US20110217355A1

US20110217355A1 - Contact lens based bioactive agent delivery system

Info

Publication number: US20110217355A1
Application number: US12/675,886
Authority: US
Inventors: Anuj Chauhan
Original assignee: Individual
Current assignee: Individual
Priority date: 2007-09-04
Filing date: 2008-09-04
Publication date: 2011-09-08
Also published as: US8075909B2; WO2009032267A2; WO2009032267A3; US20090060981A1

Abstract

A bioactive agent delivery system comprising an optically transparent contact lens having dispersed therein (1) an ophthalmically bioactive agent capable of diffusion through the contact lens and into the post-lens tear film when placed on the eye and (2) associated with the bioactive agent, an ophthalmically compatible polymeric surfactant in an amount sufficient to slow the rate of migration of the bioactive agent through the contact lens.

Description

RELATED APPLICATION

This application claims the benefit of U.S. application Ser. No. 11/896,608 filed Sep. 4, 2007, which is herein incorporated by reference in its entirety.

FIELD OF THE INVENTION

The present invention relates to methods and systems for the delivery of ophthalmic drugs and other bioactive agents to the eye.

BACKGROUND OF THE INVENTION

Providing and maintaining adequate concentrations of bioactive agents, such as drugs, for example, in the pre-corneal tear film for extended periods of time is one of the major problems plaguing methods and systems for ocular drug delivery. When they are applied as eye drops, most drugs penetrate poorly through the cornea. Drainage of instilled drug with the tear fluid, and absorption through the conjunctiva leads to a short duration of action. The additional pre-corneal factors that contribute to the poor ocular bio-availability of many drugs when instilled in the eye as drops are tear turnover and drug binding to tear fluid proteins. In addition to the above factors, the rate of corneal uptake is high at early times, but it declines rapidly. This may lead to a transient period of overdose and associated risk of side effects followed by an extended period of sub-therapeutic levels before the administration of next dose. All the above factors indicate the need for an ocular drug delivery system that will be as convenient as a drop but will serve as a controlled release vehicle [Nagarsenker, M. S., Londhe, V. Y., Nadkarni, G. D., “Preparation and evaluation of liposomal formulations of tropicamide for ocular delivery”, Int. J. of Pharm., 1990, 190: 63-71].
Topical delivery via eye drops that accounts for about 90% of all ophthalmic formulations is very inefficient and in some instances leads to serious side effects [Lang, J. C., “Ocular drug delivery conventional ocular formulations”. Adv. Drug Delivery, 1995, 16: 39-43]. Only about 5% of the drug applied as drops penetrate through the cornea and reaches the ocular tissue, while the rest is lost due to tear drainage [Bourlais, C. L., Acar, L., Zia H., Sado, P. A., Needham, T., Leverge, R., “Ophthalmic drug delivery systems”, Progress in retinal and eye research, 1998, 17, 1: 33-58]. The drug mixes with the fluid present in the tear film upon instillation and has a short residence time of about 2-5 minutes in the film. About 5% of the drug gets absorbed and the remaining flows through the upper and the lower canaliculi into the lacrimal sac. The drug containing tear fluid is carried from the lacrimal sac into the nasolacrimal duct, and eventually, the drug gets absorbed into the bloodstream. This absorption leads to drug wastage and more importantly, the presence of certain drugs in the bloodstream leads to undesirable side effects. For example, beta-blockers such as Timolol that is used in the treatment of wide-angle glaucoma have a deleterious effect on heart [TIMPOTIC® prescribing information, supplied by MERCK]. Furthermore, application of ophthalmic drugs as drops results in a rapid variation in drug delivery rates to the cornea that limits the efficacy of therapeutic systems [Segal, M., “Patches, pumps and timed release”, FDA Consumer magazine, October 1991]. Thus, there is a need for new ophthalmic drug delivery systems that increase the residence time of the drug in the eye, thereby reducing wastage and eliminating side effects.
There have been a number of attempts in the past to use contact lenses for ophthalmic drug delivery; however, all of these focused on soaking the lens in drug solution followed by insertion into the eye. In one of the studies, the authors focused on soaking the lens in eye-drop solutions for one hour followed by lens insertion in the eye [Hehl, E. M., Beck, R., Luthard K., Guthoff R., “Improved penetration of aminoglycosides and fluoroquinolones into the aqueous humour of patients by means of Acuvue contact lenses”, European Journal of Clinical Pharmacology, 1999, 55 (4): 317-323]. Five different drugs were studied and it was concluded that the amount of drug released by the lenses are lower or of the same order of magnitude as the drug released by eye drops. This happened perhaps because the maximum drug concentration obtained in the lens matrix is limited to the equilibrium concentration. In another study researchers developed a contact lens with a hollow cavity by bonding together two separate pieces of lens material [Nakada, K., Sugiyama, A., “Process for producing controlled drug-release contact lens, and controlled drug-release contact lens thereby produced”; U.S. Pat. No. 6,027,745, May 29, 1998]. The compound lens is soaked in the drug solution. The lens imbibes the drug solution and slowly releases it upon insertion in the eye. The compound lens suffers from the same limitations as the drug-soaked lens because the concentration of the drug in the cavity is the same as the concentration of the drug in the drops and thus such a lens can supply the drug for a limited amount of time. Furthermore, the presence of two separate sheets of lens material leads to smaller oxygen and carbon dioxide permeabilities that can cause an edema in the corneal tissue. The other studies and patents listed below suffer from the same limitations because they are also based on soaking of contact lenses or similar devices in drug-solutions followed by insertion into the eye [Hillman, J. S., “Management of acute glaucoma with Pilocarpine-soaked hydrophilic lens” Brit. J. Ophthal. 58 (1974) p. 674-679, Ramer, R. and Gasset, A., “Ocular Penetration of Pilocarpine:” Ann. Ophthalmol. 6, (1974) p. 1325-1327, Montague, R. and Wakins, R., “Pilocarpine dispensation for the soft hydrophilic contact lens” Brit .J. Ophthal. 59, (1975) p. 455-458, Hillman, J., Masters, J. and Broad, A. “Pilocarpine delivery by hydrophilic lens in the management of acute glaucoma” Trans. Ophthal. Soc. U. K. (1975) p. 79-84, Giambattista, B., Virno, M., Pecori-Giraldi, Pellegrino, N. and Motolese, E. “Possibility of Isoproterenol Therapy with Soft Contact Lenses: Ocular Hypotension Without Systemic Effects” Ann. Ophthalmol 8 (1976) p. 819-829, Marmion, V. J. and Yardakul, S. “Pilocarpine administration by contact lens” Trans. Ophthal. Soc. U. K. 97, (1977) p. 162-3, U.S. Pat. No. 6,410,045, Drug delivery system for antiglaucomatous medication, Schultz; Clyde Lewis, Mint; Janet M; U.S. Pat. No. 4,484,922, Occular device, Rosenwald; Peter L., U.S. Pat. No. 5,723,131, Contact lens containing a leachable absorbed material, Schultz; Clyde L. Nunez; Ivan M.; Silor; David L.; Neil; Michele L.].
A number of researchers have focused on developing ‘imprinted’ contact lenses [Hiratani H, Alvarez-Lorenzo C-“The nature of backbone monomers determines the performance of imprinted soft contact lenses as timolol drug delivery systems” Biomaterials 25, 1105-1113, 2004; Hiratani H, Fujiwara A, Tamiya Y, Mizutani Y, Alvarez-Lorenzo C-“Ocular release of timolol from molecularly imprinted soft contact lenses” Biomaterials 26, 1293-1298, 2005; Hiratani H, Mizutani Y, Alvarez-Lorenzo C— “Controlling drug release from imprinted hydrogels by modifying the characteristics of the imprinted cavities” Macromol Biosci 5,728-733, 2005: Alverez-Lorenzo C, Hiratani H, Gomez-Amoza J L, Martinez-Pacheco R, Souto C, Concheiro A—“Soft contact lenses capable of sustained delivery of timolol” J Pharm Sci 91, 2182-2192, 2002; Hiratani H, Alvarez-Lorenzo C-“Timolol uptake and release by imprinted soft contact lenses made of N,N-diethylacrylamide and methacrylic acid” J Control Release 83,223-230, 2002]. The imprinting leads to an increase in the partition coefficients and slower release of drugs, but the increase is not very substantial, and these lenses typically have an initial burst release.
A number of researchers have trapped proteins, cells and drugs in hydrogel matrices by polymerizing the monomers that comprise the hydrogel, in presence of the encapsulated species [Elisseeff, J., McIntosh, W., Anseth, K., Riley, S., Ragan, P., Langer, R., “Photoencapsulation of chondrocytes in poly(ethylene oxide)-based semi-interpenetrating networks”, Journal of Biomedical Materials Research, 2000, 51 (2): 164-171; Ward, J. H., Peppas, N. A., “Preparation of controlled release systems by free-radical UV polymerizations in the presence of a drug”, Journal of Controlled Release, 2001, 71 (2): 183-192; Scott, R. A., Peppas, N. A., “Highly crosslinked, PEG-containing copolymers for sustained solute delivery”, Biomaterials, 1999, 20 (15): 1371-1380; Podual, K., Doyle F. J., Peppas N. A., “Preparation and dynamic response of cationic copolymer hydrogels containing glucose oxidase”, Polymer, 2000, 41 (11): 3975-3983; Colombo, P., Bettini, R., Peppas, N. A., “Observation of swelling process and diffusion front position during swelling in hydroxypropyl methyl cellulose (HPMC) matrices containing a soluble drug”, Journal of Controlled Release, 1999, 61 (1,2): 83-91; Ende, M. T. A., Peppas, N. A., “Transport of ionizable drugs and proteins in crosslinked poly(acrylic acid) and poly(acrylic acid-co-2-hydroxyethyl methacrylate) hydrogels. 2. Diffusion and release studies”, Journal of Controlled Release, 1997, 48 (1): 47-56; U.S. Pat. No. 4,668,506]. Although direct entrapment of drug could lead to higher loading, in a majority of cases, the loaded drug is released rapidly from contact lenses.
Recently, it has been suggested to disperse in contact lenses nanoparticles of ophthalmic bioactive agents nanoencapsulated in a material from which the ophthalmic drug is capable of diffusion into and migration through the contact lens and into the post-lens tear film when the contact lens is placed on the eye [Gulsen D, Chauhan A—“Dispersion of microemulsion drops in HEMA hydrogel: a potential ophthalmic drug delivery vehicle”. Int J Pharm 292, 95-117, 2005., Gulsen D, Chauhan A—“Ophthalmic drug delivery through contact lenses”. Invest Ophth V is Sci 45, 2342-2347, 2004.] Also Graziacascone et al. discloses a study on encapsulating lipophilic drugs inside nanoparticles, and entrapping the particles in hydrogels. [Graziacascone, M., Zhu, Z., Borselli, F., Lazzeri, L., “Poly(vinyl alcohol) hydrogels as hydrophilic matrices for the release of lipophilic drugs loaded in PLGA nanoparticles”, Journal of Material Science: Materials in Medicine, 2002, 13: 29-32]. They used PVA hydrogels as hydrophilic matrices for the release of lipophilic drugs loaded in PLGA particles. These systems are potentially useful but display the shortcoming of burst release due to the presence of the drug outside the particles. Also, these systems required formulations of nanoparticles followed by addition of these nanoparticles to the polymerizing medium. The solution is then required to be polymerized to trap the nanoparticles in the gel. Thus this is a multistep procedure for making nanoparticle-laden contact lenses, which is not optimal. Furthermore, there is a possibility that some nanoparticles may degrade during the gel polymerization step.
The present invention seeks to overcome these obstacles utilizing surfactants to slow down the release rates of drugs from contact lenses. The use of surfactants to retard drug release rates from polymeric gels has been reported but none of these focused on creating surfactant-laden contact lenses [Rodriguez R, Alvarez-Lorenzo C, Concheiro A, “Interactions of ibuprofen with cationic polysaccharides in aqueous dispersions and hydrogels rheological and diffusional implications”, European Journal of Pharmaceutical Sciences 20 (4-5): 429-438, 2003, Rodriguez R, Alvarez-Lorenzo C, Concheiro A, “Influence of cationic cellulose structure on its interactions with sodium dodecylsulfate: implications on the properties of the aqueous dispersions and hydrogels”, European Journal of Pharmaceutics and Biopharmaceutics 56 (1): 133-142 2003, Barreiro-Iglesias R, Alvarez-Lorenzo C, Concheiro A, “Thermal and FTIR characterization of films obtained from carbopol/surfactant aqueous solutions”, Journal of Thermal Analysis and calorimetry, 68 (2): 479-488 2002, Barreiro-Iglesias R, Alvarez-Lorenzo C, Concheiro A, “Incorporation of small quantities of surfactants as a way to improve the rheological and diffusional behavior of carbopol gels”, Journal of Controlled Release, 77 (1-2): 59-75, 2001, Paulsson M, Edsman K, “Controlled drug release from gels using lipophilic interactions of charged substances with surfactants and polymers”, Journal of Colloid and Interface Science, 248 (1): 194-200, 2002, Paulsson M, Edsman K, “Controlled drug release from gels using surfactant aggregates. II. Vesicles formed from mixtures of amphiphilic drugs and oppositely charged surfactants”, Pharmaceutical Research, 18 (11): 1586-1592, 2001, Paulsson M, Edsman K, “Controlled drug release from gels using surfactant aggregates: I. Effect of lipophilic interactions for a series of uncharged substances”, Journal of Pharmaceutical Sciences, 90 (9): 1216-1225, 2001, Yan H, Tsujii K, Potential application of poly(N-isopropylacrylamide) gel containing polymeric micelles to drug delivery systems, Colloids and Surfaces B-Biointerfaces, 46 (3): 142-146, 2005]. Also most previous efforts at using surfactants to retard the drug transport rates focused on systems dilute in polymer with typical loadings less than 10%. Polymer loading is an important parameter in such systems because a higher polymer loading provides more adsorption sites for the drug that is trapped in the gel. Accordingly, micellar systems need to have higher drug partitioning to compete with the higher adsorption sites on the polymer. The systems disclosed here contain about 40% polymer in hydrated state of the gel, and these systems to our knowledge have not been disclosed previously for attenuating drug release rates.
It is therefore an object of the present invention to provide a novel bioactive agent delivery system, particularly adapted for delivering the agent to the eye.

