Soft contact lenses for controlled ocular delivery: 50...

1141ISSN 2041-599010.4155/TDE.13.81 © 2013 Future Science Ltd Ther. Deliv. (2013) 4(9), 1141–1161

Review

Although scientific events are not always easy to date with certainty, there is a consensus that the use of contact lenses (CLs) as suitable platforms for controlled delivery of ophthalmic drugs was first postulated by Otto Wichterle and his Czech colleagues in a patent published in 1965 [1–3], soon after the first patent that described a semi-indus-trial method of preparing hydrogel CLs, filed in 1963 (FiguRe 1) [4]. In the US Patent 3,220,960 relating to “cross-linked hydrophilic polymers and articles made therefrom”, Wichterle and Lim specu-lated that “medicinal active substances such as anti-biotics may be dissolved in the aqueous constituent of the hydrogels to provide medication over an extended period”. The same year, Sedlacek published first experiments on the application of eye drugs with the aid of gel-CLs [5]. Subsequent research in this field has proved the relevant advantages that medicated CLs can offer in ocular therapy compared with conventional ocular treatments. However, almost 50 years later, the off-label use of CLs as drug depots remains infrequent [6] and the first clinical studies for commercialization of medicated CLs are still ongoing [201]. Delays in the development of these combination products (medical device plus drug) are mainly, but not only, due to the different origin of their parents (if a comparison with animated beings is allowed). Industries specializing in CLs and of pharma-ceutical dosage forms have traditionally evolved separately [7]. Thus, production of medicated CLs involves strong efforts of the manufactur-ers to reorganize research and to find financial returns. Additionally, regulatory issues regarding classification and evaluation of drug-eluting CLs were not clear [8]. The last decade has witnessed a tremendous advance in the numerous challenges

that the design of medicated CLs involve, such as: the loading of adequate amounts of the drug with-out perturbing biocompatibility, optical clarity and comfort; and the release of therapeutic drug concentrations with controlled profiles during the desired timeframe [9]. The large body of scien-tific research and the implementation of specific evaluation pathways in the regulatory process as a subclass of combination products [8,10] allows forecasting that, finally, there is a hopeful future for medicated CLs 50 years after their postula-tion. This Review will try to provide the reader an overview of the sequence of improvements in the design of medicated CLs, with particular attention on advances that occurred in the last decade. First, a short general outlook on ocular drug delivery is provided for a better comprehen-sion of the needs in the field. Then, the evolution in CLs for drug release is revisited, not forcedly in a temporal order, but from the simplest combina-tion of independently developed medical devices and drug solutions, to the most elaborated designs of ad hoc combination products.

Ocular drug deliveryInfections, allergies, vascular and degenerative processes, as well as accidental damages are relatively common threats for the eye structures [11,12]. Systemic treatment involves the admin-istration of large doses able to create sufficient serum–ocular tissue concentration gradients and to counteract, at least partially, the blood–ocular barriers (blood–aqueous humor, blood–retina, and blood–vitreous humor) against drug pen-etration. Even administering high drug doses, systemic delivery only helps in the management of diseases affecting the posterior segment of the

Soft contact lenses for controlled ocular delivery: 50 years in the making

The use of contact lenses as ocular bandages for drug delivery was envisioned nearly 50 years ago by Wichterle and co-workers. Despite the therapeutic advantages that can be obtained, this application has to face up to the poor affinity shown by commercially available contact lenses for most ophthalmic drugs, resulting in small amounts of drug being loaded and short time of therapeutic levels in the eye structures. Novel strategies that appeared in the beginning of 21st century, for example coating lenses with vitamin E, incorporation of drug nanocarriers or application of molecular imprinting technology, are becoming relevant tools for development of true drug/contact lens combination products that may be available for ocular therapy in the foreseeable future.

Clara González-Chomón, Angel Concheiro & Carmen Alvarez-Lorenzo*Departamento de Farmacia y Tecnología Farmacéutica, Facultad de Farmacia, Universidad de Santiago de Compostela, 15782-Santiago de Compostela, Spain *Author for correspondence: Tel.: +34 981 563 100 Fax: +34 981 547 148 E-mail: carmen.alvarez.lorenzo @usc.es

Review | González-Chomón, Concheiro & Alvarez-Lorenzo

Ther. Deliv. (2013) 4(9)1142 future science group

eye [13,14]. The anterior segment seems to be more accessible via external topical route, although also with notable limitations [14,15]. Additionally, topical administration may be a patient-friendly alternative to the intravitreal implants and trans-scleral drug-delivery systems for some diseases affecting the back of the eye [13,15].

One of the most important challenges of ocu-lar administration is to maintain the optimal drug concentration in the site of action dur-ing prolonged periods of time [9,15]. Eye drops are the most common form to deliver drugs into the eye, but less than 5% of the instilled dose effectively penetrates into the cornea [16]. Nasolachrymal drainage, epithelial membrane barriers and nonproductive absorption lead to poor ocular bioavailability and to undesir-able systemic adsorption, which could provoke side effects [9]. Eye-drop instillation produces a rapid burst where the drug concentration is well above the therapeutic concentration, followed by a short time of optimal concentration and a subsequent period of subtherapeutic concentra-tion until the next drop is administered [15,16]. To diminish the burst and prolong the contact

time between the drug and the eye surface, many approaches have been proposed, such as the use of penetration enhancers, mucoadhesive and viscous polymers, nanoparticles, suspensions and implants [17–20]. The interest for develop-ing ocular bandages dates back to 1886 when Galezowski successfully applied gelatine discs loaded with different drugs (cocaine and sub-limate) to prevent complications after cataract surgery [21]. Nevertheless, the poor optical per-formance and the limited possibilities of obtain-ing mechanically stable, sterile bandages limited their ocular application.

The search for systems that combine efficient control of drug release without compromising optical performance (e.g., not causing blurred vision) or patient adherence with drug therapy (e.g., avoiding uncomfortable feeling of foreign body) has been concurrent with the notable development in CLs, as explained in the next section. The feasibility of using CLs as drug plat-forms opens novel possibilities in eye therapy [6,9,22]. Soft CLs are flexible, slightly crosslinked hydrogels that can uptake drugs when soaking in eye-drop solutions and then sustain drug release to the postlens tear film – that is, the lachrymal fluid trapped between the cornea and the lens. Since turnover of the postlens tear film occurs slowly, the greater drug concentration attained should promote the pass through the cornea. Moreover, the release towards the lachrymal fluid that bathes the external surface of the CL is minimized by the dry periods during the blink-ing interval, and it is estimated to be five-times smaller than the amount of drug that diffuses towards the corneal epithelium [23]. Prolonged stay of the drug on the cornea leads to higher intraocular concentration, which in turn results in remarkable improvements in pharmacologi-cal response compared with conventional topical administration [24–27]. In some cases, drug deliv-ery through CLs can even allow a decrease in the dose required to attain the desired therapeutic effect [26,28]. However, the success of commercial CLs as drug-release platforms was rapidly found to be limited to drugs with medium affinity for the polymer components. If the affinity is too low, the loading is insufficient to attain thera-peutic levels; while if the drug–polymer interac-tion is strong, irreversible binding occurs and the drug is not released to the postlens tear film. These findings meant that interest in developing medicated CLs was kept to a minimum until the beginning of the 21st century, when novel design strategies were implemented (FiguRe 2).

Figure 1. First page extracts of the patents GB1,078,715 and US3,220,960. (A) The first semi-industrial production process for manufacturing hydrophilic contact lenses and, (B) their potential application as drug delivery systems.

