Abstract
purpose. To formulate and characterize a drug-eluting contact lens designed to provide extended, controlled release of a drug.
methods. Prototype contact lenses were created by coating PLGA (poly[lactic-co-glycolic acid]) films containing test compounds with pHEMA (poly[hydroxyethyl methacrylate]) by ultraviolet light polymerization. The films, containing encapsulated fluorescein or ciprofloxacin, were characterized by scanning electron microscopy. Release studies were conducted in phosphate-buffered saline at 37°C with continuous shaking. Ciprofloxacin eluted from the contact lens was studied in an antimicrobial assay to verify antimicrobial effectiveness.
results. After a brief and minimal initial burst, the prototype contact lenses demonstrated controlled release of the molecules studied, with zero-order release kinetics under infinite sink conditions for over 4 weeks. The rate of drug release was controlled by changing either the ratio of drug to PLGA or the molecular mass of the PLGA used. Both the PLGA and the pHEMA affected release kinetics. Ciprofloxacin released from the contact lenses inhibited ciprofloxacin-sensitive Staphylococcus aureus at all time-points tested.
conclusions. A prototype contact lens for sustained drug release consisting of a thin drug-PLGA film coated with pHEMA could be used as a platform for ocular drug delivery with widespread therapeutic applications.
Topical ophthalmic solutions, or eye drops, are currently the most commonly used method of ocular drug delivery, accounting for approximately 90% of all ophthalmic medications,
1 2 but they are very inefficient. Eye drops are administered by pulse delivery, which is characterized by a transient overdose, followed by a relatively short period of effective therapeutic concentration, and then a prolonged period of an insufficient concentration or underdose. Furthermore, each drop is diluted and washed away by reflex tearing; and the drop is dispersed by blinking, so that only 1% to 7% of the dose delivered from an eye drop is absorbed by the eye.
3 The remainder is either flushed onto the patient’s cheek or drained through the nasolacrimal system, where the medication is available for systemic absorption,
4 with potentially toxic side effects. The overdosing of ophthalmic solutions contributes to the ocular and systemic side effects of some ophthalmic drugs.
5 6 Because of its small surface area and its short contact time with topical drops, the cornea itself only absorbs a fraction of the dose of a drop that is delivered to the surface of the eye.
7 Moreover, patient compliance can be problematic with ophthalmic drops, especially among the elderly.
8 The rate of noncompliance in patients with glaucoma is between 24% and 59%,
8 9 even in long-term users.
10 Simplifying the drug regimen improves compliance,
11 but there are few alternatives to eye drops for local delivery of most ophthalmic medications prescribed after surgery.
A sustained release system for ophthalmic drugs could obviate many of these shortcomings. The design criteria for such a system include comfort, biocompatibility, and, ideally, zero-order kinetics (i.e., the release of a constant amount of drug per day), for an extended period. The concept of delivering drugs specifically through a hydrogel (contact lens) was introduced as early as 1960.
12 Ninety-three percent of eye care providers indicate that they would use a drug-eluting contact lens if it was added to their treatment armamentarium,
13 and 72% of eye care providers have used bandage contact lenses (which protect the cornea and promote re-epithelialization) adjunctively with topical antibiotic drops.
13
Several researchers have designed contact lenses for drug delivery. However, achieving constant drug release (zero-order kinetics) has been a difficult challenge. The uptake and release of medications from conventional soft contact lenses has been explored.
14 15 16 17 Drugs released by the contact lenses demonstrated nonlinear kinetics: a burst of drug is delivered during the first few hours, followed by declining, subtherapeutic levels of drug release in the subsequent hours. Little, if any, drug is eluted by the second day of use. Research has also focused on the controlled release of medications from delivery systems incorporated into a contact lens’ hydrogel material,
18 19 20 21 22 including copolymerizing the hydrogel, poly(hydroxyethyl methacrylate) (pHEMA), with other monomers in an effort to control the drug’s uptake and release properties.
23 Drugs have been released from microemulsions contained in hydrogel prototype lenses.
18 19 24 “Biomimetic” and molecularly imprinted hydrogels have been used to release medications.