SUMMARY OF THE INVENTION

One embodiment of the invention relates to a bioactive agent delivery system comprising a substantially optically transparent contact lens having dispersed therein (1) at least one ophthalmically bioactive agent, said agent being capable of diffusion through said contact lens and into the post-lens tear film when said contact lens is placed on the eye and (2) associated with said bioactive agent, at least one ophthalmically compatible surfactant, said surfactant being present in an amount sufficient to attenuate the rate of migration of said bioactive agent through said contact lens.
A second embodiment of the invention is a method of administering a bioactive agent to a patient in need thereof comprising placing on the eye the above described drug delivery system.
Third and fourth embodiments of the invention concern a kit and its use for the storage and delivery of ophthalmic drugs to the eye, the kit comprising:
a) a first component containing at least one of the above described drug delivery systems, and
b) a second component containing at least one storage container for the first component, the storage container additionally containing a material that substantially prevents the diffusion and migration of the ophthalmic drug during storage.
A fifth embodiment of the invention relates to a method of manufacturing a bioactive agent delivery system of claim 1 comprising providing a monomer mixture comprising a lens-forming monomer, the surfactant and the bioactive agent and polymerizing said monomer mixture.
Sixth and seventh embodiments of the invention concern articles of manufacture comprising packaging material and the above described drug delivery system or the above-described kit contained within the packaging material, wherein the packaging material comprises a label which indicates that the drug delivery system and kit can be used for ameliorating symptoms associated with pathologic conditions of the eye.
These and other objects are achieved in the present invention.
There has thus been outlined, rather broadly, the more important features of the invention in order that the detailed description thereof that follows may be better understood, and in order that the present contribution to the art may be better appreciated. There are, of course, additional features of the invention that will be described further hereinafter.
In this respect, before explaining at least one embodiment of the invention in detail, it is to be understood that the invention is not limited in its application to the details of construction and to the arrangements of the components set forth in the following description or illustrated in the drawings. The invention is capable of other embodiments and of being practiced and carried out in various ways. Also, it is to be understood that the phraseology and terminology employed herein are for the purpose of description and should not be regarded as limiting.
As such, those skilled in the art will appreciate that the conception upon which this disclosure is based may readily be utilized as a basis for the designing of other structures, methods and systems for carrying out the several purposes of the present invention. It is important, therefore, that equivalent constructions insofar as they do not depart from the spirit and scope of the present invention, are included in the present invention.
For a better understanding of the invention, its operating advantages and the specific objects attained by its uses, reference should be had to the accompanying drawings and descriptive matter which illustrate preferred embodiments of the invention.

BRIEF DESCRIPTION OF THE DRAWINGS AND THE FIGURES

FIGS. 1-7 depict drug release rates in several embodiments of the invention.

FIGS. 8( a-c) illustrate percentage release of surfactant during water change experiments. The gel thicknesses in μm and the percentage of surfactant loaded in the gel are indicated in the legend.

FIG. 8( a) illustrates Brij 97 surfactant systems.

FIG. 8( b) illustrates Brij 78 surfactant systems.

FIG. 8( c) illustrates Brij 700 surfactant systems.

FIGS. 9( a-c) illustrate cumulative percentage release of surfactant from the hydrogels after rescaling the time. ⊖ represents √t/CP/h2 where t is time in seconds, CP is surfactant concentration in M, and h is half thickness of the gel in μm.

FIG. 9( a) illustrates Brij 97 surfactant systems.

FIG. 9( b) illustrates Brij 78 surfactant systems.

FIG. 9( c) illustrates Brij 700 surfactant systems.

FIGS. 10( a-d) illustrate effect of thickness on percentage release of drug during equilibrium experiments. τ represents √t/i2 where t is release time in hours and i=1 for thin gels and i=2 for thick gels. Gels contain 50 μg of drug.

FIG. 10( a) illustrates pure p-HEMA systems.

FIG. 10( b) illustrates Brij 97 laden gels.

FIG. 10( c) illustrates Brij 78 laden gels.

FIG. 10( d) illustrates Brij 700 laden gels.

FIG. 11 illustrates a schematic of the microstructure of the surfactant-laden gels.

FIGS. 12( a-d) illustrate effect of surfactant loading on cumulative drug release from surfactant-laden gels in PBS change experiments. Gels are 200 μm thick and contain 50 μg of drug.

FIG. 12( a) illustrates Brij 97 surfactant system.

FIG. 12( b) illustrates Brij 97 surfactant system.

FIG. 12( c) illustrates Brij 700 surfactant system

FIGS. 13( a-d) illustrate cumulative percentage release of drug from 200 μm thick surfactant-laden gels in PBS change experiments after resealing the time. γ represents √t/h2 where t is time in seconds and h is half thickness of the gel in μm.

FIG. 13( a) illustrates Brij 97 surfactant systems.

FIG. 13( b) illustrates Brij 78 surfactant systems.

FIG. 13( c) illustrates Brij 700 surfactant systems.

FIG. 13( d) illustrates Brij 98 surfactant systems.

FIG. 14 illustrates lysozyme uptake in surfactant laden and pure p-HEMA hydrogels.

FIGS. 15( a-o) illustrate cryo-SEM image for 200 μm thick gels.

FIGS. 15( a)-(c) illustrate pure p-HEMA gels

FIGS. 15( d)-(f) illustrate Brij 98 laden gels.

FIGS. 15 (g)-(i) illustrate Brij 97 laden gels.

FIGS. 15 (j)-(l) illustrate Brij 78 laden gels.

FIGS. 15 (m)-(o) illustrate Brij 700 laden gels.

FIGS. 16 (a-b) illustrates frequency dependence of moduli for 800 μm thick surfactant laden and pure p-HEMA gels.

FIG. 16( a) illustrates storage modulus.

FIG. 16( b) illustrates loss modulus.

FIG. 17 illustrates Standard Linear Solid Model used for fitting the viscoelasticity data of the surfactant-laden gels.

FIG. 18 illustrates effect of thickness on the storage and loss moduli of pure p-HEMA gels. The gel thicknesses in μm are indicated in the legends.