Soft contact lenses for controlled ocular delivery: 50 years in the making | Review

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Medicated CLs �� The origins: CLs as medical devices

Development of poly(methyl methacrylate) in 1928 and commercialization in 1933 under the name Plexiglas® as a lightweight and shatter-resistant alternative to glass opened unprec-edented ways of using polymers for industrial and daily tasks. The high biocompatibility of poly(methyl methacrylate) was accidentally dis-covered during the Second World War, when it was observed that pilots of airplains having Plexiglas windows did not suffer rejection events when small splinters incrusted in their eyes [29]. The use of poly(methyl methacrylate) for hard CLs began in 1936. Soon after, the introduc-tion of more comfortable poly(2-hydroxyethyl methacrylate) (pHEMA) in 1954 triggered the popularization of CLs for vision correction [30,31]. These acrylic-based polymers combine the high optical clarity of the glass with lower density and better mechanical properties, and served as starting point for designing new copo-lymers able to improve the quality of the lenses [32]. Conventional CLs could be classified into hydrophilic (-‘filcon’) or hydrophobic (-‘focon’) lenses according to their chemical composition and physical properties, mainly to categorize behavior of the lenses in care product solutions and in contact with the proteins in the tear film (Table 1) [33,34,202]. Some hydrophilic CLs were approved in the late 1970s for extended wear (up to 6 nights and 7 days) due to their ability to absorb and retain large amounts of water (>35% w/w), in which the oxygen can dissolve and diffuse towards the cornea. During the 1980s, the search for materials that prevent hypoxia during closed-eye periods led to the development of silicone-based, gas-permeable, hard lenses, with greater oxygen permeability, which maintained the optical clarity and the simplicity of handling of hard lenses [31,32]. To overcome the strong hydrophobicity of these materials, high gas permeable silicone hydro-gels were prepared combining the large perme-ability of silicone components with the ocular biocompatibility and comfort of hydrophilic CLs. Some of them were approved in 2001 for continuous wear (up to 29 nights and 30 days) [35]. The time of permanence allowed on the eye surface is, thus, a critical point to consider when developing drug-eluting CLs, since the release profiles have to be fitted to the daily, weekly or monthly disposable lenses. Until recently, silicone hydrogel CLs were placed into one of the four existing US FDA classes, mostly into

group I, with the exception of balafilcon A that was assigned to group III (Table 1). However, their unpredictable way of reacting with disin-fectants and care solutions led to the Interna-tional Standards Organization to recognize that silicone hydrogels do not fit adequately within the existing FDA materials grouping and to pro-pose a specific ‘group V’ for this class of high-oxygen permeability hydrogel materials [203]. Splitting of ‘group V’ into various subgroups as for the hydrophilic conventional CLs is under study [204,205].

The FDA also classifies CLs according to their intended uses as: nontherapeutic CLs (e.g., for correction of refractive ametropia, aphakia and presbyopia); specialized-use CLs (e.g., for the treatment of keratoconus); and, therapeutic CLs that serve as a tool in the management of a wide variety of ophthalmic disorders refrac-tory to other treatment modalities [33]. Inter-estingly, drug delivery is included in the list of aims of therapeutic CLs, which are also intended for relief of ocular pain, promotion of corneal healing, mechanical protection and support, or maintenance of corneal epithelial hydra-tion [36–38,206]. Drug release could be pursued for treatment of diseases that occur disregard-ing the wear of CLs (e.g., seasonal allergies) or are directly associated to their use (e.g., risk of infection or dry-eye syndrome) [39–41].

1970 1975 1980 1985 1990 1995 2000 2005 2010

Year

Nu

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er o

f d

ocu

men

ts

35

30

25

20

15

10

5

0

Figure 2. Number of publications found in SciFinder® (bars) and in PubMed (line) databases for the concept ‘contact lens and drug release’, up to June 2013.



�� The match: commercial CLs & drug solutionspHEMA hydrogel meant a revolution not only on the popularization of CLs, but also on the scope of their applications. Since Sedlacek in 1965 [5] and Kaufman and co-workers in 1970 demonstrated that drug-loaded CLs can increase drug concentration in the anterior segment of the eye, many researchers have investigated how to load the CLs commonly used to correct ametropia problems with therapeutic substances [22,42]. Namely, the design of the medical device and the therapeutic substance formulation occur separately, and depending on the pathological process of the CL wearer, the clinician prescribes or prepares certain combinations of both com-ponents. Immersion in concentrated solutions of drug (eye drops) [42–46], instillation of the drug solution in the concavity of the lenses before placed on the eye, and instillation of the eye drops on the surface of the CLs after insertion (also named splash or postsoaking) [47–52] have all been evaluated. Some general rules on the variables affecting drug uptake are summarized in the following section [37].

Loading conditions & loading/release efficiencyThe time taken for penetration of a drug into a CL depends on the mesh size of the network (determined by the crosslinking density and the degree of swelling), the molecular size of the drug

and the concentration of the drug in the load-ing solution [53]. Drug uptake increases over the course of soaking in the drug solution, rapidly in the first 30–60 min and then more slowly for 24 h [44,54]. Attainment of drug-loading equilibrium throughout the CL is critical for it to perform as a sustained-release device; as the drug diffuses out from the surface, drug molecules localized in deeper parts would progressively replenish the surface layers enabling prolonged release [50,55]. For example, polymacon and alphafilcon A lenses loaded with ciprofloxacin by soaking for 1 hour in commercial eye drop solution led to remark-ably greater drug levels in both corneal tissue and aqueous humour at 6 h after application, com-pared with the direct application of eye drops (polymacon: 8.034 µg; alphafilcon: 6.432 µg, eye drops: 0.451 µg in cornea; polymacon: 0.361 µg; alphafilcon: 0.240 µg, eye drops: 0.0071 µg in aqueous humour) [207,208]. Similarly, etafilcon A (1-Day Acuvue®) CLs loaded by immersion in eye drop solutions provided higher concentrations of antifungals, aminoglycosides and fluoroqui-nolones in aqueous humour than when the eye drops were directly instilled [56–58]. Wearing of ofloxacin-loaded or ciprofloxacin-loaded lenses for 4–5 h before cataract extraction resulted in aqueous humour concentrations above the MIC

90

of Staphylococcus epidermidis at the beginning of surgery [58]. Nevertheless, there are many drugs (chloramphenicol, epinephrine and pilocarpine)

Table 1. Classification of contact lenses according to the US FDA.

Classification Group Description Examples: material (polymer composition; brand name)

Hydrophilic I Nonionic (<1% ionized groups at pH 7.2), low water content (<50%)

Crofilcon A (MMA-GMA; CSI 38), lotrafilcon A (DMA-siloxane macromer; Focus Night & Day™)†, polymacon (HEMA-NVP-CMA; Optima FW, Plano T, Soflens® 38)

II Nonionic (<1% ionized groups at pH 7.2), high water content (>50%)

Alphafilcon A (HEMA-NVP-CMA; Soflens 66); omalfilcon (HEMA-phosphorylcholine; Proclear® Compatibles); nelfilcon A (PVA; Focus Dailies®).

III Ionic (>1% ionized groups at pH 7.2), low water content (<50%)

Balafilcon A (Siloxane macromer-NVP; Pure Vision®)†

IV Ionic (>1% ionized groups at pH 7.2), high water content (>50%)

Etafilcon A (HEMA-MA; Acuvue®-2; HEMA-MA+PVP; 1-Day Acuvue Moist), ocufilcon B (HEMA-MA; Biomedics® 1 Day); methafilcon (HEMA-MA; Kontour 55), vifilcon A (HEMA-MA-NVP; Focus® Monthly)

Hydrophobic I Without silicon and fluorine Porofocon A (CAB; RX56), arfocon (t-butyl styrene; Airlens)

II With silicon but not fluorine Pasifocon (silicone acrylate; Boston® II and IV)

III With silicon and fluorine Itafluorocon A (FSC; Equalens), tisilfocon A (FSC; Menicon Z), paflucon (FSC; Fluoroperm®), flusifocon (FSC; Fluorex), enflucon B (FSC; Boston EO)

IV With fluorine but not silicon Fluorofocon A (PPFE; 3M Fluoropolymer)After International Standards Organization 18369–1 (2009) silicone hydrogel materials should be reclassified in a novel group V. All contact lenses shown as examples of the hydrophilic ones are available for therapeutic use or have been used in an off-label manner as therapeutic lenses [33,34,202].†Group V after International Standards Organization 18369–1 (2009).CAB: Cellulose acetate butyrate; CMA: Cyclohexyl methacrylate; DMA: N,N´-dimethyl acrylamide; FSC: Fluoro-siloxane acrylate; GMA: Glycerol methacrylate; HEMA: Hydroxyethyl methacrylate; MA: Methacrylic acid; MMA: Methyl methacrylate; NVP: N-vinyl pyrrolidone; PPFE: Polyperfluoroether; PVP: Poly(vinyl pyrrolidone).



that do not show affinity for pHEMA lenses, and in those cases the soaked lenses hardly led to one tenth of the aqueous humour concentration that can be achieved using eye drops [59–61]. In general, hydrophilic drugs scarcely adsorb on CLs. Cat-ionic substances have some affinity for anionic and hydrated methacrylic acid-based CLs (lens type IV, see Table 1), anionic drugs mainly interact with nonionic and hydrated N-vinylpyrrolidone (NVP)-based CLs (lens type II), and nonionic substances adsorb on hydrophobic silicone-based lenses (lens type III or the new group V) [206].