21 25 Other researchers also investigated the use of molecularly imprinted hydrogel contact lenses for drug delivery.
26 27 Drug-containing liposomes immobilized onto the surface of contact lenses have also been studied, but demonstrated first-order kinetics.
28 29 30 Achieving sustained, long-term drug delivery at the normal physiological temperature, pH, and salinity of the human eye has remained a challenge.
A number of noncontact lens methods have also been explored, but none have achieved these goals. Ocusert (Alza Corp., Palo Alto, CA), which was designed for placement in the cul-de-sac, was the first marketed device to demonstrate zero-order kinetics.
31 Although now rarely used to treat glaucoma, pilocarpine was once delivered to the eye by Ocusert. It is no longer commercially available. Collagen shields, which absorb then slowly release a wide variety of medications,
32 are most commonly applied after the corneal epithelium is surgically removed. They help to promote corneal re-epithelialization and provide antibiotic prophylaxis.
33 However, they are not widely used by most surgeons because they are not optically clear, they are difficult and uncomfortable to self-insert (typically requiring topical anesthesia by an eye care provider), and they degrade quickly, limiting their use to 1 to 3 days.
In our design of a device that could deliver drugs to the eye with zero-order kinetics, we used a dual polymer system, composed of a polymer film containing the test compounds coated by a transparent polymer that is used in contact lenses. For the former, we used poly(lactic-
co-glycolic) acid (PLGA), a biodegradable polymer that is well known for its biocompatibility and its ability to control drug-release kinetics.
34 35 36 37 For the latter, we used pHEMA, which is not biodegradable.
38 39 Herein, we describe the formulation and characterization of such a drug-eluting contact lens for extended drug zero-order release.
PLGAs (65% lactic acid and 35% glycolic acid; Lakeshore Biomaterials, Birmingham, AL) were of 118-kDa (50:50%) and 18-kDa (50:50%) molecular mass (high and low molecular masses, respectively). Ciprofloxacin 0.2% (Cipro IV ready-for-use infusion solutions in 5% dextrose injection) was purchased from Bayer Pharmaceutical Corp. (West Haven, CT). A photoinitiator (Irgacure 2959) was obtained from Ciba Specialty Chemicals Corp. (Tarrytown, NY). Medium grade acrylic resin was obtained from the London Resin Co. (Reading, Berkshire, UK). Ciprofloxacin powder, fluorescein, HEMA and all the other reagents were purchased from Sigma-Aldrich (St. Louis, MO).
Clinical ocular-related
Staphylococcus aureus strains, obtained from the Massachusetts Eye and Ear Infirmary (MEEI) Clinical Laboratory, were recovered from human cornea, eyelid, and canaliculus infections; the minimal inhibitory concentrations for all bacterial isolates were determined by standard methods (National Committee for Clinical Laboratory Standards [NCCLS]).
40 MEEI-IB01 is a ciprofloxacin-resistant keratitis strain (minimum inhibitory concentration [MIC] >2 μg/mL). MEEI-IB03, -IB012, and -IB013 are ciprofloxacin-sensitive (MIC < 1 μg/mL) strains.
Test materials were placed in 15 mL of phosphate-buffered saline (PBS) pH 7.4 inside a 50-mL centrifuge tube and placed in a 37°C incubator with continuous shaking. The PBS was sampled and replaced completely at predetermined intervals. The amount of fluorescein released into the PBS medium was measured by using a UV/VIS spectrophotometer (Molecular Devices, Sunnyvale, CA) at a wavelength of 490 nm. Concentrations and masses of released fluorescein at each kinetic time point were calculated based on a calibration curve prepared with known fluorescein concentrations (R 2 > 0.99). Four individual contact lenses were tested for each reported formulation.
The release of fluorescein in the absence of a drug delivery device was tested by suspending 25 mg of free fluorescein powder in 15 mL of PBS. At the same time points at which release from the devices was sampled, the tubes were centrifuged, the supernatants assayed, and the pellets resuspended.
