Porous nanocomposite medical implant device

Abstract

The invention relates to a drug-eluting nanoengineered medical implant / contact device. The device comprises a nanocomposite and a drug, wherein the nanocomposite comprises a hydrophilic polymer domain, a hydrophobic polymer domain, a water pore, and a boundary charged double layer; wherein when the drug is hydrophilic, at least 80% of the drug partition in the boundary charged double layer is formed at the boundary interface of the hydrophilic polymer domain and the water pore, and when the drug is hydrophobic, at least 80% of the drug partition in the boundary charged double layer is formed at the boundary interface of the hydrophobic polymer domain and the water pore. The device is configured to maintain release of the drug with high precision and long duration.

Description

TECHNICAL FIELD

[0001] The present invention relates to a drug-eluting nanoengineered medical implant / contact device (NMID) comprising an active drug (pharmaceutical or growth ingredient), an aqueous pore, one or more hydrophilic domains, one or more hydrophobic domains, and a boundary charged double layer; the device is configured to maintain release of the drug with high precision and long duration (over a week). The present disclosure describes the preparation of porous nanocomposites with sufficient hydrophilicity, ionic permeability, and gas permeability, and methods to exploit the electrostatic forces of the charges implanted on the pore-polymer interface (interfacial charge layer) to enhance the partitioning and retention of oppositely charged drugs upon delivery by the engineered device. The NMID will be used to deliver active drugs and other functional ingredients into the body in a controlled manner. BACKGROUND

[0002] The technology of designing and manufacturing medical implants has evolved through three generations, starting from the primitive state of using only inert materials, through the second stage of addressing the body's immune resistance, usually by including biocompatible surface coatings such as PEG, and now reaching the sophisticated stage of achieving good interaction with the body's systems, not only to heal wounds, but also to aid in the complete regeneration of the body. The latest biomedical devices will be intentionally designed to pharmacologically intervene in the host body functions by delivering pre-embedded functional pharmaceutical ingredients, including but not limited to, basic building blocks, nutrients, genes, growth factors, signaling peptides, and / or stem cells, to stimulate and accelerate the desired cellular responses, healing, and regenerative reactions.

[0003] The core technology of the new biomedical devices is to nanoengineer the composite materials to formulate the device body or its external coating, to facilitate the sustainable delivery of one or more drugs at high precision dosing, nearly constant rate, and extended duration. The latest advances in contact lens material technology provide a platform for further development of this core technology. The silicone-hydrogel composite technology has improved the hydrophilicity and oxygen permeability of contact lenses and transformed them into gas-permeable hydrophilic soft tissue-like devices since its realization in 1997.

[0004] U.S. Patent No. 10,617,559 discloses a device comprising: (i) at least one drug, (ii) one or more reservoir domains, and (iii) a layer configuration barrier domain (or barrier layer) for blocking the diffusion path of the drug from the reservoir domain to the ocular surface of the subject's eye.

[0005] The prior art shows that there are three modes of transport for drug permeation through porous composites. These results show that after the lens is immersed in the tear pool, the dissolved drug in the aqueous pores is released in a burst over a period of a few hours. The second mode of release, which is the release of drug adsorbed at the domain-pore interface, continues for more than a day. Drug trapped within the solid polymer domains comes out the slowest and can continue for weeks depending on the relaxation time scale of the solid polymer (related to the glass transition temperature of the polymer). Many efforts have been directed at extending the release of the majority (80%) of the drug to more than a week, including the use of barrier coatings, dissolution-diffusion barriers, or charged surfactants.

[0006] There is a need for a drug eluting medical implant / contact device that is safe and provides sustained release of a drug. BRIEF DESCRIPTION OF DRAWINGS

[0007] Figure 1 Drug permeation through porous nanocomposites.

[0008] Figure 2 Three modes of transport for drug through porous nanocomposites.

[0009] Figure 3 Dissociation of oleic acid molecules at the pore-domain interface enhances partitioning and retention of cationic drugs.

[0010] Figure 4 Correlation of oleic acid loading with soak solution concentration for two different silicone hydrogel commercial lenses: ACUVUE and ACUVUE R2 of the linear line fit was 0.996 and 0.998 Data is represented as mean + / - standard deviation, where n = 3.

[0011] Figure 5 Cumulative % release of THCL from oleic acid loaded contact lenses. Contact lenses.

[0012] Figure 6 Cumulative % release of THCL from oleic acid loaded contact lenses. Contact lenses.

[0013] Figure 7 Cumulative % release of BUP from oleic acid loaded contact lenses. Contact lenses.

[0014] Figure 8 Cumulative % release of BUP from oleic acid loaded contact lenses. Contact lenses.

[0015] Figure 9 Cumulative release of KTF from oleic acid-loaded contact lenses. Contact lens.

[0016] Figure 10 Cumulative release of KTF from oleic acid-loaded contact lenses. Contact lens.

[0017] Figure 11 Cumulative release percentage of THCL from oleic acid-loaded contact lenses. (Daily ACUVUE) Contact lens.

[0018] Figure 12 Cumulative release percentage of KTF from oleic acid-loaded contact lenses. (Daily ACUVUE) Contact lens.

[0019] Figure 13 Cumulative release of DFNa from oleic acid-loaded contact lenses. Contact lens.

[0020] Figure 14 Cumulative release of DFNa from oleic acid-loaded contact lenses. Contact lens.

[0021] Figure 15 Cumulative release of FBNa from oleic acid-loaded contact lenses. Contact lens.

[0022] Figure 16 Cumulative release of FBNa from oleic acid-loaded contact lenses. Contact lens.

[0023] Figure 17 ACUVUE loaded with 6% by weight of oleic acid, linoleic acid, or α-linolenic acid Cumulative release of KTF from contact lenses.