Lens thicknessIn the 1970s the increase in the hydrogel thick-ness was evaluated as a way to enhance the amount of drug loaded and to prolong the release. Thick pHEMA lenses (0.7–1.3 mm) provided sustained release of fluorescein, tetra-cycline or chloramphenicol up to 24 h [62–64] and prolonged in vivo elution of diethylenetriamine, disodium ethylenediamine tetraacetate or cys-tein hydrochloride, shortening the period of cor-neal burns healing by half compared with eye drops [65]. However, the increase in thickness has negative repercussions on oxygen permeability and patient comfort [66,67].

Water contentDrugs that do not interact with polymer compo-nents of the CLs are just hosted in the aqueous phase embedded in the network. Once the load-ing equilibrium is attained, the concentration of drug in the aqueous phase (w/w dry hydrogel) of the CL is the same as that in the loading solution [68]. Therefore, the greater the volume of water in the hydrogel, the greater the amount of drug loaded, as follows [69]:

Drug in aqueous phase WV C

p

s0= c m

equaTion 1

where Vs is the volume of the aqueous phase in

the hydrogel, Wp is the weight of dried hydrogel,

and C0 the drug concentration in the loading

solution. This explains that, for example, high water content (71%) Permalens lenses provide greater tobramycin concentrations in corneal tis-sue when the antibiotic is instilled on the lens, and for longer periods of time, than low water content (38.6%) Plano T therapeutic lenses [49]; although sustained release is in general more efficient from low water content lenses [55], as explained in the following. The content in water

also has repercussions on the oxygen permeabil-ity; large content in water enhances the perme-ability of hydrophilic CLs, but diminishes that of hydrophobic silicone-based lenses [70,71].

CLs deliver weakly interacting drugs via drug diffusion at a rate that can be predicted, if sink conditions are fulfilled, using the following equation [68]:

dtdM

l8DM exp l

Dt2 2

2= -r3 c m

equaTion 2

In this expression, M∞ represents the total amount of drug released, l the lens thickness and D the diffusion coefficient of the drug. In general, the release rate becomes faster for low-molecular-weight drugs and as the water content of the lens increases [68].

A clear example of the influence of the vari-ables mentioned above can be seen by compar-ing the ability to absorb and release drugs of silicon-containing (lotrafilcon and balafilcon) and pHEMA-containing (etafilcon, alphafil-con, polymacon, vifilcon and omalfilcon) com-mercial CLs (see characteristics in Table 1) [34]. Low molecular-weight drugs, namely cromolyn sodium, ketorolac tromethamine, and dexa-methasone sodium, were rapidly loaded and released (<1 h) from any of the lenses (all with a high content in water). Similar kinetics has been reported for cromolyn sodium, ciprofloxa-cin, idoxuridine, pilocarpine and prednisolone using vifilcon, etafilcon and polymacon lenses when placed in saline fresh solution [55]. By con-trast, ketotifen fumarate, a relatively hydropho-bic histamine, was more gradually taken up/released (for 2 h) because of the low concentra-tion (0.01 mg/ml) of the soaking solution and the slow mass transfer rate of the drug in and out the lens matrix. As explained above, the amount loaded depends on the content in water of the lenses and on the feasibility of interacting with polymer components (particularly, with ionic groups). Table 2 shows the total amounts of drug loaded and released by the different lenses. In general, pHEMA hydrogels with a high content in water and ionic character showed the greatest loading capability and also released more drug. Interestingly, only a fraction of the drug loaded was released in borate saline buffer unisol 4: 12% for cromolyn sodium, 1.5–28% for ketor-olac tromethamine and 46–67% for ketotifen fumarate, indicating that the drugs were in part irreversibly bound to the lens materials.



Greater and faster release from pHEMA-based soft CLs (SCLs; etafilcon A, omafilcon A, vifil-con A and polymacon) compared with silicone hydrogel SCLs (balafilcon A, comfilcon A, galy-filcon A, lotrafilcon B and senofilcon A) was also observed when loaded with natamycin by soaking in a solution of the antifungal agent in dimethyl-sulfoxide for 24 h. No differences were observed in the amount of natamycin loaded between both types of CLs, ranging between 500 and 1200 µg/lens. pHEMA-based CLs released approximately 15–30% drug in 1 h, while silicone hydrogel CLs released 2–10%, except balafilcon A that released 21%. Hydrophobic interactions of the drug with the silicone network are responsible for the greater retention. Although the control of the release was limited to 1 hour, the amount of natamycin released may be enough to attain the MIC

90 of

Candida spp and Fusarium spp [72]. A further study was carried out with ketoti-

fen fumarate and three conventional CLs (eta-filcon A, polymacon and alphafilcon A), four daily disposable CLs (nefilcon A, omafilcon A, etafilcon A and ocufilcon B), five silicone hydrogels (lotrafilcon B, balafilcon A, comfilcon A, galyfilcon A and senofilcon A) and two sili-cone daily disposable CLs (narafilcon A and fil-con II 3) [73]. When immersed in a 0.25 mg/ml ketotifen fumarate solution in unisol 4, all the lenses demonstrated a high loading capacity on the first 2 h. Etafilcon A (both Acuvue® 2 and 1-Day Acuvue Moist), balafilcon A and ocu-filcon B – that is, those bearing ionic groups showed the greatest uptake and were also the ones that released more drug (Table 3). Interest-ingly, all CLs released a fraction of the loaded drug in 1–3 h and then the release stopped due

to irreversible binding, failing to provide a sus-tained profile under sink conditions. The influ-ence of the drug-loading concentration on the total amount loaded and released was studied for ocufilcon B lenses. The concentration gra-dients between the loading solution and the lens aqueous phase (during soaking) and between the loaded lens and the release medium (dur-ing release) can be considered the main driving forces for the total mass transfer. This study [73] and another about loading and release of a phos-pholipid for management of dry eye syndrome [74] called the attention on two relevant points of the release tests: the volume of the in vitro release medium is usually much larger than that of the postlens tear film, which is in contrast under continuous flow/drainage; this may lead to unreal in vitro gradients that accelerate release compared with the in vivo situation; and the composition of the in vitro release test medium, usually water or a buffer does not resemble the composition of the tear fluid and may cause a delay in the release. In vivo, the interaction of the drug with the lipids, proteins and salts can reduce the concentration of free drug molecules in the medium and, thus, increase the driving force for drug diffusion out of the CLs [74].

�� The matchmakers: coadjutants for drug loading and controlled release Two different approaches have been explored to enhance the amount loaded and to control drug release of commercial CLs: coating with or incorporation of a drug-containing polymer film on/into the CLs; and adsorption of an amphi-philic molecule that regulates drug interactions with the CL network.

Table 2. Effect of contact lenses composition on the amounts of drug loaded, when immersed in the drug solution, and released in borate saline buffer.

Drug (soaking solution concentration)

Lens properties Average uptake (µg/lens)

Average release (µg/lens)

Cromolyn sodium (20 mg/ml)

Ionic vs nonionicSilicon vs pHEMA hydrogelHigh vs low water content

7548 vs 81267811 vs 79068066 vs 8126

390 vs 522338 vs 517552 vs 349

Ketotifen fumarate (0.20 mg/ml)


198 vs 123128 vs 166170 vs 135

103 vs 7784 vs 8990 vs 85

Ketorolac tromethamine (0.30 mg/ml)


103 vs 9985 vs 106108 vs 90

17 vs 2013 vs 2120 vs 16

Dexamethasone sodium phosphate (0.845 mg/ml)


67 vs 6857 vs 7174 vs 57

NANANA

pHEMA: Poly(2-hydroxyethyl methacrylate); NA:Not applicable. Reprinted with permission from [34] © Elsevier.