The mass of ciprofloxacin released into the medium was quantified with high-pressure liquid chromatography (HPLC; 110 series; Agilent Technologies, Palo Alto, CA). A dC18 analytical column (4.6 × 250mm; particle size 5 μm; Atlantis; Waters Corp., Milford, MA) was used with a mobile phase mixture composed of 10 mM phosphate buffer (pH 2.1) and acetonitrile. Ratios of acetonitrile to phosphate buffer were increased from 20% to 70% over 8 minutes and then returned to 20% over the next 2 minutes. The flow rate was set at 1 mL/min. The samples were filtered through 0.45-μm syringe filters and 20 μL of the samples were injected into the pre-equilibrated column. Ciprofloxacin concentrations were determined with a UV detector set at 275 nm, by correlating the measured peak areas with those measured for a series of ciprofloxacin standards (prepared from Cipro-IV solution; Bayer Pharmaceutical Corp.) freshly prepared for each HPLC run.
The fluorescein-PLGA films
(Table 1)composed of high-molecular-mass (118 kDa) PLGA produced thin and flexible films with a uniform appearance. The films made from low-molecular-mass (18 kDa) PLGA were also regular in appearance, but were less flexible and more difficult to remove from the fluoropolymer (Teflon; DuPont) wells. The thickness of the films varied little, measuring between 200 and 250 μm. Ciprofloxacin-PLGA (1:1 ratio, high [118 kDa] molecular mass) films were also thin and flexible and had an even distribution of ciprofloxacin throughout the film on gross inspection. The ciprofloxacin films measured between 215 and 235 μm in thickness. All the films became slightly less pliable after lyophilization.
The drug-PLGA films were easily contained between layers of pHEMA, as described in the Methods section, creating a prototype contact lens. When hydrated, the lenses were flexible and had an optically clear central aperture. Entrapping drug between the pHEMA layers in the absence of a PLGA film was technically difficult and frequently led to the spread of drug into the central optical aperture. A few of the prototype contact lenses containing the films developed bubbles within the pHEMA after 4 weeks of drug release.
Production of lens prototypes where fluorescein powder was deposited alone (i.e., without a polymeric film) was problematic. The powder tended to stray into the optic axis, a problem made worse when the second covering layer of pHEMA was added. In that context, UV polymerization of the second layer was less successful than that of the first layer, forming a tacky, rough, malformed, or incomplete surface.
All devices contained approximately 20 mg of either fluorescein or ciprofloxacin.
A prototype contact lens was created with a diameter and thickness that are within the range of dimensions found in commercially available contact lenses that released a medication with zero-order kinetics over an extended period. Drug release from the contact lens was significantly influenced by both the drug-PLGA film and by the pHEMA coating.
These prototype contact lenses provided zero-order kinetic drug release for 4 weeks, which is the longest duration for which contact lenses are currently approved. These lenses can clearly provide drug release for much longer periods, a goal that may be of value in the developing world, or in patients who require very extended therapy, as in the context of keratoprostheses, where lenses may be worn for months or years.
41 42 This study, performed in release conditions with the same pH, salinity, and temperature as the human tear film, compared favorably to previous reports on contact lens drug delivery. The rate of drug release could be controlled by changing either the ratio of drug to PLGA within the polymer film or by varying the molecular mass of the PLGA. These observations are consistent with the increased barriers to drug diffusion in films containing higher PLGA contents (i.e., a lower loading level of drug)
43 or from the better film-forming, high-molecular-mass polymers.
44 Although PLGA is biodegradable, it is unlikely that degradation of the polymer contributed significantly to the rate of drug release observed during the 4 weeks of release, since 50:50 PLGA typically degrades after 1 to 2 months and 65:35 PLGA after 3 to 4 months.
45
This drug release study was performed in such a manner that the concentrations of fluorescein and ciprofloxacin were always well below their solubility saturation limits, so that there was no impedance to efflux from the devices. Although one should be cautious when extrapolating in vitro release study results to an in vivo situation, it appears that this drug-eluting contact lens design could continuously deliver a therapeutic concentration of ciprofloxacin to the eye. The prototype contact lenses released 134 μg of ciprofloxacin per day (0.09 μg per minute). The human eye produces 2 to 3 mL of tears each day, equivalent to 1 to 2 μL per minute. If this contact lens released the same amount of drug per minute in an eye producing 3 μL of tears during the same time interval, then the drug would achieve a tear concentration of 31 μg/μL. This is well above the MIC90 for ciprofloxacin (2 μg / μL) against the most common ocular flora.
