticles, i.e. 0 μg, 60 μg, 80 μg, 100 μg and 200 μg. (c) SEM micrographs of pHEMA lens and (d) nanoparticle-loaded pHEMA lens.

loading capacity. About 43 μg of DXP, corresponding to 50% release, was observed in 10 days from contact lens, whereas 50% release was observed in 3 days from chitosan nanoparticles. This suggested that the pHEMA matrix of the contact lens further slowed down the release of DXP and thus acted as an additional barrier to diffusion [25]. Dexamethasone is a highly potent long acting drug requiring a far lower dosage compared to other intermediate and short acting glucocorticoids, i.e. nearly 5 times lower than Fig. 5 Comparative FTIR spectra of HEMA monomer (a), pHEMA lens (b), DXP loaded chitosan nanoparticles (c) and nanoparticleloaded pHEMA lens (d).

prednisolone, methylprednisolone and 25 times lower than hydrocortisone, to elicit a biological response [51]. Earlier studies with human corneal epithelium have shown the regulation of membrane associated mucins and protein levels by dexamethasone in the concentration range of 0.01 to 1 μΜ [52]. Inhibition of inflammatory cytokines in human corneal epithelial cell and fibroblast cell lines by dexamethasone has been observed in the concentration range of 0.1 to 10 μΜ [53].

Author's personal copy CSNP-Laden Contact Lenses for Extended DXP Release

consideration. This will also be investigated in the planned stability study, however, this was investigated in the aforementioned work by Tieppo et al. where the effect of autoclaving on ketotifen fumarate-loaded lenses was studied [5]. There, the authors observed no additional compounds in the autoclaved lenses and only a slight increase in drug concentration, determined to be a result of water evaporation.

Fig. 6 In vitro release profile of DXP from nanoparticles loaded pHEMA lens. Results are means ± standard error of three (n = 3) independent experiments.

Commercially available DXP eye drops contains nearly 1 mg/ml of the drug, i.e. 1.9 mM. This high concentration of ophthalmic drugs in eye drops has been attributed to their poor permeability across the corneal epithelium, as only 1– 5% of the active drug reportedly penetrates the eye [54,55]. The nanoparticle-loaded pHEMA contact lens released a maximum of 48 μg of DXP. Therefore, considering the normal tear volume to be about 6 to 10 μl, assuming no tear drainage and similar release behaviour as was observed in 2 ml of PBS, 48 μg of DXP (MW. 516.4) in 10 μl of tears, would theoretically be almost 9.3 mM, which is about five times higher than the concentration provided in the commercially available eye drops. The results presented here are in agreement with an earlier report involving the loading of DXP in pHEMA lenses, where a maximum of 3.9 mM DXP release was achieved and warranted to be clinically relevant and within the therapeutic index [9]. Further, looking at the 10 days release profile, the amount of DXP released was about 4.3 μg/day. The volume of an eye drop is 50 μl containing about 50 μg of DXP, and only 2.5 μg would able to absorb in the eye considering the low bioavailability from eye drops (1–5%). Therefore it is quite clear that about 72% more bioavailability of DXP may be expected from these contact lenses. While these are promising results, it is necessary to note the issues surrounding lens packaging, as highlighted by Kim et al., where the release of dexamethasone acetate after packaging for up to 2 months was seen to drop to below therapeutically relevant concentrations [55]. However, as observed by the authors of that work, the drug release behaviour remained the same, and since the concentration of DXP shown in the present study was demonstrated to be more than five times that delivered by eye drops, it would be hoped that sufficient drug could be loaded into the lenses to still retain a suitable concentration after packaging. Further work is necessary to determine this and extended stability and storage studies are in preparation to ascertain the long term effects of packaging. This will be reported in due course. As well as packaging, the effect of autoclaving, for example, must be taken into

CONCLUSIONS The present work was directed towards the development of nanoparticle-laden bandage contact lenses for the extended release of DXP. The development of prolonged drug releasing bandage contact lenses has been suggested as a viable alternative to eye drops for the delivery of ophthalmic drugs. Though the loading of drugs has been tested by soaking the contact lens in the drug solution, the method suffers from a lack of sustained release and lens clarity. Therefore, drug-loaded colloidal particles have been suggested for the entrapment of drug in the lens material, retaining transparency and maintaining a sustained release of therapeutics. In this study, optimisation and characterisation of DXP-loaded chitosan nanoparticles led to the ultimate preparation of contact lenses loaded with 200 μg of nanoparticles, while still retaining 95% optical clarity. These lenses demonstrated continuously increasing DXP release for 10 days, followed by a gradual increase with time, eventually plateauing at 22 days. This led to a maximum release of 48 μg of DXP, which provided a calculated 72% increase in bioavailability of the drug compared to eye drops, for the first 10 days. The results presented here suggest chitosan as a suitable material for the entrapment of the negatively charged DXP. It was postulated that the sustained release was achieved not only by diffusion and swelling control but that the accompanying electrostatic interaction also contributed significantly to the controlled release. The lens material was suitably transparent, with an average light transmittance of 95–98%. As stated, DXP release from the lens was sufficient to meet the anti-inflammatory levels and as such may be a viable alternative to eye drops in suitable patients. While the strategy looks promising, it should also be taken into account that the pHEMA based lens cannot be used continuously because of their low oxygen permeability. Therefore, the lenses are recommended to be taken off at night and cleansed, to clean from both protein and lipid deposition. The effect of these steps on the property of lens is still to be assessed and cytotoxicity yet to be evaluated. All these issues are currently being addressed and the results will be presented in a future publication.

Author's personal copy Behl et al.

ACKNOWLEDGMENTS AND DISCLOSURES The authors wish to thank Science Foundation Ireland (SFI) for their support and funding through Technology Innovation Development Award.

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