DETAILED DESCRIPTION OF THE INVENTION

The present invention is predicated on the discovery that contact lenses, preferably, soft contact lenses can function as new vehicles for ophthalmic drug delivery to reduce drug loss, eliminate systemic side effects, and improve drug efficacy.
Specifically, the invention relates to the discovery that the rate of migration of bioactive agents, capable of diffusion through contact lenses and into the post-lens tear film when the contact lens is placed on the eye, is attenuated when the bioactive agent is associated with at least one ophthalmically compatible surfactant.
The contact lenses of the present invention are formed from reaction mixtures which comprise the reactive components, catalyst, other desired components, and optionally a solvent. The reaction mixtures may be cured using conventionally known conditions well known to those skilled in the art.
Hydrophilic components are those which when mixed, at 25° C. in a 1:1 ratio by volume with neutral, buffered water (pH about 7.0) form a homogenous solution. Any of the hydrophilic monomers known to be useful to make hydrogels may be used.
Exemplary of suitable hydrophilic monomers are DMA, HEMA, glycerol methacrylate, 2-hydroxyethyl methacrylamide, NVP, N-vinyl-N-methyl acrylamide, N-methyl-N-vinylacetamide, polyethyleneglycol monomethacrylate, methacrylic acid and acrylic acid, polymers or copolymers of any of the foregoing, mixtures thereof and the like.
The reaction mixtures may also comprise at least one hydrophobic component. Hydrophobic components are those which when mixed, at 25° C. in a 1:1 ratio by volume with neutral, buffered water (pH about 7.0) form an immiscible mixture.
Examples of suitable hydrophobic components include silicone containing components, fluorine containing components, components comprising aliphatic hydrocarbon groups having at least 3 carbons, mixtures thereof and the like.
The term component includes monomers, macromers and prepolymers. “Monomer” refers to lower molecular weight compounds that can be polymerized to higher molecular weight compounds, polymers, macromers, or prepolymers. The term “macromer” as used herein refers to a high molecular weight polymerizable compound. Prepolymers are partially polymerized monomers or monomers which are capable of further polymerization.
The invention is exemplified herein using soft hydrogel lenses that are made of poly 2-hydroxyethyl methacrylate p-(HEMA). However, it will be understood by those skilled in the art that the range of materials that may be employed as vehicles in the present invention is limited only by the selection of materials that may be employed in the manufacture of contact lenses and the nature of the particular ophthalmic drug to be incorporated therein. The term, “optically transparent” as used herein is intended to refer to a degree of transparency equivalent to that of p-HEMA or other material employed as a contact lens. The p-HEMA hydrogel matrix may be synthesized by any convenient method, e.g., bulk or solution free radical polymerization of HEMA monomers in presence of a cross linker such as ethylene glycol-di-methacrylate (EGDMA) [Mandell, R. B., “Contact Lens Practice: Hard and Flexible Lenses”, 2nd ed., Charles C. Thomas, Springfield, vol. 3, 1974].
Addition of the bioactive agent and surfactant to the polymerizing medium followed by polymerization could result in the formation of self assembled surfactant aggreagtes that contain the bioactive agent. If contact lenses made of this material are placed on the eye, the drug molecules will diffuse from the surfactant aggregates, travel through the lens matrix, and enter the post-lens tear film (POLTF), i.e., the thin tear film trapped in between the cornea and the lens. In the presence of the lens, drug molecules will have a much longer residence time in the post-lens tear film, compared to about 2-5 minutes in the case of topical application as drops [Bourlais, C. L., Acar, L., Zia H., Sado, P. A., Needham, T., Leverge, R., “Ophthalmic drug delivery systems”, Progress in retinal and eye research, 1998, 17, 1: 33-58; Creech, J. L., Chauhan, A., Radke, C. J., “Dispersive mixing in the posterior tear film under a soft contact lens”, I&EC Research, 2001, 40: 3015-3026; McNamara, N. A., Polse, K. A., Brand, R. D., Graham, A. D., Chan, J. S., McKenney, C. D., “Tear mixing under a soft contact lens: Effects of lens diameter”. Am. J. of Ophth., 1999, 127(6): 659-65]. The longer residence time will result in a higher drug flux through the cornea and reduce the drug inflow into the nasolacrimal sac, thus reducing drug absorption into the blood stream. In addition, due to the slow diffusion of the drug molecules through the particles, drug-laden contact lenses can provide continuous drug release for extended periods of time.
Without wishing to be bound by any theory, the inventors believe that the mechanism of attenuation of migration of the active agent is one of entrapment of the agent in a micelle or inverse micelle structure formed by the surfactant.
Suitable surfactants include any ophthalmically compatible surfactants capable of providing the necessary attenuation in release rates without affecting the optical transparency of the resulting contact lens. The examples presented here utilize nonionic surfactants but it should be understood that cationic, anionic, and zwitterionic surfactants are equally applicable. Also both linear and branched surfactants may be utilized. Exemplary of suitable surfactants are the following.
Block copolymers, which are surface active, are classified by the ratio of the hydrophilic and lipophilic segments in the molecule. A large number of commercial emulsifying agents, such as surfactants, have been assigned a hydrophilic/lipophilic balance (HLB) number. The block copolymer consists of a hydrophilic moiety (water soluble) and a hydrophobic moiety.
The preferred water soluble (hydrophilic) region of the block copolymer consists of polyethylene glycol, polyethylene oxide, polyvinyl alcohol, polyacrylamide, polymethacrylamide, poly(vinylpyrrolidone), and the like. It is most preferred that the hydrophilic moiety is polyethylene glycol, polyacrylamide, polymethacrylamide, poly (vinylpyrrolidone) or polyvinyl alcohol. The most preferred hydrophilic core is polyethylene glycol.
The hydrophobic polymer segment is attached to the hydrophilic polymer by non-hydrolyzable chemical bonds, such as carbon-carbon bonds, by amide linkage, ether linkages, ester linkages, thio linkages, amino linkages, and the like. The preferred hydrophobic polymer segments include linear and branched carbon chains (both saturated and unsaturated), poly propylene oxide, poly hydroxy butyrate, polystyrene, etc. The preferred hydrophobic polymer segments also include poly(α-hydroxycarboxylic acids) which are derived from either glycolide or lactide; poly(ω-hydroxycarboxylic acids) which are derived from either ω-lactone or δ-lactone or ε-lactone; or those derived from a copolymer of such include poly(α-hydroxycarboxylic acids) with such poly(ω-hydroxycarboxylic acids). The hydrophobic polymer segments may have an ethylenically unsaturated polymerizable group at one end which is opposite to the one at which the hydrophobic polymer segment is bonded to the hydrophilic polymer segment. Such a polymerizable group can be introduced from (meth)acrylic acid or vinylbenzyl chloride. Furthermore, such a polymerizable group may be subjected to a polymerization reaction after the formation of the polymer, and is thus brought into a polymerized (crosslinked) state. In such a state, the polymer per se is more stable.
The most preferred polymer is formed from a block copolymer which is composed of both a hydrophilic polymer segment essentially comprising poly (ethyleneglycol) [hereinafter sometimes abbreviated as PEG] and a hydrophobic polymer segment. The phrase “essentially comprising” means that PEG occupies the main portion of the hydrophilic polymer segment, and that some linking group or the like which has essentially no influence on the hydrophilicity of said segment may be contained in some amount in the PEG chain or between hydrophilic and hydrophobic polymer segments. However, it is preferable that the PEG chain consists of PEG alone.
Examples of block copolymers are found in U.S. Pat. No. 5,925,720, to Kataoka, et al., U.S. Pat. No. 5,412,072 to Sakarai, et al., U.S. Pat. No. 5,410,016 to Kataoka, et al., U.S. Pat. No. 5,929,177 to Kataoka, et al., U.S. Pat. No. 5,693,751 to Sakurai, et al., U.S. Pat. No. 5,449,513 to Yokoyama, et al., WO 96/32434, WO 96/33233 and WO 97/0623, the contents of all of which are incorporated by reference. Modifications thereof which are prepared by introducing thereon a suitable functional group (including an ethylenically unsaturated polymerizable group) are also examples of block copolymers from which surfactants of the present invention are preferably prepared.
The contact lens may be formed by the polymerization of any suitable reactive components or mixtures of reactive components known in the art to produce contact lenses and, in which the bioactive agent may be microemulsified by the surfactant and in which the bioactive agent migrates when placed on the eye. The lens material may be hydrophilic or hydrophobic.
The lens-forming or reactive components include monomers, prepolymers and macromere that are polymerizable by free radical polymerization, generally including an activated unsaturated radical, and most preferably an ethylenically unsaturated radical.
Additional additives, which are generally known in the art may also be included. Additives include but are not limited to wetting agents, ultra-violet absorbing compounds, tints, pigments, photochromic compounds, release agents, combinations thereof and the like. The additional compounds may be reactive or non-reactive.
An especially preferred class of materials are hydrogel copolymers. A hydrogel is a crosslinked polymeric system that can absorb and retain water in an equilibrium state. Accordingly, for hydrogels, the monomer mixture will typically include at least one hydrophilic monomer and a crosslinking agent (a crosslinker being defined as a monomer having multiple polymerizable functionalities). Suitable hydrophilic monomers include: unsaturated carboxylic acids, such as methacrylic and acrylic acids; acrylic substituted alcohols, such as 2-hydroxyethylmethacrylate and 2-hydroxyethylacrylate; vinyl lactams, such as N-vinyl pyrrolidone; and acrylamides, such as methacrylamide and N,N-dimethylacrylamide. Typical crosslinking agents include polyvinyl, typically di- or tri-vinyl monomers, such as di- or tri(meth)acrylates of diethyleneglycol, triethyleneglycol, butyleneglycol and hexane-1,6-diol; divinylbenzene; and others known in the art.
Another class of lens-forming monomers are those that form silicone hydrogel copolymers. Such systems include, in addition to a hydrophilic monomer, a silicone-containing monomer. Such bulky monomers specifically include methacryloxypropyl tris(trimethylsiloxy)silane, pentamethyldisiloxanyl methylmethacrylate, methyldi(trimethylsiloxy)methacryloxymethyl silane, 3-[tris(trimethylsiloxy)silyl]propyl vinyl carbamate, and 3-[tris(trimethylsiloxy)silyl]propyl vinyl carbonate, monomethacryloxypropyl terminated mono-n-butyl terminated polydimethylsiloxane, mono-(3-methacryloxy-2-hydroxypropyloxy)propyl terminated, mono-butyl terminated polydimethylsiloxane, bis-3-methacryloxy-2-hydroxypropyloxypropyl polydimethylsiloxanes, 3-methacryloxy-2-hydroxypropyloxy)propylbis (trimethylsiloxy)methylsilane, mixtures thereof and the like.
Another suitable class are multifunctional ethylenically “end-capped” siloxane-containing monomers, especially difunctional monomers. Other silicone-containing monomers include the silicone-containing monomers described in U.S. Pat. Nos. 5,034,461, 5,610,252 and 5,496,871, the disclosures of which are incorporated herein by reference. Many other silicone-containing monomers are well-known in the art.
The invention is illustrated by the following non-limiting examples wherein the formulations are based on adding surfactants to polymerizing mixtures. In the first set of Examples the surfactants explored here are linear ethoxylated surfactants (commonly referred as Brij) containing the same alkyl chain length (CIS) and increasing numbers of ethoxylate (EO) units (10, 20, and 100). Here, are reported formulations and fabrication processes for surfactant-laden HEMA based soft contact lenses, and several examples for the loading of ophthalmic drugs such as cyclosporine, and the release of the drugs under conditions that simulate ocular conditions. The results show that the Brij surfactants are very effective at providing extended release of cyclosporine for a period of about 20-25 days. Amongst the three types of Brij surfactants explored, Brij 78 provides the best release behavior. The surfactant-laden gels were subjected to the same processing conditions as contact lenses such as extraction, autoclaving, storage, and these steps did not impact the release rates from the gels. Thus, Brij surfactant laden gels are excellent candidates for the delivery of cyclosporine. These systems may also be adapted to deliver other ophthalmic drugs.
In the second set of Examples the microstructure of the gel is investigated with particular focus on the micellar-aggregates, and the mechanisms that impact the partitioning of the drugs in the aggregates. Also investigated is the effect of surfactant loading on gel physical properties relevant to contact lenses such as transparency, modulus, protein binding, wettability, and water content. These experimental results provide useful data and insight towards delivering CyA to eyes through contact lenses, and also in designing suitable contact lenses for delivering other ophthalmic drugs. Additionally, disclosed are fundamental issues related to drug and surfactant transport in the gels including a first disclosure of the microstructure and physical characterization of the surfactant-laden hydrogels with a large polymer fraction as large as 40% and polymer mesh sizes as small as 2 nm.