[0024] Figure 18 The cumulative release of DFNa from contact lenses loaded with 6.5 mg / mL sphingosine. Contact lens.

[0025] Figure 19 ACUVUE Cumulative release percentage of KTF from commercial silicone hydrogel lenses. Fatty acids in the contact lenses were loaded at a fixed immersion concentration of 7.5 mg / mL. The control group was 176.3 μg ± 18.6 μg. 7.5 mg / mL stearic acid, which is 214.3 μg ± 38.7 μg, for 7.5 mg / mL Myristic acid, 516.4 μg ± 4.2 μg.

[0026] Figure 20 Ketotifen uptake (in the range of 70 μg - 375 μg / CL KTF loading) was linearly dependent on the solution loading concentration.

[0027] Figure 21 With respect to contact lenses loaded with 2% oleic acid (oleic acid loading conditions of 15 mg / ml ethanol), ketotifen loading and release had consistent delivery kinetics profiles over a wide range (70 μg - 375 μg / CL). DETAILED DESCRIPTION

[0028] Definitions

[0029] As used herein, "medical implant / contact device" refers to a device that is in contact with a body part or implanted in a body environment to intervene in a specific body function or condition. Examples are, but are not limited to, surgically implanted body parts, stents, chemotherapy pumps, tissue growth scaffolds, contact lenses, dental implants, skin pads, wound healing bandages, etc.

[0030] As used herein, "drug" or "drug compound" refers to a molecule with activity that causes a physiological change in a subject, such as a pharmaceutical drug or a nutrient.

[0031] "Charged drug or drug compound" means that the molecule is cationic or anionic under normal physiological conditions. They can belong to the class of compounds with pKa, pKb, or isoelectric point, including but not limited to proteins, small peptides, molecules with other Lewis acid or base groups,

[0032] "Hydrophobic / hydrophilic" ratio is a measure of the hydrophobicity of a material based on its solubility and partitioning between the hydrophobic octanol and the hydrophilic water. It is commonly referred to as the octanol / water material constant logP, where P is the partition coefficient ratio of the soluble component between octanol and water (Sangster J, J Phys Chem Ref Data 1989; 18: No. 3). High logP values mean hydrophobicity, and low or negative values mean hydrophilicity.

[0033] "Junction potential" is the Coulomb potential that arises due to the diffusion of electrons or ions at the boundary of a conductor, semiconductor, or electrolyte solution. Examples are metal junctions, P-N junctions, liquid junctions, or cell aqueous media junctions.

[0034] As used herein in the present device, a "boundary charged double layer" is composed of a negatively charged layer and a positively charged layer, and is formed at the boundary interface of a hydrophilic / hydrophobic polymer domain and a water pore. The boundary charged double layer is generally composed of a surface Stern layer and a diffuse Debye layer, like those of a colloidal particle or an ionic crystal. For example, the boundary charged double layer is formed by the charge of the head group of the boundary charge modifier and the charge of (ii) the counterion of the drug or the charged head group.

[0035] A "boundary charge modifier" is a compound that is implanted onto a water-solid interface by physical dissolution or chemical bonding to generate a boundary charged double layer on the water-solid interface and to form a junction potential by interfacial ion dissociation or exchange in an aqueous medium.

[0036] A "good solvent" for a polymer is a solvent in which the molecules and monomeric segments of the polymer have good energetic interactions, such that the solvation effect of the solvent in the polymer causes the polymer coil to swell and its imbibition causes the polymer to swell.

[0037] Porous nanocomposites for drug delivery

[0038] The inventors have discovered a new generation of nanoengineered contact / implant biomedical devices with the following advantageous functions: (a) compatibility with and permeability by body fluids, (b) efficient exchange of oxygen and carbon dioxide, (c) strong adhesion of cells to the hydrophobic domains (hydrophobic bonding), (d) drug release rate that can be prolonged by composite material engineering.

[0039] The present invention provides a medical device comprising a drug and a nanocomposite. In a first aspect, the medical device comprises a hydrophilic polymer domain, a hydrophobic polymer domain, a water pore, and a charged double layer; wherein at least 80% of the drug partition in the charged double layer is formed in the interphase of the hydrophilic polymer domain and the water pore when the drug is hydrophilic, and at least 80% of the drug partition in the charged double layer is formed in the interphase of the hydrophobic polymer domain and the water pore when the drug is hydrophobic.

[0040] In a second aspect, the medical device comprises a hydrophilic polymer domain, a water pore, and a charged double layer; wherein at least 80% of the drug partition in the charged double layer is formed in the interphase of the hydrophilic polymer domain and the water pore.

[0041] In the present device, each layer in the boundary charged double layer is formed from (i) the charge of the head group of the boundary charge modifier, and (ii) the charge of the counterion of the drug or the charged head group, wherein the boundary charge modifier is a molecule with a charged head group and a hydrophobic tail and is fixed at the boundary charged double layer throughout the service life of the device.

[0042] The present invention also provides a method for manufacturing such devices.

[0043] The present disclosure provides the design of device composites consisting of aqueous pores, hydrophilic polymer domains (monomer logP < 1), and hydrophobic polymer (monomer logP > 3) domains, where the aqueous pores and hydrophilic polymers can facilitate wetting by body fluids (e.g., tear fluid or blood), while the hydrophobic domains can enhance gas (e.g., oxygen or carbon dioxide) permeation and adhesion of hydrophobic species (e.g., cells or proteins). The size of the pores and domains are controlled at the nanoscale to increase pore-surface area and interfacial interactions of the device. Various hydrophilic and hydrophobic polymer materials can be used in the present invention.