Key Term

Film-loaded contact lens: Contact lens that wraps a drug-containing polymer film that can provide controlled release.



The first approach has been tried with cipro-floxacin- and econazol-impregnated poly(lactic-co-glycolic) acid (PLGA), films that were added to the CL monomers solution before polymeriza-tion [75,76]. Since the films are not transparent to UV-vis light, a clear optical aperture should be left in the CLs. After a short burst, the film-containing CLs demonstrated zero-order release kinetics under infinite sink conditions for over 4 weeks, being able to inhibit the growth of sen-sitive strains of microorganisms. Drug release rate can be modulated by changing either the PLGA molecular mass or the drug/PLGA ratio.

In a comparative study carried out with Acu-ity CLs, antimicrobial agents were loaded by three different procedures: soaking, application of a coating of fibrin containing the drugs to the inner surface of the CLs, and sealing of the drug-containing fibrin film between two CLs (film-loaded CL) [77]. The CLs coated with or containing the films were able to provide sus-tained-release profiles, different from the rapid release recorded for the drug-soaked CLs. Once again the films resulted in alteration of the opti-cal performance of the CLs, and the two CLs sandwiched together would have a thickness not valid for practical use. Therefore, only partial coating of the CLs, near to the edges, could be feasible for clinical application in the prevention or treatment of ocular infections [77].

Adsorption of an amphiphilic molecule onto commercial CLs for enhancement of drug

interactions and regulation of the release pro-cess seems to be a simpler approach, with minor incidence on the optical performance of the CLs. For example, cationic surfactants (e.g., cetal-konium chloride) have been shown to adsorb through hydrophobic interactions on CL net-works (1-Day Acuvue), while the cationic moi-eties are free to interact with anionic drugs, such as dexamethasone 21-disodium phosphate used to treat persistent macular edema in retina. The surfactant and the drug were simultaneously loaded in the CLs by soaking in a solution of both components. As the amount of surfactant increased, the drug loading was larger and the release rate became slower [78]. Drug release was prolonged from 2 to 50 h in CLs loaded with 10% surfactant.

The use of vitamin E may be a more physio-logically friendly alternative to the use of cationic surfactants. This lipophilic vitamin has been shown to be suitable for protecting the eye from light-induced pathologies [79]. Vitamin E can be loaded in the CLs by soaking before loading of the drug or, alternatively, the loading of both vita-min E and the drug can simultaneously occur. The vitamin increases the CL/loading medium partition coefficient and also acts as a barrier for drug diffusion during release. As a consequence, drug uptake and release performance of both silicon hydrogel and pHEMA-based CLs can be notably improved. For example, loadings of 10% and 40% vitamin E prolonged the release

Table 3. Total amount of ketotifen loaded and released, plateau times and percentage release for all the commercial lenses.

Lens Total ketotifen uptake after 24 h (µg/lens)

Uptake plateau time (min)

Total ketotifen release after 24 h (µg/lens)

Release plateau time (min)

Release† (%)

Etafilcon A (Acuvue® 2)

461.78 ± 27.45 60 284.51 ± 29.83 120 61.7 ± 7.2

Polymacon 363.98 ± 41.41 60 114.3 ± 9.84 120 31.7 ± 4.6

Alphafilcon A 382.52 ± 15.32 20 138.98 ± 7.85 120 36.4 ± 3.3

Nelfilcon A 223.52 ± 14.33 360 40.36 ± 4.13 5 18.1 ± 1.4

Omafilcon A 237.56 ± 50.84 30 106.72 ± 6.03 60 48.8 ± 8.1

Etafilcon A (1-Day Acuvue Moist)

502.48 ± 10.19 120 249.93 ± 13.39 180 49.7 ± 2.7

Narafilcon A 350.81 ± 17.19 30 133.22 ± 13.61 120 38.1 ± 5

Ocufilcon B 502.94 ± 71.27 120 261.17 ± 13.82 240 47.5 ± 2.4

Filcon II3 309.03 ± 22.12 25 124.21 ± 12.61 60 40.1 ± 1.7

Lotrafilcon B 362.29 ± 40.13 180 158.98 ± 11.13 240 44 ± 2.4

Balafilcon A 515.85 ± 47.29 240 227.64 ± 14.71 120 45.4 ± 2.6

Comfilcon A 373.5 ± 56.33 60 110.38 ± 8.88 120 30.1 ± 5

Galyfilcon A 257.44 ± 57.96 180 139.34 ± 12.33 60 36.1 ± 1.6

Senofilcon A 339.82 ± 12.71 240 122.49 ± 1.61 240 49 ± 9†Percentage release is the total ketotifen uptake delivered divided by total release.Reproduced with permission from [73] © Wolters Kluwer Health.



of timolol by a factor of about five and 400, respectively, for NIGHT&DAY™ lens. Simi-lar results have been obtained for other drugs including fluconazole, dexamethasone 21-diso-dium phosphate and cyclosporine A (FiguRe 3) [80,81]. Vitamin E-loaded silicone CLs (lotrafil-con B; O

2OPTIX®) provided continuous anes-

thetics release for about 1–7 days and could be very useful for postoperative pain control after corneal surgery, such as the photorefractive keratectomy procedure for vision correction [82].

The success of the use of vitamin E as co-adjuvant of loading and controlled release of timolol in Acuvue® TruEyeTM has recently been proved in vivo using Beagle dogs affected by spontaneous glaucoma [83]. The in vivo release experiments were performed for: eye drops (150 µg/drop) placed twice a day; 60 µg drug-loaded control lenses replaced daily; 200 µg drug-loaded control lenses continuously worn for 4 days; and 200 µg drug-loaded vitamin E-containing lenses continuously worn for 4 days. Any CL reduced the intraocular pressure to a greater extent compared with the eye drops using a minor amount of timolol, and vitamin E-containing lenses were the only CLs able to provide a prolonged decrease of the intraocular pressure over the 4 days, in agreement with its most sustained release of the drug [83].

�� The specific design: CLs prepared ad hoc for drug delivery Commercial CLs present three important drawbacks to use as drug-delivery systems: the amount of drug loaded by the lenses is limited due to the fact that their affinity for the major-ity of the drugs is very low; most of the drug contained in the soaking solution is not loaded neither released by the CLs; and the release of the drug is too rapid and the period of time where therapeutic levels are maintained is quite short. Therefore, the design of true combination products comprising a CL and a drug requires novel approaches [6,84]. Below, the three most commonly investigated are described.

Incorporation of drug-loaded nanocarriers into CLs during manufacturingThis approach is based on the capability of nano-metric micelles, particles and liposomes to host drugs and to regulate their release. It is expected that, if an adequate proportion of nanocarriers is added to the components of CLs during manu-facturing, the nanocarriers will remain trapped in the CL structure, acting as reservoirs for sus-tained release. If the number and the size of the nanocarriers are sufficiently low, the CL stays transparent. Intensive research carried out in Chauhan’s group has demonstrated the feasibil-ity of this approach for a variety of drugs and nanostructures. Incorporation of microemul-sions or liposomes to pHEMA-based hydrogel CLs usually led to biphasic release profiles, with an initial burst of drug that was free in the network followed by a sustained release of drug remaining encapsulated. Microemulsion droplets and liposomes were shown to be respon-sible for the control of drug release, with minor effects from the crosslinking density and the lens thickness on the kinetics of the process [85,86]. Nevertheless, microemulsion droplets and lipo-somes exhibited limited physical stability and required the storage of the CLs in a drug-satu-rated buffer to prevent premature release. Immo-bilization of liposomes on the surface of the CLs has been tried by other groups, but the many steps involved in the process have prevented the scaling-up to date [87,88]. Drugs tested and their release conditions are specified in Table 4.

Incorporation of drug-loaded micelles and polymeric nanoparticles seems to be more promising. Micelle-containing CLs have been shown to be able to stand autoclaving and pack-ing in multipurpose solution, and still provide extended release of cyclosporine A at therapeutic

Time (day)

Dru

g r

elea

se (

%)

100

90

80

70

60

50

40

30

20

10

06050403020100

0 g vitamin E/g pure lens

0.1 g vitamin E/g pure lens

0.2 g vitamin E/g pure lens

Figure 3. Cyclosporine A release profile from vitamin E-loaded Acuvue® Oasys™ lenses. The drug was loaded in the lenses by soaking in 10 ml of 17 µg/ml drug solution for 30–60 days. The release profiles were measured in 1.75 ml sample that was replaced with fresh phosphate buffer pH 7.4 at every measurement. Data are plotted as mean ± standard deviation (n = 3). Reproduced with permission from [81] © Elsevier.