EXAMPLE SET ONE

Example 1

Synthesis of gels containing Brij 97

To prepare Brij 97 laden gels with 1.5% surfactant loading in dry state, 0.2 g of surfactant was dissolved in 10 ml of DI water, and then stirred at about 600 rpm at room temperature till the surfactant completely dissolved in the water phase. Separately, 3.5 mg of CyA was dissolved in 2.7 ml of HEMA monomer and stirred at 600 rpm for a period of 5 hours. Next, 15 ml of the crosslinker (EGDMA) and 2 ml of surfactant solution were added to the drug containing HEMA solution. The solution was then degassed by bubbling nitrogen for 10 minutes. Next, 6 mg of the initiator (TPO) was added and the solution was stirred at 300 rpm for 10 minutes to ensure complete dissolution of the initiator. The mixture was then poured in between two glass plates that were separated from each other by 200 (for thick gels) or 100 mm (for thin gels) thick sheet. The polymerization reaction was performed under UV light for 40 minutes. To prepare gels with higher surfactant loading, the amount of surfactant added to the 10 ml DI water was increased. Specifically, 0.2, 0.6, 1.5 g of Brij 97 was added to fabricate gels with 1.5%, 3.5%, 8% surfactant loading in dry gel, respectively. To synthesize HEMA gels without surfactants, 2 ml of the surfactant solution was replaced by 2 ml DI water, and the drug was directly added to the mixture of HEMA, EGDMA and DI water. To synthesize HEMA gels loaded with other surfactants, the same procedure as described above was followed except that Brij 97 was replaced by an equal amount of the desired surfactant.

Example 2

Drug Release Studies

The gels prepared by the procedure described above were cut into pieces that were about 40 mg in weight. In some cases these gels were soaked in a large volume of volume/PBS to extract the unreacted monomer. This step is referred to as the initial extraction. The drug release studies reported below were conducted with a 40 mg gel which was soaked in 3.5 ml PBS and the PBS was replaced every 24 hours. An extraction step was not performed in these experiments. For comparison, pure HEMA gels were loaded with the same amount of drug as the surfactant laden gels, and drug release studies was also performed from these gels. The release of cyclosporine from pure HEMA gels and that from the Brij 97 laden gels with 8% surfactant loading (based on weight of surfactant in dry gel) is shown in FIG. 1 [Comparison of release from Brij 97 surfactant laden gels and HEMA gels. The gels were 200 mm thick in dry state and the amount of drug in the gels was 50 mg. Data is represented as mean, and the error bars represent the standard deviation.
The cyclosporin release from HEMA gels last only about 6-7 days but the surfactant-laden gels release drug for about 25 days. This clearly demonstrates a significant reduction in delivery rate and an increase in the duration of release on addition of surfactant to the gels. It is speculated that the surfactant molecules may be forming aggregates such as micelles, and since cyclosporine is a hydrophobic molecule, a larger fraction of the entrapped drug may be present in the hydrophobic domains of the surfactant aggregates. Thus, the drug concentration in the aqueous phase in the gel may be much smaller for the surfactant containing gels compared to the pure HEMA gels, and the lower free drug concentration may be causing the slower release. Additionally, the drug may face a barrier to diffuse out from the hydrophobic domains, and this may cause an additional reduction in the drug release rates.

Example 3

Dependence of the Release Rates on the Surfactant Loading

As stated above, the results shown in FIG. 1 are for a system that had 8% surfactant loading in the dry gel. To investigate the effect of the surfactant loading on the drug release rates, it was decided to prepare gels with two different surfactant loadings. Table 1 shows the dry gel weight percentages of surfactant and cyclosporine for the two gels with different surfactant loadings.
Drug release experiments were performed on these two gels with protocols described above, and the results are compared below in FIG. 2 [Effect of surfactant concentration on drug release profiles from Brij 97 laden gels. The gels were 200 lam thick in dry state and the amount of drug in the gels was 48 mg].
As shown in the figure, the release rates depend strongly on the surfactant loading. The gel with 8.4% surfactant loading releases CyA for about 500 hours but the gel with 1.4% surfactant loading releases the drug for about 300 hours, which is only marginally longer than the duration of release for HEMA gels. The reductions in release rates due to an increase in surfactant concentration may be attributed to an increase in the number and/or size of surfactant aggregates that may be forming in the gel.

Example 4

Effect of Gel Thickness on Drug Release Profiles

The results presented above were obtained with 200 μm thick gels. Typical contact lenses are about 100 μm thick, and so it was decided to explore the effect of gel thickness on release profiles. To explore this issue, two different sets of drug containing surfactant-laden gels were synthesized. One set of these gels were about 200 μm thick and the others were about 100 μm thick. The surfactant loading for both of these gels was 8% w/w for dry gel. The drug was loaded into the gels by directly dissolving it into the HEMA monomer. It is noted that the weights of both the thick and then thin gels were about same because the cross sectional area of the thin gel was double that of the thick gel. As shown in FIG. 3 [Effect of thickness on drug release profiles for Brij-97 laden gels. The thick and the thin gels were 200 mm and 100 mm thick in dry state, respectively, and the amount of drug in the gels was 47 mg. These gels had 8% surfactant (w/dry gel w)], the release rates are faster for the thin gels, but the thin gels also exhibit extended release.

Example 5

Effect of processing conditions on drug release: In order to evaluate the suitability of the Brij 97 surfactant-laden gels as contact lenses, it was decided to fabricate gels with the same thickness as contact lenses, and take these gels through processing conditions very similar to those used for typical contact lenses. The results of these studies are shown below.
(1) Synthesis: The synthesis procedures were identical to those described earlier for preparing Brij 97 laden gels. Gels were prepared with three different surfactant loadings (1.5, 3.5 and 8.3%). All the gels used in these studies were about 100 μm thick and did not contain any drug. The drug was loaded later by soaking the gels in aqueous drug solutions.
(2) Extraction: The unreacted monomer was extracted from the gels by soaking gels that weighed about 40 mg in 10 ml of water at 50° C. The DI water was replaced every 5 minutes for 5 times. So the total duration for the extraction step was 25 minutes.
(3) Drug Loading: After extraction, each gel was soaked in 4 ml of cyclosporine solution in DI water at a concentration of 12 μg/ml for a period of 12 days. At the end of the loading phase, the concentration in the solution was measured. The drug uptake by the gel was then determined by calculating the difference between the initial and the final drug amounts in the solution.
The results for the drug loaded into two sets of controls (pure HEMA gels) and two sets of 1.5%, 3.5% and 8% surfactant containing gels are shown in Table 1.

TABLE 1

Drug uptake by Brij 97-laden gels during soaking

	Drug in solution	Drug remaining in the	Amount of drug inside
Sample	initially (mg)	solution after 12 days (mg)	the gel system (mg)

PureHEMA1	48	31.8	16.2
PureHEMA2	48	27.1	20.9
8% surfactant1	48	17.4	30.6
8% surfactant2	48	19	29
3.5% surfactant1	48	22.9	25.1
3.5% surfactant2	48	27.1	20.9
1.5% surfactant1	48	23.6	24.4
1.5% surfactant2	48	22.7	25.3

(4) Autoclaving: After drug loading, each gel was soaked in 1.5 ml of DI water and autoclaved for 15 min at 121° C.
(5) Shelf storage: After autoclaving, the samples were stored at room temperature for a period of 10 days. After the 10 day period, the concentration in the aqueous phase was measured to determine the amount of drug that was released from the gel during the autoclaving and shelf storage. By subtracting this amount from the amount of drug taken up by the gel, the remaining amount of drug left in the gel was determined. The results for the drug retained by the gels after the storage are shown in Table 2.

TABLE 2

Summary of drug release studies from Brij 97-laden gels

Amount of drug inside the gel	Amount of drug released	Amount of drug remaining	Amount of drug released during
system before autoclaving (mg)	during shelf storage (mg)	inside the gel (mg)	drug release experiments (mg)

PureHEMA1	16.2	6.2	10	6.8
PureHEMA2	20.9	6.2	14.7	5.6
8%	30.6	5.2	25.4	22.1
surfactant1
8%	29	5.7	23.3	22.1
surfactant2
3.5%	25.1	4.5	20.6	19.2
surfactant1
3.5%	20.9	4.4	16.5	20.8
surfactant2
1.5%	24.4	6.9	17.5	13.9
surfactant1
1.5%	25.3	7	18.3	15.3
surfactant2

(6) Drug release:
In the final step, each gel was submerged in 3.5 ml of PBS, which was replaced every 24 hours, and the concentration of the drug was measured by HPLC. The elution time of cyclosporine that diffused out of the gels after autoclaving was compared with the control to ensure that the drug did not degrade during the processing steps. The drug release profiles for the cumulative % release as a function of time are plotted in FIG. 4 [Effect of the surfactant loading on drug release profiles for Brij 97 surfactant laden gels. All the gels were about 100 mm in thickness. The amount of drug in each gel and the total cumulative release are noted in Table 3].
Also, total cumulative release from the gels is listed in the last column in Table 2. A comparison of the total cumulative release (5th column in Table 2) with the amount of drug retained by the gel (4th column in Table 2) shows that almost the entire amount of the drug retained by the surfactant-laden gels diffuses out during the drug release experiments. However, for pure HEMA gels there is a significant difference between these values. We believe that the discrepancy may be due to an overestimation of the drug loaded into the gel because of neglect of drug absorption on the surface of the glass vials. This underestimation is more important for HEMA gels because the bulk drug concentration is higher for HEMA gels and so there is a larger adsorption on the glass surface.

Example 6

Effect of Hydrophilic Chain Length

To understand the mechanisms involved in transport of drug in the surfactant laden gels, it was decided to fabricate gels with three Brij surfactants with the same hydrophobic group but different lengths of the hydrophilic (EO) group.
The procedures for preparing these gels and performing the drug release experiments are the same as those detailed above. It is noted that drug was loaded into the gels by dissolving it in the HEMA phase and that the initial monomer extraction was not conducted for these gels. The drug release profiles for 100 μm thick gels with 8% surfactant loading are shown in FIG. 5 [Effect of the length of the hydrophilic group (EO) on drug release profiles for Brij surfactant laden gels. All the gels were about 100 mm in thickness and the amount of drug in the gels was 48 mg] for the three types of surfactants. Results show that the drug release from Brij 700 gel s (100 EO units) is similar to that from pure HEMA gels. This may be expected because the fraction of hydrophobic segment is very small for Brij 700 and so there are almost negligible hydrophobic regions in Brij 700 gels. Interestingly, even though Brij 78 (20 EO units) has less hydrophobic fraction compared to Brij 97 (10 EO units), Brij 78-laden gels release drug at a much slower rate than the Brij 97 gels.