[0044] The present disclosure also outlines the compositional design of NMIDs for loading and delivering pharmaceutically active ingredients, where one component domain with high drug affinity is used as its reservoir, while other domains (with much lower drug solubility) are utilized as barriers to modulate the rate of drug release into body fluids. The water content is tunable to further modulate the overall drug delivery rate.

[0045] The present disclosure finally provides a method to significantly enhance or reduce the release of charged drugs by utilizing engineered domain morphologies (e.g., size, geometry, and surface area) and inclusion of boundary charge modifiers at the polymer-aqueous pore interface.

[0046] The porous nanocomposites of the present invention can be further bioengineered for various contact or in-situ medical devices, including but not limited to drug delivery matrices, implant devices, tissue engineering scaffolds, chemotherapy pumps, wound dressing patches, cosmetic masks, etc. The following sections show how the inventors can tailor the porous nanocomposites to fabricate devices that respectively achieve the following features: (a) high compatibility and permeability with body fluids, (b) efficient exchange of oxygen and carbon dioxide, (c) strong adhesion (hydrophobic bonding) of cells to the hydrophobic domains, (d) tunable drug release rates by composite engineering.

[0047] In one embodiment, the porous nanocomposite consists of: aqueous pores (10-80 wt% or 10-40 wt%), hydrophilic polymer domains made of hydrophilic monomers with logP in the range of -1 to 1 (10-50 wt%), hydrophobic polymer domains made of hydrophobic monomers with logP > 3 (10-50 wt%), and physically or chemically included boundary charged double layers (1-20 wt% or 1-10 wt%).

[0048] In another embodiment, the porous nanocomposite consists of: water pores (10-80 wt% or 10-40 wt%), hydrophilic polymer domains (10-90 wt%) made from hydrophilic monomers with logP in the range of -1 to 1, and physically or chemically included boundary charged double layers (1-20 wt% or 1-10 wt%).

[0049] The composition and morphological features (e.g., hydrophobicity / hydrophilicity ratio, domain size shape and orientation) are tailored in each application based on the required or desired specifications of body fluid wetting, oxygen exchange rate, cell adhesion and growth.

[0050] Porous nanocomposites compatible with body fluids, permeable to oxygen, CO2, and / or superior in promoting cell adhesion / growth Composite domain morphology for controlling drug delivery rate

[0051] In one embodiment, the hydrophilic domain component of the composite is selected from any of the hydrophilic components listed in Table 1A, Table IB, Table 1C to increase the hydrophilicity of the device, and the hydrophobic domain component is selected from any of the hydrophobic components listed in Table 1A, Table IB or Table 1C to enhance the oxygen permeability (Dk > 100) of the device to four to five times higher than that of a pure hydrogel device (Dk -20). The nanocomposite contact lens can be made from these acrylic oligomers or prepolymers by a thermal curing (100-120 °C) process. The porosity (e.g., 30-40 vol%) is created by a solvent / co-solvent system used to homogeneously mix two different types of polymers, followed by a washing / solvent exchange process to infuse water into the pores. The hydrogel and aqueous pores allow easy penetration of body fluids, while the flexible, low tack hydrophobic polymer is highly gas permeable.

[0052]

[0053]

[0054]

[0055] Table 1C shows some of the main monomers of commercial contact lenses. Table 1C is from "Dynamic Contact Angle Analysis of Silicone Hydrogel Contact Lenses," Michael Leonard Read, Philip Bruce Morgan, Jeremiah Michael Kelly, and Carole Maldonado-Godina, J Biomater Appl, March 10, 2010 online. The following additional abbreviations are used in Table 1C. mPDMS: mono-functional methacryloxypropyl terminated polydimethylsiloxane; EGDMA: ethylene glycol dimethacrylate; TEGDMA: tetraethylene glycol dimethacrylate; TPVC: tri-(trimethylsiloxy methylsilyl) propyl vinyl carbamate; NVA: N-vinylacetamide; PBVC: poly(dimethylsiloxy) di(silbutanol) bis(vinyl carbamate); M3U: alpha omega-bis(methacryloxyethyl imino carboxy ethyloxy propyl)-poly(dimethylsiloxane)-poly(trifluoropropylmethylsiloxane)-poly(methoxy-poly(ethylene glycol) propylmethyl-siloxane; FMM: alpha-methacryloxyethyl imino carboxy ethoxy propyl-poly(dimethylmethylsiloxy)-butyl dimethyl silane; TAIC: 1,3,5-triallyl-1,3,5-triazine-2,4,6(1H,3H,5H)-trione; IBM: isobornyl methacrylate; HOB: 2-hydroxybutyl methacrylate; NMNVA: N-methyl-N-vinylacetamide; AOE: 2-allyloxyethanol.

[0056]

[0057] In one embodiment, the hydrophilic component can be selected from monomers, oligomers, or prepolymers that contain at least one hydrophilic group of a hydroxyl, alkyl glycol, amine, lactam, carboxyl, or sulfonic acid group. The hydrophobic component can be composed of monomeric groups with a log P > 3. These immiscible components can be pre-mixed within a compatible solvent / co-solvent system, or pre-reacted into block copolymers prior to final curing into a uniform nanocomposite.

[0058] In one embodiment, the hydrophilic polymeric domains are fabricated to be ten times smaller than the average size of the hydrophobic domains, such that the hydrophobic domains are completely immersed in a continuous mixture of the hydrophilic domains and liquid water. This morphology of hydrophobic domains immersed in a hydrophilic continuum enhances the hydrophilicity, contact angle, and capillary wicking of the composite.

[0059] In one embodiment, the hydrophobic domains are in the shape of elongated needles, or are part of an interpenetrating network with a hydrophilic phase to enhance the oxygen permeability Dk beyond 100.