Table 4. In vitro release conditions of tests used to evaluate the performance of some drug-loaded contact lenses.

Drug Lens Medium (volume; ml)

Temperature and motion

Medium replacement

Duration of study

Ref.

Natamycin Balafilcon A, comfilcon A, galyfilcon A, senofilcon A, lotrafilcon B, etafilcon A, omafilcon A, polymacon, vifilcon A

Unisol 4 borate buffered saline pH 7.4 (5 ml)

32 ± 1ºC with continuous shaking

Replaced completely at 24 h

48 h [70]

Ketotifen Etafilcon A, polymacon, alphafilcon A, nelfilcon A, omafilcon A, etafilcon A, ocufilcon B, lotrafilcon B, balafilcon A, comfilcon A, galyfilcon A, senofilcon A, narafilcon A, filcon II 3

Unisol 4 borate buffered saline pH 7.4 (6 ml)

34ºC with continuous shaking

30 µl samples were extracted for measurements, without replacement

24 h [71]

Ciprofloxacin PLGA-coated pHEMA PBS pH 7.4 (15 ml) 37ºC with continuous shaking

Replaced completely at predetermined intervals

4 weeks [73]

Vancomycin and gentamicin

Drug in fibrin gel to coat lenses or to be trapped between two lenses

PBS (200 µl) filling the concavity of the lens

37ºC, static Replaced completely at predetermined intervals

72 h [75]

Dexamethasone 21-disodium phosphate

pHEMA PBS (6 ml) Room temperature, static

None 400 h [76]

Timolol, fluconazol, dexamethasone 21-disodium phosphate

Vitamin E-coated NIGHT&DAY lens

PBS (2.00 ml) Room temperature, soaking

None Up to 500 h [78]

Cyclosporine A Vitamin E-coated 1-DAY Acuvue® and silicone lenses


None Up to 50 days

[79]

Bupivacaine, lidocaine, tetracaine

Vitamin E-loaded lotrafilcon B (O2OPTIX) lens



Timolol Acuvue TruEye™ with vitamin E



Lidocaine Particle-laden pHEMA lens

Water (volume not specified)

Room temperature (movement not specified)

None 10 days [83]

Lidocaine Microemulsion-laden pHEMA lens

Water (volume not specified)

Room temperature (movement not specified)

None 10 days [84]

Levofloxacin Levofloxacin-loaded liposomes immobilized on hioxifilcon B lens

150 mM NaCl at pH 7.4 or a solution of 0.5% w/v Triton X-100 in water (3 ml)

37ºC (movement not specified)

None 140 h [86]

Cyclosporine A Brij 97 microemulsions or micelles – laden in pHEMA lens


Replaced every 24 h 20 days [87]

Cyclosporine A Brij surfactant-laden pHEMA lens


Replaced every 24 h 20 days [88]

1VI: 1-vinylimidazole; 4VI: 4-vinylimidazole; AAc: Acrylic acid; AM: Acrylamide; DEA: N,N-diethylacrylamide; DEAEM: Diethylaminoethyl methacrylate; DMA: N,N-dimethylacrylamide; HEAA: N-hydroxyethyl acrylamide; HEMA: 2-hydroxyethyl methacrylate; MAA: Methacrylic acid; MAPTAC: Methacrylamidopropyltrimethyl-ammonium chloride; MMA: Methylmethacrylate; MOEP: 2-methacryloxyethyl acid phosphate; NVP: N-vinylpyrrolidone; PBS: Phosphate-buffered saline; pHEMA: Poly(2-hydroxyethyl methacrylate); PLGA: poly(lactic-co-glycolic) acid; SiMA: 1-(tristrimethyl-siloxysilylpropyl)-methacrylate; TRIS: Tris(trimethylsiloxy)sililpropyl methacrylate.



Table 4. In vitro release conditions of tests used to evaluate the performance of some drug-loaded contact lenses (cont.).



Medium replacement

Duration of study

Ref.

Dexamethasone acetate

Silica shell crosslinked micelles laden in pHEMA lens

PBS (4.5 ml) 37ºC (movement not specified)

Replacement or not of aliquots taken for measurement

30 days [89]

Timolol Drug nanoparticles loaded in Acuvue Oasys™ lens

PBS (1.75–3.50 ml) 25–100ºC (movement not specified)

None 50 days [90]

Azulene p(HEMA-MAPTAC) optionally with MOEP and MAA

0.9% NaCl (5 ml) 37ºC, soaking Replaced every 2 h 8 h [91]

Naphazoline p(HEMA-MAPTAC) optionally with MOEP and MAA

0.9% NaCl (10 ml) 37ºC, soaking Not specified 40 h [92]

Ofloxacin pHEMA with MAA or MOESA and a silyl-containing monomer

0.9% NaCl (2 ml) Room temperature, soaking

Replaced at 1, 2, 4, 8 and 24 h

72 h [93]

Timolol pHEMA-silicone IPNs PBS (10 ml) 37ºC, 100 rpm Replaced at certain time intervals

16 days [95]

Puerarin pHEMA/b-CD lens Water (10 ml) 37ºC, 100 rpm 5 ml was replaced at certain time intervals

12 h [97]

Diclofenac pHEMA with pendant b-CD units

Artificial lachrymal fluid (10–20 ml)

25ºC, static conditions

None 20 days [98]

Timolol maleate p(DEA-MAA), pDEA, pHEMA, p(SiMA-DMA), and p(MMA-DMA) lenses

0.9% NaCl (10 ml) 37ºC, static conditions

None, or 0.6 ml was replaced at certain time intervals

13 days [24,109,110]

Timolol maleate pHEMA with MAA or MMA

0.9% NaCl pH 5.5, phosphate buffer pH 7.4 or artificial lacrimal fluid pH 8


None 12 h [108]

Timolol maleate p(DMA-TRIS-MAA) lenses 0.9% NaCl (10 ml) 25ºC, static conditions

0.6 ml were replaced at certain time intervals

7 days [111]

Timolol maleate p(HEMA-AAc) lenses 0.9% NaCl (2 or 8 ml) 25ºC, static conditions

None 22 days [112]

Norfloxacin p(HEMA-AAc) lenses Artificial lacrimal fluid (10–15 ml)


None 24 h [113]

Diclofenac p(HEMA-DEAEM-PEG200DMA)

Artificial lacrimal fluid (450 ml for infinite sink conditions or microfluidic device)

34ºC and 45 rpm for infinite sink conditions; or continuous physiological flow

None/flow conditions (3 µl/min)

140 h [114]

Prednisolone acetate

p(HEMA-MAA) lenses 0.9% NaCl or artificial lachrymal fluid (10 ml)

37ºC (movement not specified)

None 48 h [115]

Ciprofloxacin P(HEMA-TRIS-AAc) lenses Artificial lachrymal fluid (2 ml)

34ºC, shaking None 14 days [116]

Ketotifen fumarate

pHEMA lenses combining AAc, AM, NVP and PEG200DMA

Water or artificial lachrymal fluid solely or containing lysozyme (30 ml)

120 rpm None 5 days [117]




rates [89]. The hydrophobicity of the micellar core plays a relevant role on the drug partition. For example, in a study carried out with Brij surfac-tants (C

18H

x[OCH

2CH

2]

yOH), namely Brij 78

(x = 37, y = 20), Brij 97 (x = 35, y = 10), Brij 98 (x = 35, y = 20) and Brij 700 (x = 37, y = 100), cyclosporine A showed a high partitioning inside the micelles of Brij 700 and Brij 78. By contrast, the micellar partition coefficients of dexametha-sone and dexamethasone-21-acetate were smaller and the lenses could not effectively extend the release [90]. Reinforcement of the micellar struc-ture has been evaluated for loading of dexa-methasone acetate. The drug was encapsulated in methoxy poly(ethylene glycol)-block-polycap-rolactone micelles and the shell of the micelles was crosslinked, resulting in stable rod-like struc-tures. pHEMA hydrogels were prepared with 2, 4 and 6% wt of crosslinked dexamethasone-loaded micelles and with a thickness of 100 or 200 µm. Compared with control pHEMA hydrogels to which noncrosslinked micelles were added and that released 97% drug in 10 h, the hydrogels with the drug encapsulated in the crosslinked micelles prolonged the release for several days (FiguRe 4; see Table 4 for release conditions) [91].