Example 7

Effect of Brij 78 Loading on Release Behavior

Since Brij 78 systems seem to be more suitable than Brij 97 systems for extended delivery of cyclosporine; it was decided to explore the dependency of release rates on the surfactant loading for Brij 78 laden gels. Three different surfactant concentrations (1.4%, 4.5% and 8% w/dry gel w) were explored and the release profiles are shown in FIG. 6
Effect of the surfactant loading on drug release profiles for Brij 78 surfactant laden gels. All the gels were about 200 μm in thickness and the amount of drug in the gels was 50 mg]. The results in FIG. 6 show that an increase in the surfactant concentration has a significant effect on the drug release rates. Importantly, even with 1.4% surfactant loading, these systems release drug for about 400 hours, and if the loading is increased to 4.5%, the release duration increases to more than 900 hours.

Example 8

Effect of Thickness on Drug Release from Brij 78 Laden Gels

The results presented above were obtained with 200 μm thick gels. We further deCided to explore the effect of gel thickness on release profiles for Brij 78 laden gels. The drug was loaded into the gels by dissolving it in the HEMA solution before polymerization. Drug release profiles for 100 μm thick and 200 μm thick Brij 78 laden gels are compared in FIG. 7 [Effect of thickness on drug release profiles for Brij-78 laden gels. The thick and the thin gels were 200 mm and 100 mm thick in dry state, respectively, and the amount of drug in the gels was 48 mg. These gels had 8% surfactant (w/dry gel w)]. The thin gels also exhibit extended release of the drug.
Brij surfactant laden systems for ophthalmic drug delivery by contact lenses have been demonstrated. HEMA gels loaded with three different types of Brij surfactants were synthesized. These surfactants have the same hydrophobic segment but different lengths of hydrophilic (EO) segment. Cyclosporine was loaded into these systems by dissolving it in HEMA before polymerization, and in some cases, by soaking the surfactant-laden gels in aqueous cyclosporine solutions. Experiments were conducted to study the effects of surfactant type, concentration and gel thickness on release profiles.
Furthermore, for Brij 97 systems, experiments were performed to simulate the various processes in contact lens manufacturing such as extraction, autoclaving and storage. Results from all these studies show that the Brij 97-laden gels release cyclosporine in PBS for a period of about 20-25 days. The release duration depends weakly on thickness, but strongly on surfactant concentration. These systems are very promising and can be used for extended release of cyclosporine from contact lenses. Brij 78 systems are even more promising because these systems release cyclosporine for longer periods of time compared to Brij 97 systems. In fact with only 4.5% surfactant loading, Brij 78 loaded gels release cyclosporine for about 900 hours. Brij 78 surfactants have other advantages over the Brij 97 surfactants such as a longer chain length, which is expected to reduce the flux of the surfactant from the gel into the eye. Furthermore, Brij 78 surfactants have been used in ocular studies as cornea permeability enhancers, and so these are not expected to cause any toxic response in the eyes. While the examples reported here were conducted with cyclosporine, other drugs could also be dissolved in the HEMA, but the release rates may not be as slow as those for cyclosporine, particularly if the drug molecules are much smaller than cyclosporine. It will be appreciated by those skilled in the art that these systems can be made suitable for other drugs by using mixed surfactants that will pack more tightly. It is also noted that in addition to surfactants, other self assembling molecules such as lipids, and block-co-polymers could be used to create domains that could trap and slowly release hydrophobic drugs. It is further noted that similar ideas could be used to create hydrophilic domains in silicone contact lenses which are hydrophobic in nature. Hydrophilic drugs can then be trapped and slowly released from these hydrophilic domains in the silicone contact lenses.

EXAMPLE SET TWO

Materials

Hydroxy ethyl methacrylate (HEMA) monomer, ethylene glycol dimethacrylate (EGDMA), Dulbecco's phosphate buffered saline (PBS), dexamethasone (DMS), dexamethasone acetate (DMSA), Acetonitrile, lysozyme from chicken egg white, HPLC grade water, Brij 97, Brij 98, Brij 78 and Brij 700 were purchased from Sigma-Aldrich Chemicals (St Louis, Mo.). 2,4,6-trimethylbenzoyl-diphenyl-phosphineoxide (Darocur TPO) was kindly provided by Ciba (Tarrytown, N.Y.). Cyclosporine A (CyA) was purchased from LC Laboratories (Woburg, Mass.). All the chemicals were reagent grade. Acetonitrile was filtered after receiving and all the other chemicals were used without further purification.

Preparation of Surfactant Laden Gels

Surfactant laden gels were prepared by polymerizing the monomer solution containing surfactant and drug mixed in specific ratio. Briefly, 0.25, 0.6, 1.5 g of surfactant was dissolved in 10 ml DI water to make three different surfactant solutions (corresponding to 2%, 4%, 8%, surfactant in dry gel respectively). Separately, 3.5 mg of drug was dissolved in 2.7 ml of HEMA monomer and stirred at 600 rpm for a period of 5 hours. Next 15 μl of the crosslinker and 2 ml of surfactant solution were added to the 2.7 ml of drug loaded monomer. The solution was degassed by bubbling nitrogen gas through it for 10 minutes followed by addition of 6 mg of UV initiator (TPO) and stirring the solution for 10 minutes. The solution was then poured between two glass plates separated by a spacer and the gel was cured by irradiating UVB light (305 nm) for 40 min from an Ultraviolet transilluminiator UVB-10 (Ultra•Lum, Inc.). Four different spacers, 100, 200, 400 and 800 μm in thickness were utilized to synthesize gels of various thicknesses. Control, drug loaded p-HEMA gels without surfactants were prepared by following procedures identical to those described above except that the 2 ml surfactant solution was replaced by 2 ml DI water.

Drug Release Experiments

After polymerization, each gel was removed from the glass mold and was cut into smaller pieces that weighed about 40 mg in dry state. These 40 mg gels were used in all experiments described below. As thickness of the gel was varied, size of the gel piece was adjusted to maintain similar weight for all the gels used in the study. Two sets of experiments were performed for the drug release studies. In the first set of experiments, gel was soaked in 3.5 ml of PBS and measurements were taken until equilibrium was reached for the drug. In the second set, gel was soaked in 3.5 ml PBS and PBS was replaced every 24 hours, mimicking perfect sink conditions for the release experiments. Equilibrium experiments were conducted for all the three drugs explored in this study (CyA, DMS, and DMSA), whereas, PBS replacement experiments were performed for CyA only.

Drug Detection

CyA concentration was measured using a HPLC (Waters, Alliance System) equipped with a C₁₈reverse phase column and UV detector [Kim C, Ryuu S, Park K, Lim S, Hwang S. Preparation and physicochemical characterization of phase inverted water/oil microemulsion containing cyclosporin A. International Journal of Pharmaceutics 1997; 147:131-134]. The mobile phase composition was 70% acetonitrile and 30% DI water, and the column was maintained at 60° C. The flow rate was fixed at 1.2 ml/min and the detection wavelength was set at 210 nm. The retention time for CyA under these conditions was 4.5 minutes, and the calibration curve for area under the peak vs. concentration was linear (R²=0.995).

Surfactant Release Experiments

The rates of surfactant release were measured in 3.5 ml of DI water with water replacement after each measurement to maintain perfect sink conditions. The surfactant concentration in the release medium was determined by measuring surface tension (σ), which was then related to the concentration through a σ(C) calibration curve. The surface tension was measured by using a Wilhelmy plate (sand blasted platinum plate) attached to a Scaime France Microbalance which was further connected to a Stathan Universal transducer (SC001). A detailed description of the process for measuring surfactant concentration by surface tension measurements has been reported earlier [Kapoor Y, Chauhan A. Drug and surfactant transport in Cyclosporine A and Brij 98 laden p-HEMA hydrogels. Journal of Colloid and Interface Sciences 2008; 322:624-633].

Lysozyme Sorption

A lysozyme solution was prepared by adding 40 mg of lysozyme to 40 ml of PBS. The 8% surfactant laden gels (about 40 mg in weight) were soaked in 3.5 ml of lysozyme solution and the amount of lysozyme that was taken up by the hydrogels was monitored by UV detection in the wavelength range 240-340 nm. The concentration of lysozyme was evaluated following a similar protocol as reported above for DMS and DMSA.
Preparation and cryo-SEM of Hydrogels
All samples were soaked in 1×PBS buffer for at least 24 hours. The hydrogel samples were trimmed down to approximately 1 cm×1 cm in size and mounted vertically on the cryo-SEM sample holder with a small amount of Tissue-Tek adhesive (Sakura). The samples were rapidly plunged into liquid nitrogen at a temperature below −190° C. (Gatan, Alto 2500), withdrawn into a vacuum transfer device under the protection of high vacuum, and transferred into the cryo-preparation chamber where the temperature was maintained at −130° C. and the anticontaminator at around −188° C. The hydrogel samples were freeze fractured using the flat edge of a cold knife maintained at −130° C. and sublimated for 5 minutes at −95° C. to etch away surface water and expose the internal structural features. After sublimation, the temperature of the stage was adjusted back to −130° C. and the samples were sputter coated with platinum at 11 mA for 100 seconds. The samples were then transferred into the main chamber of the Field Emission SEM (Hitachi S-4800) via an interlocked airlock and mounted onto a cold stage module (−130° C.) fitted to the SEM stage. Images were acquired at a voltage of 2 kV.

Dynamic Mechanical Analysis

A dynamic mechanical analyzer (DMA Q800, TA instruments) was used to determine the mechanical properties of different surfactant laden systems synthesized above. For this study 400 μm and 800 μm thick gels were utilized to avoid breaking of the gel during the experiment. A hydrated gel was mounted on the clamp and the gel was kept submerged in DI water at room temperature during the experiment. Gel response in form of storage and loss modulus of the gel was determined by applying tensile force in the longitudinal direction while keeping the gel tightly screwed between the clamps by applying a preload force of 0.01 N. To determine the linear viscoelastic range, strain test were first conducted at a frequency of 1 Hz followed by frequency sweep (1-35 Hz) measurements performed for all the samples at 20 micron strain.

Surface Contact Angle Measurements

Surface contact angles were measured for all the surfactant laden systems with 8% loading to investigate the effect of surfactant release on wettability. The contact angles were measured by captive bubble technique with a prop Shape Analyzer (DSA100, KRUSS). This technique was preferred over sessile drop technique to eliminate contact angle change due to sample drying during measurements. A 200 μm thick gel was mounted on a glass slide which was then placed on a water filled cuvette with the lens submerged in water. An air bubble was created by an inverted syringe inside the cuvette, and allowed to detach and rise till it came in contact with the gel, and then the contact angle (θ) was measured. The gels were presoaked in a PBS buffer solution for one day before the experiment.