[0060] In another embodiment, the device surface comprises hydrophobic domains that are a fraction of the size of a cell (e.g., domain size between 10 nm to 1 micron or 10-100 nm) and are surrounded by hydrophilic domains and water pores that are even smaller in size (e.g., ten times smaller than the size of the submerged hydrophobic domain) to allow for enhanced cell adhesion to the device surface via hydrophobic bonding (enhanced accumulation in aqueous environment).

[0061] Figure 1

[0062] The present invention transports active functional ingredients to the host body environment with engineered precision and rate via porous nanocomposites. In this disclosure, we explore the transport of active ingredients (drugs) in such composites, where one class of domains with high drug affinity is used as its reservoir, while other domains (with significantly lower drug solubility) are utilized as barriers to modulate the drug release rate. The drug elution through the connected nanopores within such dual-domain composites is analogous to the molecular elution in a chromatographic column, where the drug retention time can be precisely modulated by its attraction to the affinity column components.

[0063] The transport (or release from) of drugs through porous hydrophobic-hydrophilic composites can be best modeled by first studying its permeation within each individual component, then effectively averaging over the entire composite (i.e., applying the effective medium theory of transport coefficients, see “Effective Medium Theory, Principles and Applications”, 2nd Edition, Tuck C. Choy, Oxford University Press, 2016). The drug permeability in each domain is generally determined by the product of its solubility (S) and diffusivity (D) in that domain, i.e., permeability = S x D. In the composites of interest, the water phase in the pores generally has the highest diffusion constant compared to the other two polymer domains (typically by several orders of magnitude), but this does not necessarily mean that its permeation constant is the highest, since the water solubility of the drug can be lower. Many drugs with significant hydrophobic segments can have the highest drug permeability at the interfacial region due to the higher affinity to the interfacial hydrophobic domain and the higher diffusion speed in the surrounding aqueous medium. Figure 2Drug with a charge head and a hydrophobic tail is shown to adhere strongly to the hydrophobic polymer pore surface and thus can have high partitioning and permeation constants at the interfacial region due to enhanced affinity. Drug is expelled from the porous nanocomposite through three modes: fluidic mode through the pore water (fastest), interfacial mode through the solid boundary layer followed by the water pore (medium), and solid mode through the bulk polymer. Figure 7 It is shown that although diffusion through the pore fluid is generally the fastest mode of transport, the significantly higher partitioning of the charged hydrophobic drug at the interface can make the interfacial drug release the most dominant mode, especially when the interface is attached with a modifier such as oleic acid with opposite charge.

[0064] Our modeling results are validated by the disclosed examples, which have shown that the most effective way to control drug delivery through porous nanocomposites is by increasing the partitioning coefficient of the drug and thus its loading accumulation at the polymer-pore interface region. Our innovative strategy is to make the interfacial release of the drug the most dominant mode (more than 80% of the total capacity) by including a boundary charge modifier at the polymer-water interface to greatly increase the partitioning coefficient of the drug in the boundary layer compared to that in the pore water or the solid polymer. The present invention enhances the delivery of the drug from the interfacial region of the porous nanocomposite.

[0065] By reducing the solid domain size to the nanoscale, the interfacial area of the porous composite can be scaled up to more than 100 m2 / gram. The present invention discloses a porous nanoporous composite with such high pore surface area that can be further prepared by implanting a boundary charged double layer through solvent loading or chemical addition of a boundary charge modifier at the polymer-water interface in the nanocomposite. The boundary charge migration can generate a Coulombic junction potential at the modified aqueous interface, such that the strong electrostatic attraction between the boundary charge and the oppositely charged drug enhances the partitioning, accumulation, and retention of the drug. As shown in the examples in the experimental section, such nanofabricated devices have long-range (1 / r 2 ) electrostatic forces between the drug molecules and the interfacial charge in the boundary region as well as all the nanoscale pores, enabling one of the most effective mechanisms to control the precision and duration of drug delivery through the nanoporous composite. In one embodiment, the anionic charge of the surface-implanted oleic acid prolongs the release of the cationic bupivacaine (BUP) by 50-fold. In another embodiment, the prolongation is 60-fold. Figure 8 and Composite morphology to facilitate control over drug delivery

[0066] Retention of charged drugs using electrostatic interactions at reservoir-pore interfaces

[0067] ​Based on matching the logP of the hydrophobic segment of the drug, the hydrophilic polymer (logP < 1) phase or the hydrophobic polymer (logP > 3) phase can be selected as the drug reservoir phase. The domain size of the drug reservoir polymer is preferably on the nanometer scale (10 nm - 100 nm) to maximize the domain-pore interface area and drug partitioning.

[0068] Increased drug loading (partitioning) at pore interfaces through charge interactions

[0069] Any compound with an end group that can exchange ions (including hydronium ions H30 + ) in aqueous media can be used as a boundary charge modifier. Compounds with pKa, Pkb, isoelectric point, and dissociable ions (e.g., basic ions, halide ions, quaternary ammonium ions, etc.) can be grafted onto the polymer-water boundary to create a charge layer by ion dissociation or hydronium ion adsorption. Polymer boundary modification can be achieved by chemical reaction with charge modifiers (e.g., reactive acrylic, amine, or imine moieties). Alternatively, boundary charge modifiers with higher affinity to the polymer can be grafted physically by the polymer solvation process as disclosed below. Such molecules are most stable in the case where the hydrophobic tail resides in the high affinity polymer reservoir and the charged head group is both immersed in the aqueous phase and surrounded by water molecules as well as counter ions. In the device of the present invention, the boundary charge modifier and the charged drug form a charged bilayer at the interface of the water pores and the hydrophobic / hydrophilic polymer.