Ciprofloxacin showed complete killing of
S. aureus at inocula of 10
5 cells. The study demonstrated no change in the bactericidal activity of the ciprofloxacin in the release medium collected at the beginning, middle or end of the 4-week release study
(Table 2) . In addition, the ciprofloxacin released from all the lenses at the conclusion of the month demonstrated the same bactericidal activity
(Table 2) . With the use of inocula greater than 10
6 cells, rare ciprofloxacin-resistant bacteria emerged in all release medium tested. This is consistent with previous reports that more resistance emerges with higher bacterial loads.
46 47 48 Therefore, some conditions, such as corneal ulcers with purulent discharge, may require a higher ciprofloxacin concentration than was found in our release medium. However, as already mentioned, a higher drug concentration may very well be obtained in the eye, particularly in the post–contact lens tear lake—the fluid space between the cornea and the contact lens. Ultimately, in vivo studies are needed to confirm the final concentration of ciprofloxacin in a living eye and in the post–contact lens tear lake.
Ciprofloxacin was used here because of its broad-spectrum antibacterial properties and commercial availability, and because it is well studied as a topically applied ophthalmic antibiotic.
49 Because ciprofloxacin is associated with greater antibacterial resistance and the development of corneal crystal deposits after long-term use,
50 future studies may employ other fluoroquinolone antibiotics. Medications from other classes of antimicrobials, including antifungals, steroids, and antiallergy and -glaucoma medications, could also be incorporated into the contact lens drug-eluting platform given the flexibility of the film-forming method to incorporate drugs with a range of hydrophobicities and solubilities.
We used PLGA and pHEMA because both have been well studied and are FDA approved for ocular use. pHEMA is a polymer that has been extensively investigated and used by the contact lens industry since the 1960s.
38 39 PLGA’s safety profile has been well documented, and the U.S. Food and Drug Administration has approved PLGA for use in ocular drug delivery, including its use in intravitreous, subscleral, and subconjunctival routes of administration.
34 35 36 37 PLGA has also been incorporated into a device intended for topical drug delivery to the eye. Huang et al.
51 found no ocular irritation from a timolol-PLGA film placed in the inferior fornices of rabbit eyes.
Clearly, there are many aspects of this prototype drug-eluting contact lens that may require optimization before commercialization. Although contact lenses are available that are as thick as 1 mm (and a diameter up to 24 mm; e.g., Precision Sphere; Kontour, Inc., Hercules, CA),
52 like our prototype, thinner lenses would probably be better for comfort, oxygen diffusion, and other parameters. The polymeric footprint could be modified to improve oxygen diffusion, if that turns out to be a problem. For example, a thin, crescent-shaped film could be incorporated into the superior periphery of a large-diameter contact lens. The thin film would have a negligible effect on oxygen permeability and would be cosmetically undetectable, since it would be covered by the upper lid. We used the contact lens material pHEMA, which has lower oxygen permeability than some of the newer silicon hydrogels that constitute the material for some contact lenses.
39 Such materials could also be used for drug-eluting lenses. Finally, the manufacturing process could be optimized in many respects to minimize cost and increase efficiency.
One important concern in contact lens development is shelf life. Conventional contact lenses are often stored for many months at room temperature. It would be expected that the PLGA component of the prototype contact lens described here would degrade substantially and variably during that time. That problem could be circumvented by using a nondegradable polymer. (We used this particular polymer system primarily because of its safety and the fact that its drug-releasing characteristics are extremely well understood.) Even so, drug elution into the liquid storage medium would continue. One way to resolve that problem would be to store the lenses in medium that contains a concentration of drug adequate to stop drug efflux from the lens. Another would be to store the lenses in a dehydrated state, so that there would be no drug efflux, and no polymer degradation. Whether the latter approach would be detrimental to the lens’ optical properties remains to be determined.