Since sterilization and conditioning processes may trigger premature leakage of a certain amount of the preloaded drug, incorporation of the nanocarriers after silicone hydrogel CL syn-thesis has also been recently explored [92]. Timo-lol was encapsulated in propoxylated glyceryl triacrylate nanoparticles, and then the nanopar-ticles incorporated to the CLs either during or after synthesis. This approach involved the soak-ing of Acuvue Oasys lenses in 3% solution of the particles in ethanol for 24 h, followed by soaking in buffer phosphate pH 7.4 for other 24 h to remove the ethanol. CLs loaded either during

or after synthesis showed similar timolol release profiles (FiguRe 5). However, nanocarrier-laden CLs stored into packaging medium in a refrig-erator for 2 days experienced an important pre-mature leakage and, thus, led to lower amounts released. In vivo evaluation (Beagle dogs) of Acu-vue Oasys lenses freshly loaded with timolol-con-taining particles demonstrated their capability to reduce intraocular pressure for 3 days – that is, the period of time that the release is faster and therapeutic levels can be attained [92].

Table 4. In vitro release conditions of tests used to evaluate the performance of some drug-loaded contact lenses (cont.).



Medium replacement

Duration of study

Ref.

Ketotifen fumarate

pHEMA lenses combining AAc, AM, NVP and PEG200DMA

Artificial lacrimal fluid (30 ml for infinite sink conditions or microfluidic device)

120 rpm for infinite sink conditions; or continuous physiological flow

None/flow conditions (3 µl/min)

7 days [119]

Acetazolamide pHEMA or p(DMA-NVP) lenses combining 1VI or 4VI and HEAA with a Zn2+ source

0.9% NaCl (5–10 ml) 25ºC, static conditions

None 8 days [120,122]


0 5 10 15 20 25 30Time (days)

100

80

60

40

20

0

Cu

mu

lati

ve D

MS

A r

elea

se (

%)

DMSA-gel

2% DMSA-SSCM-gel

4% DMSA-SSCM-gel

6% DMSA-SSCM-gel

Figure 4. Dexamethasone release profiles from 200 µm thick hydrogels containing 0, 2, 4 or 6% of silica shell crosslinked micelles. The hydrogel without micelles attained the plateau at 50 h, releasing 90% of the drug loaded. Hydrogels that contained crosslinked micelles loaded with dexamethasone were able to control the release over 30 days. Reproduced with permission from [91] © John Wiley and Sons.

Key Term

Nanocarrier-laden contact lens: Contact lens with trapped drug nanostructures (e.g., liposomes, micelles, vesicles or particles) that regulate drug release from the contact lens matrix.



Incorporation of monomers with enhanced affinity for the drugCopolymerization of the monomers used to fab-ricate CLs with small proportions of other mono-mers bearing specific groups to interact with cer-tain drugs has also received some attention. The nature and the proportion of functional mono-mers should be chosen taking into account the compromise between the achievement of the func-tionality and the maintenance of the morphology, dimensions, optical clarity and biocompatibility of the CLs. pHEMA hydrogels copolymerized with ionizable monomers, for example, meth-acrylamide propyltrimethylammonium chloride, 2-methacryloxyethyl phosphate or methacrylic acid (MAA), have been shown to be able to cap-ture/release small molecules using ion-exchange reactions [93,94]. When the CL is immersed in a solution of an ionizable or protonizable drug, for example, azulene or naphazoline, this can interact with the oppositely charged groups of the net-work. Once applied on the eye, the CL releases the

drug by an exchange with Cl- or Na+ ions of the tear fluid. As an example, pHEMA-based hydro-gels prepared incorporating an anionic monomer (MAA or 2-methacryloyloxyethyl hydrogen suc-cinic acid, MOESA, in 1–10% wt) and a silyl-containing monomer (3-methacryloxypropyltris(trimethylsiloxy)silane (MPTS), have shown enhanced ability to take up ofloxacin, compared with neutral and cationic (methacrylamide prop-yltrimethylammonium chloride) hydrogels. The CLs were soaked in 0.3% ofloxacin and 0.85% NaCl pH 6.5 solution for 3 h. The amount of ofloxacin loaded was dependent on the pH, being higher at pH 6.5 and 7 and lower at pH 5 and 8 due to the zwitterionic character of the drug. The release was sustained for several hours in 0.9% NaCl. Anionic hydrogels that did not bear silyl-groups released 85% drug in 4 h (FiguRe 6), while those having silyl-groups showed a slower release [95]. p(HEMA-MPTS) lenses containing MAA could also uptake other antibiotics (gati-floxacin and moxifloxacin) at a greater extent than

Time (days)

Time (days)

∆ IO

P (

OS

-OD

)

8

7

6

5

4

3

2

1

0

6543210

-1

-2

-3

Baseline

Contact lenses Dru

g r

elea

se (

µg

)

0 5 10 15

160

140

120

100

80

60

40

20

0

Figure 5. Pharmacodynamic effect of extended wear of the Acuvue® Oasys™ lenses loaded with timolol-containing nanoparticles. The difference between the IOP of the untreated (OS) and the treated eye (OD) is plotted on the y-axis as mean ± standard deviation (n = 10 for the control study and n = 10, 6, 6, 6, and 4 for days 1–5 for the lens study). The inset demonstrates the in vitro drug release profiles from the same lenses at approximately 37°C. Data are shown as mean ± standard deviation (n = 3). IOP: Intraocular pressure; OD: Right eye; OS: Left eye. Reproduced with permission from [92] © Elsevier.

Key Term

Cyclodextrin-grafted contact lens: Contact lens with pendant cyclodextrin moieties suitable for hosting drugs and regulating their release through complex inclusion formation.



etafilcon A and polymacon, and provided in vivo drug concentrations in the cornea and the aqueous humor that were higher than those attained with commercial eye drops [96].

Interpenetration of hydrophilic pHEMA and hydrophobic silicone networks has been explored as a way to prepare timolol-eluting CLs [97]. Increasing the content in silicone net-work up to 35%, the oxygen permeability was notably enhanced (40.5–118 barrer) while the content in water (40–26%) and the uptake of timolol decreased (35.2–26 mg/g when soaked in 2 mg/ml, and 16.14–13.30 mg/g when soaked in 1 mg/ml). The release rate (see conditions in Table 4) was also slower for CLs with larger content in silicone network.

An alternative is to increase the affinity of the CL networks for relatively hydrophobic drugs without inducing irreversible binding consists in the use of cyclodextrins. These oligosaccharides form inclusion complexes with a variety of drugs and maintain or even enhance this capability when they are fixed in a polymeric network. The affinity of the cyclodextrins for the drug can then be exploited as the driving-force for the loading and the control mechanism for the release [98]. Cyclodextrins can be incorporated during CL synthesis, as a monofunctionalized b-cyclodex-trin [99], or in a second step once the CL has been prepared [100].

pHEMA networks incorporating monofunc-tionalized b-cyclodextrin have shown improved capability to uptake puerarin, an active sub-stance that may alleviate glaucoma and ocular hypertension. In in vivo experiments (in rabbits), pHEMA/b-CD3 (the CL with the highest cyclo-dextrin content) exhibited longer mean residence times of puerarin in tear fluid than those provided by pHEMA CLs and 1% puerarin eye drops. The puerarin concentration in the aqueous humour of rabbit reached a maximum of 0.81 µg/ml after wearing the pHEMA/ b-CD3 CLs soaked in 0.802 mg/ml solution for 4.81 h (FiguRe 7) [97]. However, copolymerization of cyclodextrin monomers usually results in an increase in the rigidity of the network and a change in water content [101]. Postfunctionalization with cyclodex-trins of already preformed lenses aims to avoid changes in the mechanical and optical features of the starting CLs [100]. Cyclodextrin units can be anchored to pHEMA networks having small proportions of glycidyl methacrylate. To do that, the hydroxyl groups of cyclodextrin react with the epoxy groups of glycidyl methacrylate under mild conditions. Pendant cyclodextrins have

been shown to be able to reduce the friction coef-ficient of the hydrogels, to enhance 15-fold the amount of diclofenac loaded, to prevent a prema-ture discharge of the drug in multipurpose stor-age liquids and to sustain drug release (Table 4) for 2 weeks [100]. If loaded with miconazole, the cyclodextrin-grafted CL completely prevented Candida albicans biofilm formation in vitro [102].