[0070] Based on our modeling and experimental results, we have set general guidelines for selecting the appropriate boundary charge modifier (BCM) for each medical implant / contact device when delivering charged drugs over an extended period (more than a week), namely: (1) the water solubility of the physically grafted BCM must be low to minimize its leaching during the storage time of the device or over the extended drug delivery period; (2) the BCM is preferably neutral in physiological pH to avoid unintended electrostatic interactions with host body cells or other functional components outside the device upon leaching from the device; (3) the BCM is preferably metabolically inert or friendly to minimize the impact of accidental discharge or its accumulation over long periods of use.

[0071] Based on the above evaluation, we have chosen to use long chain (C 8-24Fatty acids as anions BCM in long-term release of cationic drugs from porous nanocomposite medical implant / contact devices. Organic fatty acids are biocompatible and generally have low toxicity. In fact, some omega-3 fatty acids (such as alpha-linolenic acid (ALA), eicosapentaenoic acid (EPA), and docosahexaenoic acid (DHA)) have important health benefits. However, our intent in using these compounds is to control the delivery of oppositely charged drugs, and our approach to loading them into the pores of the porous nanocomposite is designed to minimize their discharge in the body.

[0072] For anionic drugs, we need to include insoluble and biocompatible cationic BCM to the pore surface of the nanocomposite. In one embodiment, we choose sphingosine to extend the release of DFNa. The structure of sphingosine is similar to oleic acid, except with a primary amine head instead of a carboxylic acid. It is non-toxic and is a metabolic byproduct.

[0073] While all of these compounds are common surfactants for oil-water interfaces, their use as boundary charge modifiers to create a boundary charged double layer in the medical devices of the present invention would require a special treatment protocol as the inventors have discovered due to the low solubility of them in polymers and the low critical micelle concentration for micelle formation in water.

[0074] Common oil-water surfactants such as fatty acids or alkyl amines are not natural surfactants of polymers. The inventors have discovered that their packing density at the water-polymer boundary can be enhanced by physical implantation or chemical reaction (as opposed to at the water-oil boundary) in order to scale up their subsequent charge interaction with charged drugs. The physical implantation method is disclosed in one embodiment in the next section. The boundary charge modifier is physically carried into the polymer by infiltration, and the surface loading density is scaled up by solvent swelling of the host polymer.

[0075] Creation of charge double layers at the interface of aqueous pores and hydrophobic polymers

[0076] By reducing the drug reservoir domain size to the nanoscale, we scale up the interfacial area and drug packing therein dramatically. By including a charge double layer at the polymer-pore interface, we scale down the drug chemical potential dramatically by creating an oppositely charged interface and thus improve drug partitioning and loading in the interfacial region.

[0077] Figure 3

[0078] The present invention further provides a method of directly loading water-insoluble fatty acids into polymeric domains using a good solvent for the polymer, which also has a higher solubility for the fatty acid than water. A "good solvent" for a polymer is one in which the molecules have a good energetic interaction with the monomeric segments, such that the solvation effect in the polymer causes the polymer coils to swell and their imbibition causes the polymer to swell. The loading solvent is preferably a good solvent for the polymer, such that the drug loading can be accelerated and even supersaturation in the case of solvent swelling of the polymer. Such good solvents can load a significant amount of fatty acid molecules into the polymer domains, resulting in a significant increase in boundary modifier aggregation at the domain-pore interface. These PBCM fatty acid molecules migrate from the bulk polymer phase to the pore interface, where they can even be more stable in the case of anionic dissociation and formation of a charged double layer, especially in the presence of a cation other than hydronium. This charged double layer serves to enhance partitioning of cationic drugs and retain them for extended release. Examples

[0079] The present invention provides a method for making a nanocomposite comprising hydrophilic polymer domains, hydrophobic polymer domains, water pores, and a boundary charged double layer. The method comprises the steps of: (a) soaking a starting nanocomposite comprising hydrophilic polymer domains, hydrophobic polymer domains, and water pores in a solution comprising an appropriate amount of a boundary charge modifier dissolved in a good solvent that swells the hydrophilic polymer or the hydrophobic polymer, and (b) carrying the boundary charge modifier into the hydrophilic polymer domains or the hydrophobic polymer domains by a swelling process. The method optionally comprises the following steps after soaking: (c) removing excess solution from the surface of the nanocomposite, and (d) drying the nanocomposite. As an example, the hydrophobic polymer is silicone and the good solvent is ethanol. As an example, the starting nanocomposite can be a commercial contact lens, including those shown in Table 1C.

[0080] The present invention further provides a method for making a nanocomposite comprising hydrophilic polymer domains, water pores, and a boundary charged double layer. The method comprises the steps of: (a) soaking a starting nanocomposite comprising hydrophilic polymer domains and water pores in a solution comprising an appropriate amount of a boundary charge modifier dissolved in a good solvent that swells the hydrophilic polymer, and (b) carrying the boundary charge modifier into the hydrophilic polymer domains by a swelling process. The method optionally comprises the following steps after soaking: (c) removing excess solution from the surface of the nanocomposite, and (d) drying the nanocomposite. As an example, the hydrophilic polymer is HEMA and the good solvent is methanol or a mixture of water and ethanol. As an example, the starting nanocomposite can be a commercial contact lens, including those shown in Table 1C.​

[0081] The soaking step of both methods is performed for an appropriate period of time at an appropriate temperature. For example, the soaking is for 4 hours to 2 days, or 8 hours - 24 hours, or 16 hours - 24 hours, at 10°C - 40°C or 20°C - 30°C or at room temperature.