Transmittance Measurements

The transparency of all the surfactant laden hydrogels was quantified by measuring the transmittance of 100 μm thick hydrated gels at 600 nm using a UV-VIS spectrophotometer (Thermospectronic Genesys 10 UV).

Equilibrium Water Content

The gels of known weight were soaked in 3.5 ml of DI water, and the dynamic weight was measured as a function of time. The excess water from the gel surface was removed before each measurement by dabbing with Kimwipes (Fischer Scientific). The equilibrium water content (EWC) of the surfactant laden gels was calculated by determining amount of water uptake per dry gel weight, i.e.,
$\begin{matrix} % EWC = \frac{W_{WET} - W_{DRY}}{W_{DRY}} . & (1) \end{matrix}$

Statistical Analysis

Linear regression analysis to determine slopes, correlation coefficients and confidence intervals was performed in JMP developed by SAS (Cary, N.C.). The 95% confidence intervals (CI) were utilized to compare release rates.

Results and Discussion

Surfactant Release from the Hydrogels
Table 3 lists the surfactants utilized in this study and their relevant physical properties obtained from literature. The value of ‘fh’ is also listed, this is defined as the fraction of hydrophobic chain length of the surfactant and is calculated by taking the ratio of number of carbons in the hydrophobic tail of the surfactant to the total number of carbons present in the surfactant.

TABLE 3

Physical properties of the surfactants explored in this study.

		Molecular
Name	Chemical Formula	weight	HLB	fh

Brij

97	POE (10) oleyl ether	C₁₈H₃₅(OCH₂CH₂)₁₀OH	709	12.4	0.47
Brij 98	POE (20) oleyl ether	C₁₈H₃₅(OCH₂CH₂)₂₀OH	1149.5	15	0.31
Brij 78	POE (20) stearyl ether	C₁₈H₃₇(OCH₂CH₂)₂₀OH	1151.5	15	0.31
Brij 700	POE (100) stearyl ether	C₁₈H₃₇(OCH₂CH₂)₁₀₀OH	4670	18.8	0.08

The rate of surfactant release from surfactant-laden contact lenses needs to be measured because an excessive release could lead to toxicity. Additionally, the rate of surfactant release impacts drug release. As described in the previous section, the surfactant concentration in the release medium was determined by measuring the surface tension of the solution, which is related to the bulk concentration (Data not shown). A model for surfactant release from hydrogels laden with surfactant aggregates has been proposed earlier [Kapoor Y, Chauhan A. Drug and surfactant transport in Cyclosporine A and Brij 98 laden p-HEMA hydrogels. Journal of Colloid and Interface Sciences 2008; 322:624-633]., and it predicts the following equation to describe surfactant release at short times,
$\begin{matrix} % Surfactant Release = \sqrt{2 D_{S} C^{*}} \sqrt{\frac{t}{C_{P} h^{2}}} & (2) \end{matrix}$
where D_Sis the surfactant diffusivity, C* is the critical aggregation concentration (CAC), i.e., the concentration beyond which the surfactant aggregates inside the gel, t is time, Cp is the concentration of the surfactant present as aggregates inside the hydrogel, and h is the half-thickness of the hydrogel. The above equation is valid for diffusion of any solute that is loaded in the gel above saturation limit and so a fraction of the solute precipitates into aggregates. This equation is the equivalent of the Higuchi equation that is commonly utilized to model drug release from gels [Higuchi T. Rate of release of medicaments from ointment bases containing drugs in suspension. Journal of Pharmaceutical Sciences 1961; 50(10):874-875].
To validate the model and to understand the mechanism of surfactant transport, surfactant release studies were conducted from gels of different surfactant loadings (approximately 2%, 4% and 8% w/w in drug gel) and different gel thicknesses (˜100 and 200 μm in dry state). FIG. 8 a-c show the surfactant release from 100 μm thick and 200 μm thick gels with three different surfactant concentration explored for 200 μm thick gels. The data is re-plotted in FIG. 9 a-c with time axis rescaled to
$θ \equiv \sqrt{\frac{t}{C_{p} h^{2}}},$
where t is time in seconds, Cp is concentration in M and h is half-thickness of the gel in μm. The model (Equation 2) predicts that the rescaled plots should overlap for all thicknesses and surfactant loadings, and the plots should be linear with slope the of √{square root over (2D_SC*×100)}, which agrees with all the experimental results for Brij 97 and Brij 78 laden hydrogels within 95% CI. However, for the Brij 700 surfactant, the data matches the model only for 4% and 8% surfactant loading, while thickness scaling and surfactant release from gels containing 2% surfactant in dry gel do not follow the predicted behavior. The fact that association of Brij 700 with the gel matrix is strongly dependant on the gel thickness may be due to large number of ethylene oxide (EO) units in the surfactant. Also, the CAC values for the Brij 700 laden system might be closer to 2% surfactant loading where the model assumption that C*<<C_Pdoes not hold.
To verify this hypothesis, Brij 700 release was also conducted from gels with 1% surfactant loading and the results are plotted in FIG. 8( c). The percentage release from both 2% and 1% surfactant loaded gels overlapped within 95% CI showing that release of surfactant at these concentrations is independent of initial surfactant loading suggesting that there are no or insignificant number of aggregates inside the gel for surfactant loadings as large as 2%, validating our hypothesis.
The slope indicated in FIG. 9( a-c) represents the value of √{square root over (2D_SC*×100)}. We calculated the value of D_SC* for all the systems and they are reported in Table 4.

TABLE 4

D_SC* for all systems obtained from fitting
of the surfactant release data.

		D_SC*
System	Slope	(μm)²M/s × 10⁶

Brij 97	0.4802 ± 0.010	11.52
Brij 98	1.1000 ± 0.013	60.80
Brij 78	0.1846 ± 0.050	1.70
Brij 700	0.0613 ± 0.013	0.19

The value of this parameter for Brij 98 surfactant was calculated earlier, and is also reported in Table 4 for comparison [Kapoor Y, Chauhan A. Drug and surfactant transport in Cyclosporine A and Brij 98 laden p-HEMA hydrogels. Journal of Colloid and Interface Sciences 2008; 322:624-633]. As mentioned above, the critical aggregation concentration for Brij 700 in p-HEMA gels is more than 1%, and so the release data from gels loaded with 1% (w/w) can be fitted to the following equation to obtain the surfactant diffusivity [Ritger P L, Peppas N A. A simple equation for description of solute release 1. Fickian and non-Fickian release from mom-swellable devices in the form of slabs, spheres, cylinders and discs. Journal of Controlled Release 1987; 5:23-36].
$\begin{matrix} % Release = \sqrt{\frac{4 D_{S} t}{π h^{2}}} \times 100. & (3) \end{matrix}$
Here D_Sis the diffusivity of the surfactant and h is the half-thickness of the gel. The diffusivity of Brij 700 from these hydrogels was determined to be 7.85×10⁻¹⁷m²/s and subsequently the value of C* was evaluated from the already determined parameter ‘DsC*’ (Table 4) to be 2.4±1.1 mM, which is equivalent to 1.12±0.51% surfactant in a dry gel. This value is in reasonable agreement with the hypothesis C* for Brij 700 is approximately 2%, which was based on the overlapping percentage release data for 1% and 2% Brij 700 loaded gels.

CyA Release—Equilibrium Experiments

A large fraction of hydrophobic drugs such as CyA are expected to partition inside the surfactant aggregates. During the drug release process, the drug molecules have to first diffuse through the surfactant head region into the p-HEMA gel, and subsequently diffuse through the gel. The head group of the surfactants may offer resistance to transport of the molecules, and this transport could potentially be rate limiting. However if the resistance to transport from the surfactant aggregates to the p-HEMA gel is small, the concentrations inside the aggregates will be in equilibrium with that in the p-HEMA gel, and in this case, diffusion through the p-HEMA gel will be rate controlling. If transport through the gel is rate controlling, the time scale for drug release scales as the square of the gel thickness, and if the transport across the aggregates is rate controlling, the time scale for release should be independent of the gel thickness. To investigate the rate limiting step, drug release studies were conducted from 100 μm and 200 μm thick gels with 8% surfactant loading. It is noted that the gel weight and the fluid volume in the release medium was kept the same for gels of both thicknesses. To determine the rate limiting process, percentage release of the drug against ‘τ’ is plotted where
$τ = \sqrt{\frac{Release Time}{i}},$
i=1 for 100 μm thick gels and 2 for 200 μm thick gels in FIGS. 10( a-d). The data for different thicknesses for all the systems overlaps proving that the transport of drug in all cases is controlled by diffusion through the gel, and that the drug concentrations inside the aggregates and in the p-HEMA gel are in equilibrium. Data from pure p-HEMA gels is also shown to assert that the rate limiting step is diffusion for drug transport from these hydrogels in absence of surfactants.
The data in FIGS. 10( a-d) also show that the time required to reach equilibrium is much greater for surfactant laden gels then for pure p-HEMA gels and Brij 78 systems take the longest time to equilibrate. Also, the percentage of drug that diffuses out, until equilibrium is attained, is different for each system, with the relative order Brij 97 laden gels <Pure p-HEMA<Brij 78<Brij 700<Brij 98. To explain these observations, a qualitative picture of the system is proposed in FIG. 11. In the gel matrix, surfactants can exist in three forms: a) free form, not interacting with other surfactants or the polymer b) as micellar aggregates or c) interacting with the polymer matrix. Similarly, the drug also exists in three different forms: a) free form, b) inside micellar aggregates, c) adsorbed on the polymer. Although diffusion of drug through the gel is the rate controlling step, only the drug present in the matrix, which is a small fraction of the total drug, can directly diffuse. This creates a depot effect which prolongs the total release duration. The partitioning of the drug will likely depend on the hydrophobicity of the core of the surfactant aggregates, which will depend on the surfactant chain length and the packing in the aggregates. Also, due to surfactant diffusing out during equilibrium experiments, the drug solubility can also increase in the release medium. It is thus probable that the presence of surfactants in the release medium can affect the amount of drug diffusing out. This interaction between the drug and the surfactant in the release medium should be dominant only above CMC values and it is possible that for Brij 97 system, the surfactant concentration in the bulk phase does not reach CMC, while for other systems it does reach above CMC value, leading to smaller percentage drug release for Brij 97 laden gels than pure p-HEMA systems. Essentially, the combined interactions of drug molecules with the surfactants present inside the gel matrix and outside in the release medium would then determine the equilibrium percentage release from the hydrogel.