Molecular imprinting techniqueThe success of the incorporation of monomers with enhanced affinity for the drug can be further improved applying molecular imprinting, which pursues the optimization of the spatial distribu-tion of the functional monomers to achieve the maximum efficiency of the interactions between the drug and the polymeric network. The tech-nique consists of using the drug molecules as tem-plates during polymerization in order to induce the arrangement of the monomers as a function of their affinity for the drug. The conformation of the monomers is fixed during polymerization and, once the drug molecules that acted as templates are removed, cavities (named imprinted pockets) with the size and the most suitable chemical groups to interact with the drug are obtained. Imprinted networks can recognize with greater affinity and specificity the drug molecules during soaking, compared with the networks synthesized in the absence of template (nonimprinted networks).

Am

ou

nt

of

rele

ased

dru

g (

mg

/g le

ns)

Release time (h)

45

40

35

30

25

20

15

10

5

00 5 10 15 20 25

0% 1% 3% 5% 10%

Figure 6. Ofloxacin release profiles from contact lenses with 0, 1, 3, 5 and 10% wt of methacrylic acid. 85% of the drug loaded was released from the contact lenses within 4 h. Reproduced with permission from [95] © John Wiley and Sons.



This technology, which was initially developed for endowing rigid highly crosslinked polymeric systems with the ability to recognize target spe-cies [103,104], has been successfully adapted to the synthesis of medicated CLs [105,106].

The formation of the imprinted pockets strongly depends on the stability and solubility of the functional monomers–template assemblies during polymerization. If the molar ratio in the complex is not appropriate or if the assemblies dissociate to some extent during polymerization, the functional monomers would not correctly arrange around the template molecule, resulting in a small difference between imprinted and non-imprinted (conventional) CLs [107–109]. More-over, the low crosslinking density of CLs and their swelling in the preservation liquids or in the lachrymal fluid may compromise the physical stability of the imprinted pockets. This drawback can be overcome maximizing the interactions of drugs and the network in terms of number of binding points that can synergistically favor the binding process. Research on drug-imprinted CLs has focused on several therapeutic groups, namely b-adrenergic antagonists (timolol), antimicrobials (norf loxacin), antihistamines (ketotifen), carbonic anhydrase inhibitors (acetazolamide) and anti-inflammatory drugs.

CLs imprinted for timolol were the first successfully proven in vivo (in rabbits) [26]. A

detailed study of the influence of the nature of the functional monomer and the backbone monomer, the degree of crosslinking and the drug:functional monomers ratio on loading and release performance of the lenses has been car-ried out [110–114]. Timolol-imprinted ultrathin N,N-diethylacrylamide-based lenses (14 mm diameter and 0.08 mm center thickness) loaded 34 µg drug per lens, while the conventional lenses took up 20 µg drug per lens. Timolol levels in rabbits’ lachrymal fluid were monitored after the insertion of imprinted and nonimprinted lenses or after the instillation of timolol eye drop solu-tions of 0.068% (total dose 34 µg) or of 0.25% (commercial solution, total dose 125 µg; FiguRe 8) [26]. The imprinted and the nonimprinted CLs displayed the maximum ocular level at around 5 min, followed by monoexponential decay, which was prolonged for 180 min in the case of the imprinted CLs, compared with the 90 min of the nonimprinted ones. Timolol applied as drops was flushed out of the eyes in less than 60 min. Furthermore, imprinted CLs led to 3.3-fold and 8.7-fold area under the timolol concentration–time curve than the nonimprinted ones and eye drops, respectively. These results proved that imprinted CLs reduce the precorneal elimination of the drug compared with the eye drops and, as a consequence, a much smaller amount of drug is needed to achieve the desired therapeutic levels.

Pu

erar

in c

on

cen

trat

ion

in t

ear

flu

id (

µg

/ml)

Time (min)

250

200

150

100

50

0-50 0 50 100 150 200 250 300 350 400

Pu

erar

in c

on

cen

trat

ion

in t

ear

flu

id (

µg

/ml)

Time (min)

0-50 0 50 100 150 200 250 300 350 400

200

400

600

800

1000

1200

1400HEMA CLspHEMA/β-CD3 CLs

1% puerarin eye dropspHEMA/β-CD3 CLs

Figure 7. Dependence of puerarin concentration in tear fluid on the time worn. (A) pHEMA/b-CD3 CLs presoaked in 0.334 mg/ml puerarin solution, pHEMA CLs used as control and (B) presoaked in 0.802 mg/ml puerarin solution, 1% eye drops used as control. Each point represents the mean ± standard deviation (n = 4–5). CL: Contact lens; pHEMA: Poly(2-hydroxyethyl methacrylate). Reproduced with permission from [99] © Elsevier.

Key Term

Drug-imprinted contact lenses: Contact lens in which the monomers adopt a designed spatial arrangement by using drug molecules as templates during polymerization. Such an arrangement allows the creation of pockets (imprinted cavities) with the size and chemical groups suited for hosting with the target drug.



pHEMA-based hydrogels prepared with norfloxacin:acrylic acid (AAc) molar ratios rang-ing from 1:2 to 1:16 revealed that imprinted hydro-gels synthesized using the optimal 1:4 molar ratio (as estimated from calorimetric titration) released norfloxacin at a rate 3.5-times lower than non-imprinted hydrogels, providing a 24-h controlled release. Imprinted hydrogels of various contents in AAc and different thicknesses exhibited similar loading/release behavior, evidencing the robustness of the imprinting technique [115]. The relevance of choosing the adequate template:functional mono-mer molar ratio has also been demonstrated for CLs imprinted for diclofenac [116], prednisolone acetate [117] and ciprofloxacin [118]. In this later study, imprinted p(HEMA-TRIS) lenses prepared with high functional monomer (MAA):template ratio (8:1 and 16:1) sustained release for longer, although released less drug than those prepared with lower monomer/template ratio (4:1).

The Byrne group has designed ketotifen-imprinted CLs using functional monomers that have chemical groups that resemble the structure of the amino acids present in the physiological histamine H1-receptor [119]. The guiding hypoth-esis was that if antihistamines have high affin-ity for the H1-receptor, a hydrogel with similar chemical functionality would bind the anti-histamine tightly, increasing the loading and delaying release kinetics. In fact, CLs prepared with pHEMA (92 mol%), crosslinker (5 mol%) and the most biomimetic monomers (3 mol%), namely AAc, NVP and acrylamide (AM), designed as poly(HEMA-co-AAc-co-AM-co-NVP-co-PEG200DMA), demonstrated six- and three-fold greater loading than conventional CLs and networks containing one or two functional monomers, respectively. Release of therapeuti-cally relevant amounts of ketotifen in artificial lachrymal fluid containing lysozyme (see Table 4 for experimental details) was sustained for several days [120,121]. In vivo studies (in rabbits) carried out with ketotifen-imprinted poly(HEMA-co-AAc-co-AM-co-NVP-co-PEG200DMA) CLs (100 µm thickness, 11.8 mm diameter) evi-denced the possibility of attaining a constant tear film concentration of 170 ± 30 µg/ml (well above the effective concentration of 30 µg/ml) for 26 h, after an initial burst in the first 4 h with C

max of 214 µg/ml. Nonimprinted CLs provided

a Cmax

of 140 µg/ml in the first hour and then drug concentration decreased exponentially within 10 h (FiguRe 9). Eye drops quickly pro-vided a C

max of 143 µg/ml, but the drug disap-

peared from the tear film in less than 45 min.