[0082] The following examples further illustrate the application. These examples are intended merely to be illustrative of the application and are not to be construed as limiting.

[0083] Oleic acid loading in commercial silicone hydrogel contact lenses

[0084] All contact lenses used in the following experiments (including the comparison with J&J clinical trial data) were sourced from VisionDirect BV (York Business Park, YO26 6RB, United Kingdom).

[0085] Figure 4

[0086] The following two commercial silicone hydrogel contact lenses were tested: ACUVUE® (46% water + 54% Narafilcon A) and ACUVUE® (38% water + 62% Senofilcon A). is a daily disposable contact lens, while is an approved extended wear contact lens. The lenses were rinsed with deionized water and then air dried before being soaked in 4 mL of 19 mg / mL, 27 mg / mL or 40 mg / mL oleic acid in ethanol. The duration of the soak was 24 hours at room temperature. After the loading step, the contact lenses were removed and excess fatty acid-ethanol solution was blotted off the lens surface. The lenses were washed in deionized water for 1 hour and then air dried overnight.

[0087] The loading of OA was determined by weighing the dry lenses before and after the fatty acid loading period, as Example 2 Drug loading in pristine lenses shown. The oleic acid loading was linearly related to the concentration of the loading solution. Furthermore, between concentrations of 15 mg / mL and 27 mg / mL, the affinity for oleic acid was higher For higher concentrations, the affinity for oleic acid was the same for both contact lenses. Therefore, the loading of oleic acid in the lenses can be adjusted by this correlation.

[0088] Release kinetics of the cationic drug tetracaine hydrochloride (THCL) from silicone hydrogel contact lenses loaded with oleic acid

[0089] The drug was loaded by soaking the lenses in the drug-PBS solution. The soaking duration for the pure lenses was for 24 hours at room temperature. After the loading period, the lenses were removed and excess drug solution on the surface was removed by blotting with filter paper.

[0090] Figure 5 Figure 6

[0091] Figure 5 and Figure 6 It is shown and Drug release of THCI in commercial lenses. The cumulative % release of drug was calculated by dividing the amount of drug released in a defined time by the total amount of drug loaded per lens. From Release kinetics of the cationic drug bupivacaine hydrochloride (BUP) from silicone hydrogel contact lenses loaded with oleic acid it is seen, The control contact lenses released 80% of THCI in less than 5 hours and about 100% in less than 1 day. After a release period of 100 hours, the lenses with 5.1% OA (oleic acid), 13.5% OA and 28.6% OA released about only 47%, 30% and 17% of THCI, respectively. The control contact lenses released 80% of THCI in less than 5 hours and about 100% in less than 1 day. After a release period of 100 hours, the lenses with 5.1% OA (oleic acid), 13.5% OA and 28.6% OA released about only 47%, 30% and 17% of THCI, respectively.

[0092] For lenses Figure 7 ), the control lenses released 100% of THCI in less than 1 day, as in the case of . After 100 hours, the lenses loaded with 9.1% OA released 49% of THCI, while 21.4% OA released 35%. At the release period of 400 hours, these percentages increased for 9.1% OA from 49% to 62%, and for 21.4% OA from 35% to 47%. and The presence of oleic acid in both contact lenses significantly reduced the burst release of the drug and prolonged the release kinetics of THCL. Moreover, the effect of the presence of oleic acid on the release kinetics was greater in than in .

[0093] Figure 8 Release kinetics of the cationic drug ketotifen fumarate (KTF) from silicone hydrogel contact lenses loaded with oleic acid

[0094] Figure 9 and Figure 10 It is shown and the release kinetics of the anesthetic drug BUP in The control lenses released 80% of the loaded BUP in about 5 hours. For the lenses loaded with oleic acid 13.5% OA lenses release 80% of the BUP within 300 hours. Release duration based on 80% BUP release, The presence of oleic acid can extend the delivery of BUP by 60-fold. For the case of lenses with 28.6% OA, only about 60% of the loaded BUP is released after 450 hours. For the case of 21.4% OA, the control lenses release 80% of the loaded BUP in about 2 hours. Lenses with 9.1% OA release 80% of the BUP in 100 hours. This indicates that oleic acid can extend the release of BUP by 50-fold in lenses with 21.4% OA. For the case of 28.6% OA, about 75% of the loaded BUP is released after 450 hours.

[0095] Example 6 Oleic acid (OA) loading in conventional hydrogel pHEMA contact lenses Example 7 Release kinetics of the cationic drug tetracaine hydrochloride (THCL) from one-day ACUVUE® contact lenses loaded with oleic acid

[0096] Figure 11 and Figure 12 The effect of oleic acid on the release kinetics of KTF, a relatively selective non-competitive antagonist of the histamine Hl receptor, is shown. The control lenses release 80% of the loaded KTF in less than 50 hours. For 5.1% OA, 13.5% OA and 28.6% OA, the lenses loaded with oleic acid release only 40%, 30% and 18% of the loaded KTF, respectively, after 100 hours. For the case of 21.4% OA, the control lenses also release more than 90% of the loaded KTF in less than 50 hours, as in the case of 28.6% OA. After 100 hours, 9.1% OA and 21.4% OA release 45% and 30% of the loaded KTF, respectively. For both 9.1% OA and 21.4% OA, the release kinetics are still extended after 1500 hours. In addition, the effect of oleic acid on the lenses is more pronounced than for the lenses.