CyA Release—PBS Change Experiments

FIGS. 12( a-c) show the percentage release of the drug with different surfactant loading for all the surfactant laden gels that are 200 μm in thickness and loaded with 50 μg of drug. As expected, the percentage release of the drug decreases as the surfactant loading is increased inside the hydrogel. A model for drug release from hydrogels laden with surfactant aggregates has been proposed earlier, [Kapoor Y, Chauhan A. Drug and surfactant transport in Cyclosporine A and Brij 98 laden p-HEMA hydrogels. Journal of Colloid and Interface Sciences 2008; 322:624-633] and it predicts the following equation to describe drug release at short times,
$\begin{matrix} % Drug Release = β^{*} \sqrt{\frac{t}{h^{2}}} \times 100 Where & (4) \\ β^{*} = \frac{(1)}{\sqrt{\begin{matrix} \frac{π (1 + {KC}_{P})}{4 D} - (C^{*} D_{S}) \frac{K}{D} \sqrt{\frac{π (1 + {KC}_{P})}{4 D}} + \\ {(C^{*} D_{S})}^{3 / 2} \frac{K (1 + {KC}_{P})}{D^{2} \sqrt{2 C_{P}}} \end{matrix}}} & (5) \\ \frac{K_{m} M W_{S} f_{h}}{ρ} . & (6) \end{matrix}$

Here,

D≡Diffusivity of drug
Km≡Partition coefficient of the drug defined as ratio of drug concentration inside the micelle to drug concentration in the hydrogel.
MW_S≡Molecular weight of the surfactant
ρ≡Density of micelles
To establish the validity of the model and to determine all the model parameters, the percentage release of the drug against ‘γ’ was re-plotted where
$γ = \sqrt{\frac{t}{h}}$
where t is time in seconds and h is half thickness of the gel in μm. The re-plotted data for systems loaded with Brij 98 surfactant is provided in FIGS. 13( a-d). The slopes between various surfactant loadings for each system differ as the 95% CI for the slopes do not overlap. The release data was fitted to a straight line to obtain β* and then Equations (5) and (6) were used to obtain Km. The parameters determined for each system are listed in Table 6.

TABLE 6

Partition coefficient of CyA for all the surfactant systems

	Amount of
System	surfactant	Slope	K (M⁻¹)	Km

Brij

97	2%	11.08 ± 0.83	14.14	48.9 ± 6.3
	4%	9.22 ± 0.2	18.35
	8%	7.76 ± 0.33	16.40
Brij 78	2%	6.37 ± 0.43	142.20	458.9 ± 61.5
	4%	4.51 ± 0.40	163.10
	8%	3.19 ± 0.33	186.10
Brij 98	2%	9.17 ± 0.2	99.46	261.0 ± 24.4
	4%	7.46 ± 0.16	96.46
	8%	5.71 ± 0.07	83.10
Brij 700	2%	13.87 ± 2.32	—	675.8 ± 11.4
	4%	7.26 ± 0.67	258.48
	8%	5.33 ± 0.17	252.47

The values of Km are relatively independent of the surfactant loading, thus validating the model. In these calculations, the diffusivity of the drug was taken to be 1.44×10⁻¹⁴m²/s and p was taken to be 1000 kg/m³. The value of K for 2% Brij 700 loading in the system could not be determined which is expected since it was shown earlier that at this concentration is close to the CAC value for this surfactant. Furthermore, in FIG. 12( c) the percentage release of CyA from p-HEMA gels is shown to overlap that from the 2% Brij 700 loaded gels, again suggesting that there was no significant partitioning inside surfactant aggregates due to an insignificant number of such aggregates inside the hydrogel.
The partition coefficient of the drug between the surfactant aggregates and the p-HEMA gels is highest for the Brij 700 laden gels and smallest for the Brij 97 laden hydrogels. Even though the value of Km is highest for Brij 700, these systems do not attenuate drug release significantly for two reasons. First, the value of C* is large (˜2%) implying that only a small amount of surfactant is available for forming aggregates. Second, due to the large molecular weight, the volume fraction of the hydrophobic core which provides the site for drug partitioning is small.
The values of the partition coefficients clearly suggest that an increase in the hydrophilic chain length leads to an increase in the partition coefficient. This is most likely due to an improved shielding of the hydrophobic core from water on increasing the hydrophilic chain length. Also, a comparison of Brij 78 and Brij 98 systems shows that there is a significant increase in the partition coefficient of the drug if the hydrophobic tail of the surfactant is saturated (Brij 78). This could be attributed to the fact that an unsaturated chain is more rigid than a saturated chain, and an increase in rigidity will likely lead to reduced packing resulting in a reduced shielding of the core from water, and a consequent reduction in the partitioning of hydrophobic drugs.
The results reported above show that amongst the systems explored, Brij 78 is the most suitable candidate for extended drug delivery. The toxicological response of this surfactant on the ocular surface has been investigated in rabbits and it is reported that administration of 0.1 ml of 0.2% (20000 μg/ml) eye drops does not cause toxic effects [Saettone M F, Chetoni P, Cerbai R, Mazzanti G, Braghiroli L. Evaluation of ocular permeation enhancers: in vitro effects on corneal transport of four β-blockers, and in vitro/in vivo toxic activity. International Journal of Pharmaceutics 1996; 142:103-113]. Assuming a bioavailability of about 2%, about 40 μg of the surfactant delivered in the drop reaches the cornea without causing any toxicity. The same study also showed that exposure to 0.05% Brij 78 solution for about 5 hours does not lead to any significant increase in corneal hydration again suggesting that Brij 78 has negligible toxicity even with extended exposure. The surfactant release studies reported here demonstrate that about 10% of the Brij 78 loaded in the 100 micron thick gel with 8% loading is released in a period of 10 days, which corresponds to around 300 μg of surfactant or equivalently an average release of about 30 μg/day. When a contact lens is placed on an eye, at the most half this amount, i.e., 15 μg/day will be released into the post lens tear film, which is the thin tear film in between the cornea and the contact lens. Ocular conditions are likely not perfect-sink conditions and so the amounts released would be less than this level. Thus, it might be expected that the lenses loaded with Brij 78 will not cause any toxicity even if the lenses are worn continuously for a few days.

Uptake of Lysozyme in the Hydrogels

Binding of tear proteins to contact lenses is undesirable as it can lead to increased bacterial binding to the lenses. Lyzozyme is the main protein present in tears, and it is frequently used as the test protein to investigate protein binding to contact lenses [Lord M S, Stenzel M H, Simmons A, Milthrope B K. Lysozyme interaction with poly(HEMA)-based hydrogel. Biomaterials 2006; 27:1341-1345]. It may be speculated that the presence of hydrophobic regions inside the surfactant-laden contact lenses can lead to increased protein binding. To test this hypothesis, hydrogels (200 μm thick) were soaked in lyzozyme solution, and the mass of drug bound to the lens was determined by assaying the free lyzozyme concentration through UV-Vis spectrophotometry. Lysozyme uptake by the surfactant laden gels and pure p-HEMA gels is shown in FIG. 14. It is observed that the presence of surfactant in the gels does not have any significant effect on lysozyme adsorption in the gel matrix.
Microstructure of Hydrogels: cryo-SEM Imaging
FIGS. 15( a-o) show a series of SEM images of the cross-sections of p-HEMA and surfactant laden gels. The pure p-HEMA gels have a uniform structure with no visible pores, which is expected because pores in p-HEMA gels are a few nm in size. The microstructure of surfactant laden gels is in sharp contrast to pure p-HEMA gels, as they show a uniform distribution of pores with pore sizes ranging from 40-50 nm. The size of these pores is much larger than the expected micelle size suggesting that the structure of the surfactant aggregates inside these pores is likely more complex than micelles. The volume fraction of the pores is much larger than the surfactant loading, which implies that these pores are likely water rich environments, and so the presence of these pores should lead to an increased water content in the gels. To investigate this issue further, the area fraction of pores in the hydrogels was determined by image analysis using ImageJ (National Institute of Health) software and these values are listed in Table 8.

TABLE 8

Physical properties of the surfactant
laden and pure p-HEMA hydrogels.

		EWC_Pred	EWC	Transmit-	Contact
System	α	(%)	(%)	tance (%)	Angle

Brij
97	31.0	67.6	64.1	99.8	27.9 ± 1.06
Brij 98	29.1	66.7	67.2	99.2	24.9 ± 1.4
Brij 78	41.0	72.3	70.3	99.9	19.8 ± 0.78
Brij 700	46.4	74.8	70.4	99.5	27.2 ± 0.59
Pure	—	—	53.0	98.9	30.3 ± 0.18
p-HEMA

If we assume that after hydration these pores are filled with water, then the equilibrium water content of the systems can be predicted by the following equation
$\begin{matrix} {EWC}_{pred} (%) = (\frac{{EWC}_{p - HEMA} (1 - α)}{100} + α) \times 100 & (8) \end{matrix}$

Where

α=fraction of area occupied by pores, and
EWC_p-HEMA=Water uptake in pure p-HEMA hydrogel.
The values of water content (EWC_Pred) from Equation 8 are also listed in Table 8. A quantitative analysis of the pore structure along with the equilibrium water content shows that these pores are essentially filled with water when the gels are hydrated. Together with the 50 nm size suggests that the surfactant structures could possibly be vesicles.

Physical Properties

The surfactant-laden gels were characterized by several techniques to investigate the impact of surfactant loading and the resulting porous structure and also to determine the suitability of these systems as a typical contact lens. The studies set forth in this Example Set Two used gels which contained 8% surfactant per dry gel weight.

Mechanical Properties

The storage moduli (G′) and the loss moduli (G″) for the five systems explored in this Example Set Two are plotted as a function of frequency in FIGS. 16( a-b). All the gels explored for studying the mechanical properties were 400 μm and 800 μm thick. The results show a negligible effect of surfactant loading on the mechanical properties. The elastic modulus G′ continuously increases with increasing frequency and the values of the modulus at frequencies approaching zero for all the systems is close to the desired value for commercial lenses [Alvarez-Lorenzo C, Hiratani H, Mez-Amoza J L G, Martinez-Pacheco R, Souto C, Concheiro A. Soft contact lenses capable of sustained delivery of timolol. Journal of Pharmaceutical Sciences 2002; 91(10): 2182-2192]. The loss modulus first increases and then decreases with frequency. The mechanical response of a contact lens plays an important role in lens settling, lens shape, and the pressure distribution on the post-lens tear film. It may thus be useful to obtain a model that can describe the mechanical properties of the lenses so that this model can be utilized in modeling lens deformation in the eyes due to application of the forces during blinking. To model the frequency dependence of the storage and loss modulus, three parameter ‘Standard Linear Solid Model’ was used, as illustrated in FIG. 17. In this model there is an elastic spring connected to a viscous dashpot in series like a Maxwell model, with an addition of an elastic spring in parallel. To determine the expressions for G′ and G″, it is instructive to consider application of a periodic strain ε₀e^iωt. The stresses and strains in the individual elements are then related by the following expressions
ε₁=ε₂+ε₃=ε₀ e ^iωt (9)
ε₂E₂=ε₃μ (10)
The complex modulus G (=G′+iG″) is the ratio of the stress and the strain, and it can be obtained from the following expression
$\begin{matrix} G^{'} +  G^{″} = \frac{ɛ_{1} E_{1} + ɛ_{2} E_{2}}{ɛ_{0} e^{ ω t}} & (11) \end{matrix}$
Using Equations 9, 10 and 11 we can evaluate G′ and G″ to be,
$\begin{matrix} G^{'} = \frac{E_{1} E_{2}^{2} + (E_{2} + E_{1}) ω^{2} μ^{2}}{E_{2}^{2} + ω^{2} μ^{2}} & (12) \\ G^{″} = \frac{E_{2}^{2} ω μ}{E_{2}^{2} + ω^{2} μ^{2}} & (13) \end{matrix}$
The experimental data averaged from all the systems was fitted to the model, and the model fit is also plotted in FIGS. 16( a-b), and the parameters are listed in Table 7.
TABLE 7

Parameters obtained by fitting Standard Linear

Solid Model to the experimental data.