500

400

300

200

100

0

0 30 60 90 120 150 180

Time (min)

Tim

olo

l co

nce

ntr

atio

n (

mM

)

Imprinted contact lenses (34 µg)Nonimprinted contact lenses (21 µg)

0.25% timolol (125 µg)

0.068% timolol (34 µg)

Figure 8. Timolol tear fluid concentration–time profiles after application of drug-loaded imprinted and nonimprinted contact lenses or instillation of eye drops. The doses were 34 µg for imprinted contact lenses, 21 µg for nonimprinted contact lens, and 34 µg and 125 µg when 0.068% or 0.25% timolol eye drops were instilled, respectively. Each point represents the mean ± standard deviation (n = 3–5). Reproduced with permission from [26] © Elsevier.

0 2 4 6 8 10 12 14 16 18 20 22 24 26

300

250

200

150

100

50

0

Time (h)

Co

nce

ntr

atio

n o

f K

eto

tife

n te

ar fl

uid

(µ

g/m

l)

Imprinted contact lensesNonimprinted contact lensesEye drops

Figure 9. In vivo ketotifen fumarate tear fluid concentration profile from imprinted poly(2-hydroxyethyl methacrylate-co-acrylic acid-co-acrylamide-co-N-vinylpyrrolidone-co-PEG200DMA) lenses (squares), nonimprinted lenses (circles) and eye drops (0.035% solution Zaditor®; triangles) in white New Zealand rabbits. Reproduced with permission from [27] © Elsevier.



Therefore, ketotifen-imprinted CLs rendered notably higher mean residence time, area under the concentration time curve and bioavailability than conventional CLs and eye drops [27].

The bioinspired selection of functional mono-mers has also been applied to the design of CLs imprinted for carbonic anhydrase inhibitors (CAIs; e.g., acetazolamide, ethoxzolamide) [122]. The active site of carbonic anhydrases consists of a cone-shaped cavity that contains a Zn2+ ion coor-dinated to three histidine residues in a tetrahedral geometry with a solvent molecule as the fourth ligand [123]. Acetazolamide inhibits the activity of the enzyme by binding to the Zn2+ ion. pHEMA-based CLs prepared with 4-vinylimidazole (4-VI) that resembles histidine and a source of Zn2+ ions (FiguRe 10) showed network/water partition coef-ficient (K

N/W) values three-times greater than

those prepared with 1-vinylimidazole, and also provided a better control of drug release [122]. Sim-ilar improvements were observed when pHEMA was replaced by other the backbone monomers (NVP/DMA) maintaining the biomimetic func-tional groups [124]. Nonbioinspired imprinted networks have also been shown to be useful for loading and release of dorzolamide [125].

Conclusion & future perspectiveThe idea of using CLs as delivery systems of ophthalmic therapeutic agents was postulated 50 years ago, but these combination products can still be considered to be in their early youth. Since the first attempts to use conventional com-mercial CLs as not only vision correctors but also as platforms for management of ocular patholo-gies, the CLs have revealed a limited affinity for ophthalmic drugs and consequently poor capability for controlling drug release. Notable advances regarding materials to use as compo-nents of CLs have been mainly focused on the improvement of oxygen permeability, visual acu-ity and wearer comfort, paying little attention to a potential application for drug delivery. Never-theless, the efforts of clinicians to screen, among the commercially available ones, the CLs with better performance as drug platforms have gen-erated information of great interest for further advances in the field. In fact, certain commer-cial CLs could behave as interesting bandages, which can be readily loaded just by soaking, for ensuring drug ocular levels up to a few hours, compared with the more limited performance of eye drops and other topically applied drug formulations. Controlled release could be suc-cessfully attained for demulcent macromolecules

Model for biomimetichydrogels

The active site of human carbonic anhydrase II, interacting with acetazolamide

N

N N N

N

N

Zn

HNN

N

S OO

S

O

HN

HO

O

S

N-H

OO

Thr 199

His 119

His 94

His 96

Hydrogels with receptor-like binding sites take up and control the release of carbonic anhydrase inhibitors

N

N N N

N

N

S

N-H

O

O

S

O

O

O

O

OO

N

OO

HEMA

HN

OH

4-VI HO

ZnMA2

OH4-VI

OH

HEAA

HN

Zn

Figure 10. The biomimetic approach followed for the design of acetazolamide-imprinted networks. HEAA: N-(2-hydroxyethyl)acrylamide; HEMA: 2-hydroxyethyl methacrylate. Reproduced with permission from [122] © American Chemical Society.



that have smaller diffusion coefficients than low molecular weight drugs [126–128].

The last decade has witnessed the dawn of novel approaches for development of true drug/CL combination products. Among oth-ers, two approaches have been revealed as the most promising ones since they do not compro-mise the optical features of the CLs: sequential or simultaneous soaking of commercial CLs in drug solution and vitamin E solution, in such a way vitamin E enhances drug/network parti-tion coefficient and regulates the diffusion in/out of the drug and water; and application of the molecular imprinting approach during synthesis of extended-wear CLs. The first approach could benefit from the fact that the synthesis process of the CLs does not need to be necessarily altered, and thus any CL could be loaded with any drug, although the loading involves several steps and premature discharge in the multipurpose solu-tions cannot be completely avoided. The second approach requires a detailed knowledge about the interactions of the target drug with the monomer constituents of the CLs to create efficient drug/CL pairs. Imprinted CLs are so far the ones that can most prolong the release profiles.

It is obvious that many questions still remain to be answered before wider clinical use of drug/CL

combination products, mainly regarding when and how the CLs should be loaded and, if pos-sible, reloaded and, more importantly, what dose and release rate are the most convenient for each pathology/CL wearer. The popularization of the use of CLs (currently ~125 million wearers; [209]) and the already proven clinical benefits that CLs can provide to ocular therapeutics are undoubt-edly strong driving forces for the emergence of the first drug/CL combination products in the market in the very near future.

Financial & competing interests disclosureThe preparation of this article and the work of the authors mentioned herein were supported financially by MICINN (SAF2011–22771), Xunta de Galicia (10CSA203013PR) Spain and FEDER. C González-Chomón acknowledges an FPU grant from MEC, Spain. C Alvarez-Lorenzo and A Concheiro are associated with the design of novel materials for ocular drug delivery, including medicated soft contact lenses, covered by patent numbers WO03/090805 and ES2335958. The authors have no other relevant affiliations or financial involvement with any organization or entity with a financial interest in or financial conflict with the subject matter or materials discussed in the manuscript apart from those disclosed.

No writing assistance was utilized in the production of this manuscript.

Executive summary

The value of contact lenses as ocular drug-delivery platforms

�� Drugs are loaded onto contact lenses (CLs) by soaking in drug solutions.

�� Sustained release to the postlens tear film, slower turnover and minimized loss lead to greater and prolonged drug concentrations in aqueous humor compared with topical drug formulations.

�� Less dose may be required to attain the therapeutic effects, as well as excellent optical acuity and biocompatibility.

Limitations of commercial CLs

�� There is limited loading of hydrophilic, nonionic drugs.

�� Irreversible binding of ionic drugs to pHEMA-based CLs and of hydrophobic drugs to silicone hydrogels can occur.

�� Drug is delivered too rapidly.

Coadjutants for improving drug loading/release performance of conventional CLs

�� Incorporation of the drug as a polymer film during CL synthesis. An optical window is needed due to opacity of the film.

�� Adsorption of an amphiphilic molecule, for example, vitamin E, that favors drug interaction with the CL and delays drug release is a simpler and more efficient method.

Novel approaches for drug/CL combination products

�� Drug-loaded nanocarriers dispersed in the CL matrix can efficiently regulate drug release. Physical stability of the nanocarrier and avoidance of premature leakage have to be ensured.

�� Random copolymerization with ionic monomers or grafting of cyclodextrins to CL components may notably improve drug/network partition coefficient.

�� Drug-template polymerization (molecular imprinting) of the CL monomers results in high-affinity domains with the chemical groups more suitable for specific uptake of the drug of interest, this results in high loading and prolonged release.

Regulatory aspects

�� The primary mode of action will determine the regulatory process as device (Center for Devices and Radiological Health) or as drug (Center for Drug Evaluation and Research).

Key Term

Drug/contact lens combination products: Product comprising two regulated components (drug plus contact lenses) that are combined or mixed in a single entity to be used jointly as vision correctors and an ocular drug-delivery system.



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Soft contact lenses for controlled ocular delivery: 50...

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