[0097] Figure 11

[0098] One Day ACUVUE ​​​(58% water + 42% Etafilcon A) is a conventional hydrogel commercial lens composed primarily of poly-2-hydroxyethyl methacrylate (pHEMA). The lenses were rinsed with PBS and then air dried prior to use. The dry lenses were soaked in 4 mL of 100 mg / mL oleic acid in a mixture of ethanol and deionized water (75 / 25 ethanol / water). Since the swelling of the conventional hydrogel lenses in ethanol is negligible, it was decided to soak the lenses in a mixture of ethanol and deionized water since the lenses swell significantly in deionized water. After the loading step, the contact lenses were removed and the excess fatty acid-ethanol solution was blotted off the surface of the lenses. The lenses were washed in PBS for 1 hour and then air dried overnight. The loading of OA was determined by weighing the dry lenses before and after the fatty acid loading period.

[0099] We found the maximum oleic acid loading to be 2.9 wt.%. For example, when the oleic acid soak concentration was increased from 100 mg / mL to 500 mg / mL, the one-day ACUVUE® The oleic acid loading in the lenses remained at 2.9 wt.%.

[0100] Release kinetics of the anionic drug diclofenac sodium (DFNa) from silicone hydrogel contact lenses loaded with oleic acid Figure 13 Figure 14

[0101] For the one-day ACUVUE® contact lenses, a maximum oleic acid loading of 2.9 wt.% was observed. Conventional hydrogel contact lenses such as the one-day ACUVUE® are not composed of hydrophobic domains, but rather are composed primarily of pHEMA hydrophilic polymers. As a result, these lenses have a higher water content than silicone hydrogel contact lenses, and the lack of hydrophobic silicone domains reduces oxygen permeability and absorption of hydrophobic molecules such as oleic acid. Figure 12 and Figure 14 The effect of oleic acid on the release kinetics of KTF and THCL in the one-day ACUVUE® contact lenses was described. The control lenses released 90% of the THCL in less than 24 hours. For the lenses loaded with 2.9% OA, 60% of the THCL was released after 250 hours. For KTF Release kinetics of the anionic drug flurbiprofen sodium (FBNa) from silicone hydrogel contact lenses loaded with oleic acid ), the control lenses released 80% of the drug in about 48 hours. For the case of the lenses loaded with 2.9% OA, 70% of the KTF was released after 500 hours. Thus, even though the amount of oleic acid loaded in the lenses was limited to 2.9 wt.%, the release prolongation achieved for THCL and KTF was still significant. It should be noted that since the lenses could not take up higher amounts of oleic acid, we believe that the drug release prolongation achieved at 2.9% OA is likely the maximum prolongation that can be achieved for the one-day ACUVUE® contact lenses.

[0102] Figure 15 Figure 16

[0103] Diclofenac sodium is one of the nonsteroidal anti-inflammatory drugs (NSAIDs) that carry a negative charge at physiological pH. Figure 15 and Example 10 Comparison of release kinetics of KTF from ACUVUE® contact lenses loaded with three different fatty acids - oleic acid, linoleic acid, or alpha-linolenic acid It shows the effect of oleic acid on and The effect of DFNa release from contact lenses. (The text then lists DFNa release parameters, which are not directly related to the preceding sentence and are omitted from the translation.) lens( Figure 17 It releases 80% of DFNa in approximately 5 hours, while The control lens released 80% within 60 hours. For ( Figure 17 Lenses containing oleic acid released 80% of DFNa within 4 hours, while control lenses released 80% of DFNa within 20 hours. Therefore, DFNa is released faster in oleic acid-loaded contact lenses compared to lenses without oleic acid. Based on the 80% drug release time, DFNa is released more rapidly in oleic acid-loaded contact lenses compared to control lenses. The release from the lens is 12 times faster. For Compared to the control lens, DFNa is released 5 times faster in lenses containing oleic acid.

[0104] Example 11 Release kinetics of the anionic drug (DFNa) from silicone hydrogel contact lenses loaded with sphingosine Figure 18

[0105] Flurbiprofen sodium (FBNa) is a nonsteroidal anti-inflammatory drug used after cataract surgery. Figure 18 and Example 12 Comparison of release kinetics of ketotifen fumarate from ACUVUE® contact lenses loaded with two different saturated fatty acids - myristic acid (tetradecanoic acid) and stearic acid (octadecanoic acid) The effect of oleic acid on FBNa release is shown. Figure 19 Among them, those with OA of 13.5% and 28.6% The lens released 80% of FBNa within 2 hours, while the control lens released 80% within 12 hours. For In this case, oleic acid-loaded lenses released 80% of FBNa within 2 hours, while control lenses released 80% within 6 hours. Similar to DFNa, FBNa was released more rapidly in oleic acid-loaded contact lenses. Based on the 80% drug release time, for oleic acid-loaded lenses... and The release of FBNa was 6 times and 3 times faster than the control, respectively.

[0106] From the results obtained for DFNa and FBNa, it can be concluded that the presence of the anionic oleic acid in the silicone hydrogel contact lenses accelerates the drug release of the anionic drugs due to the repulsive ionic interactions. Furthermore, it was demonstrated that the effect of oleic acid on accelerating the release kinetics of DFNa and FBNa was more pronounced than for the control.

[0107] Example 13 Oleic acid (OA) loading in pristine silicone hydrogel contact lenses Example 14 Oleic acid (OA) loading in pristine conventional hydrogel pHEMA contact lenses Example 15 Oleic acid (OA) loading in pristine conventional hydrogel pHEMA contact lenses

[0108] Oleic acid is a long-chain unsaturated fatty acid with one double bond. Linoleic acid and a-linolenic acid are unsaturated fatty acids with the same carbon number as oleic acid but with two and three double bonds, respectively. These three fatty acids are oily liquids at room temperature. However, due to the different number of double bonds, we hypothesized that their packing density at the aqueous interface of the hydrophobic domain would affect the drug release kinetics.