Parameter Value

E₁ 0.55 MPa

E₂ 0.90 MPa

μ 0.0076 MPa · s

The data fits the model reasonably for lower frequencies, while there is a clear deviation from the model at frequencies greater than 25 s⁻¹. A more generalized Maxwell model is needed for the data at higher frequencies but physiological blink frequencies are about 10 s⁻¹and so the model proposed above is may be adequate. The loss moduli of the gels may partially be attributed to water flow during the gel stretching. The contribution of water flow to the moduli could be explored by measuring the moduli for gels of different thicknesses. Since water transport depends on the gel thickness, a significant dependence of the loss modulus on the thickness will indicate that water transport is an important contributor to the loss modulus. The storage and the loss modulus for pure p-HEMA gels are relatively independent of the gel thickness as shown in FIG. 18. This suggests that water transport during gel stretching does not contribute to the loss modulus for the p-HEMA gels. The results were similar for the surfactant-laden gels and are not shown in FIG. 18 for clarity of presentation. Transparency, Equilibrium water content, and surface contact angle of gels
All the gels explored in this study were visibly transparent and clear, and 100 μm thick hydrated gels had transmittance values larger than 98.5% (Table 8) at a wavelength of 600 nm, and so are suitable for contact lens application. It was also observed that the surfactant laden gels had a higher transparency than the pure p-HEMA gel likely due to the higher water content.
The EWC of contact lenses is crucial as it likely impacts comfort, and also an increase in the EWC leads to an increase in the oxygen permeability of p-HEMA contact lenses. The EWC values of all the gels obtained by hydrating 200 μm thick gels are listed in Table 8. All the surfactant systems had higher water content than pure p-HEMA gels, and thus are more suitable for contact lens application. The water content appears to be a function of length of hydrophilic part of the surfactants and it increases as the length of (EO) groups of the surfactants increase. The measured values match the EWC predictions based on analysis of SEM images. This suggests that water content inside the hydrogel increases due to formation of pores which are filled with water and since the pore size is less than 40 nm for all the surfactant laden hydrogels, these remain visibly transparent after polymerization.
The contact angles were measured for all systems to determine the effect of surfactants on the wettability of the hydrogels. Thickness of the gels utilized for measuring the contact angle was 200 μm. It has been shown previously that surfactants can significantly alter the contact angles of hydrogels [Tonge S, Jones L, Godall S, Tighe B. The ex vivo wettability of soft contact lenses. Current Eye Research. 2001; 23(1):51-59; Ketelson H A, Meadows D L, Stone R P. Dynamic wettability properties of a soft contact lens hydrogel. Colloids and surfaces B. Biointerfaces 2005; 40:1-9.]. The captive bubble technique was utilized to determine the contact angle to ensure that the gel remains hydrated during the measurements. The values of contact angles listed in Table 8 are all lower than those for the p-HEMA gels likely due to the presence of surfactant on the surface making surfactant laden gels more suitable for contact lens application.

SUMMARY

Brij surfactant laden p-HEMA hydrogels were explored for ophthalmic drug delivery by contact lenses. The surfactants explored herein have the same length of the hydrophobic group but differed in chain length of the hydrophilic unit (EO) and in presence of an unsaturated group in the tail region of the surfactant. Ophthalmic drugs were loaded in these systems by direct addition to the polymerizing mixture. The results reported here clearly prove that the duration of CyA release can be significantly reduced by incorporation of surfactants inside the gel matrix by as much as a factor of five compared to pure p-HEMA gels. The mechanism of reduction in release rates is through a preferential partitioning of the drug into the surfactant domains that form inside the gels. The rate controlling step is diffusion through the gel but its rate is reduced due to a reduction in the free drug concentration. The concentration and type of the surfactant plays an important role as an increase in the hydrophilic chain length increases partitioning into the hydrophobic cores of the surfactant aggregates and the presence of a double bond in the hydrophobic chain reduces the partitioning. Both these effects likely occur due to the impact of these factors on shielding of the hydrophobic core from the water molecules. Amongst the surfactants explored here, Brij 78 is most promising for extended release of CyA from p-HEMA contact lenses due to a high partition coefficient of 458.9±61.5 for partitioning between the hydrophobic cores and the p-HEMA matrix. The partition coefficient is even higher for Brij 700 due to the large hydrophilic chain length but the total partitioning is small due to a smaller fraction of the hydrophobic segment and also due to a relatively large value of the critical aggregation concentration in the gel. Furthermore, Brij 78 surfactants have been used in ocular studies as cornea permeability enhancers, and have been shown to have negligible toxicity at concentrations as large as 2% (w/w) [Saettone M F, Chetoni P, Cerbai R, Mazzanti G, Braghiroli L. Evaluation of ocular permeation enhancers: in vitro effects on corneal transport of four β-blockers, and in vitro/in vivo toxic activity. International Journal of Pharmaceutics 1996; 142:103-113]. Freeze fracture SEM imaging provides a direct evidence of the presence of pores in the surfactant-laden hydrogels. The pores are about 50 nm in size and are filled mostly with water, which suggests that the surfactant aggregates are perhaps more complex than micelles, and could possibly be vesicles.
The surfactant-laden gels were characterized to determine their suitability as contact lenses materials. All the systems were clear and transparent, and had storage moduli suitable for contact lens applications. The water content for all the surfactant laden gels was much higher than that for pure p-HEMA hydrogel, which is encouraging as it will increase oxygen permeability and perhaps comfort. The wettability of the lenses also improved due to surfactant entrapment with maximum improvement with Brij 78, which could also have beneficial effects.
Because silicone hydrogels are a mix of silicone and hydrogel materials, the creation of both hydrophobic and hydrophilic domains is contemplated. This may be useful for extended release of both hydrophobic and hydrophilic drugs. In addition to the surfactants explored here, other surfactants and self assembling molecules such as lipids, and block-co-polymers could be used to create domains that could trap and slowly release hydrophobic and/or hydrophilic drugs. Thus, the studies disclosed herein provide conclusive evidence that p-HEMA contact lenses loaded with Brij 78 surfactant have physical properties suitable for contact lens applications, and also for extended drug delivery.
Having now described a few embodiments of the invention, it should be apparent to those skilled in the art that the foregoing is merely illustrative and not limiting, having been presented by way of example only. Numerous modifications and other embodiments are within the scope of one of ordinary skill in the art and are contemplated as falling within the scope of the invention and any equivalent thereto. It can be appreciated that variations to the present invention would be readily apparent to those skilled in the art, and the present invention is intended to include those alternatives. Further, since numerous modifications will readily occur to those skilled in the art, it is not desired to limit the invention to the exact construction and operation illustrated and described, and accordingly, all suitable modifications and equivalents may be resorted to, falling within the scope of the invention. Each reference cited herein is hereby incorporated by reference in its respective entirety.

Claims

1. A bioactive agent delivery system comprising a contact lens having dispersed therein (1) an ophthalmically bioactive agent, said agent being capable of diffusion through said contact lens and into the post-lens tear film when said contact lens is placed on the eye and (2) associated with said bioactive agent, at least one ophthalmically compatible polymeric surfactant, said surfactant being present in an amount sufficient to attenuate the rate of migration of said bioactive agent through said contact lens.

2. A bioactive agent delivery system of claim 1 wherein said polymeric surfactant comprises a block copolymer having a hydrophobic segment and a hydrophilic segment

3. A bioactive agent delivery system of claim 2 wherein said bioactive agent is hydrophobic and is associated with the hydrophobic segment of said surfactant.

4. A bioactive agent delivery system of claim 1 wherein said contact lens comprises a polymer formed from a reaction mixture comprising at least one hydrophilic monomer.

5. A bioactive agent delivery system of claim 4 wherein said hydrophilic monomer is an unsaturated carboxylic acid; acrylic substituted alcohol; vinyl lactam or acrylamide.

6. A bioactive agent delivery system of claim 5 wherein said hydrophilic monomer is selected from the group consisting of methacrylic or acrylic acid, 2-hydroxyethylmethacrylate, 2-hydroxyethylacrylate, N-vinyl pyrrolidone, methacrylamide, N-vinyl-N-methacetamied, N,N-dimethylacrylamide, and mixtures thereof.

7. A bioactive agent delivery system of claim 1 wherein said bioactive agent is hydrophilic and is associated with the hydrophilic segment of said surfactant.

8. A bioactive agent delivery system of claim 1 wherein said contact lens comprises a hydrophobic material.

9. A bioactive agent delivery system of claim 8 wherein said hydrophobic material is selected from the group consisting of a silicone, silicone containing prepolymers and macromers, polydimethylsiloxane, pentamethyldisiloxanyl methylmethacrylate, tris(trimethylsiloxy)methacryloxy propylsilane, methyldi(trimethylsiloxy)methacryloxymethyl silane, monomethacryloxypropyl terminated mono-n-butyl terminated polydimethylsiloxane, mono-(3-methacryloxy-2-hydroxypropyloxy)propyl terminated, mono-butyl terminated polydimethylsiloxane, bis-3-methacryloxy-2-hydroxypropyloxypropyl polydimethylsiloxanes, 3-methacryloxy-2-hydroxypropyloxy)propylbis(trimethylsiloxy)methylsilane, 3-[tris(trimethylsiloxy)silyl]propyl vinyl carbamate, methyl methacrylate, ethylene glycol di-methacrylate, and mixtures thereof.

10. A method of manufacturing a bioactive agent delivery system of claim 1 comprising providing a monomer mixture comprising a lens-forming monomer, the surfactant and the bioactive agent and polymerizing said monomer mixture.

11. A method of manufacturing a bioactive agent delivery system of claim 1 without introducing a bioactive agent before polymerization.

12. A method of loading drug into the system of claim 11 by soaking the device in at least one drug solution.

13. A method of loading drug into the system of claim 11 by packaging the device in drug solutions.

14. A method of administering an ophthalmically bioactive agent to the eye comprising contacting the eye with the bioactive agent delivery system of claim 1.