[0109] We tested 20 mg / mL fatty acid ethanol soak concentrations in ACUVUE Oasys® contact lenses, as Example 16 High precision control of KTF delivery dose and time outlined. At the same fatty acid soak concentrations, we found similar fatty acid loadings, i.e., 6.1 wt% for oleic acid, 6.2 wt% for linoleic acid, and 6.0 wt% for a-linolenic acid. From Figure 20 we see that for oleic acid, linoleic acid, and a-linolenic acid, 37%, 44%, and 53% of KTF was released after a 100-hour release period, respectively. Interestingly, the differences in release kinetics of the three fatty acids became more pronounced after a 200-hour release period. As the number of cis double bonds increased, the twisting and bending of the chains also increased, disrupting their packing in the monolayer at the interface. The differences in charge density of the three fatty acids at the aqueous interface can explain the differences in KTF release kinetics, with the longest release time achieved for oleic acid. However, all fatty acid loaded lenses significantly prolonged KTF compared to the control.

[0110] Figure 21 Figure 17

[0111] ACUVUE Oasys® commercial lenses were rinsed with deionized water and then air dried before use. ACUVUE Oasys® contact lenses were soaked in 4 mL of 6.5 mg / mL sphingosine ethanol. The duration of the soak was 24 hours at room temperature. After the loading step, the contact lenses were removed and excess sphingosine-ethanol solution was blotted off the lens surface. The lenses were washed in deionized water for 1 hour and then air dried overnight.

[0112] Diclofenac sodium was loaded onto lenses by immersing them in 5 mL of a 0.2 mg / mL drug-PBS solution. The immersion period for the pure lenses was 24 hours at room temperature. After the loading period, the lenses were removed and excess drug solution was blotted off with filter paper. To determine the amount of drug loaded onto each lens, the drug concentration in each immersion solution was measured using a UV-Vis spectrophotometer (Varian Can's 50 Bio, Walnut Creek, CA, USA) before and after the immersion period. Following the drug loading step, an in vitro release assay was used to test the lenses.

[0113] Figure 10 The effect of sphingosine on DFNa release is shown. Figure 20 Among them, a 6.5 mg / mL sphingosine-loaded solution The lens released only 40% of DFNa within 100 hours, while the control lens released 80% within 12 hours. Sphingosine also increased DFNa uptake in the control lens from 187.1 ± 16.1 μg to 590.0 ± 0.7 μg.

[0114] Figure 21 Figure 1 ​

[0115] ACUVUE (Refractive power -3.5) Contact lenses were rinsed with deionized water and then air-dried overnight before use. One batch of contact lenses was immersed in 4 mL of 7.5 mg / mL myristic acid ethanol. Another batch of contact lenses was immersed in 4 mL of 7.5 mg / mL stearic acid ethanol. A third batch of lenses was used as a control (no fatty acid loading). The immersion time was 24 hours at 23°C. After the loading step, the contact lenses were removed and excess fatty acid-ethanol solution was aspirated from the lens surface. The lenses were washed in deionized water for 1 hour and then air-dried overnight. All contact lenses were optically clear in the hydrated state. The percentage by weight of fatty acids retained in the contact lenses was determined by recording the increase in weight of the dried lenses after the immersion and drying processes. For both stearic acid and myristic acid, approximately 1.5% ± 0.5% by weight of the dried lenses consisted of fatty acids.

[0116] Ketotifen fumarate (KTF) was loaded by soaking the lenses in 5 mL of a 0.3 mg / mL KTF PBS solution. After the loading period (24 hours at 23 °C), the lenses were removed and excess drug solution on the surface was removed by blotting with filter paper. To determine the amount of KTF loaded in each lens, the KTF concentration of each soaking solution was measured before and after the soaking period using a UV-visible spectrophotometer (Varian Cary 50 Bio, Walnut Creek, CA, USA). After the drug loading step, the lenses were tested using an in vitro release experiment. The drug release experiment was performed by soaking the drug loaded lenses in 3 mL of Dulbecco PBS, pH 7.4 and temperature 23 °C. During the release experiment, 1 mL of release sample was removed at predetermined time intervals and 1 mL of fresh PBS was refilled into the release medium. The amount of drug released was measured using a UV-spectrophotometer (Varian Cary 50 Bio) at a wavelength of 300 nm. The drug release experiment was performed in triplicate for each different case. Thus, the KTF uptake for the control lenses was determined to be 176 pg, while the uptake for the stearic acid loaded lenses was slightly increased to 214 pg. In contrast, the contact lenses with myristic acid exhibited a significantly increased KTF uptake of 516 pg.

[0117] ​ The effect of stearic acid and myristic acid on the release kinetics of KTF from the ACUVUE® contact lenses is shown. Here, we define the term "release duration" as the time to 70% cumulative drug release. The release duration for the control lenses (no fatty acid) was 5 hours, while the release duration for the lenses loaded with stearic acid and myristic acid was 6 hours and 60 hours, respectively. The release duration of between 8 and 24 hours is consistent with the daily wear schedule of the lenses.

[0118] ​

[0119] In this example, the medical device is the commercial silicone hydrogel ACUVUE® contact lenses (Johnson & Johnson Vision Care, Inc., Jacksonville, Fla.) and the fatty acid PBCM is oleic acid. The lenses are composed of HEMA hydrophilic polymer, silicone hydrophobic polymer, and the water content in the aqueous pores is 38%. Ethanol is a good solvent that is able to swell the silicone phase. The silicone hydrogel commercial lenses were rinsed with deionized water and then air dried before use. The ACUVUE® Contact lenses were soaked in 4 mL of 19 mg / mL or 33 mg / mL oleic acid ethanol. ACUVUE Contact lenses were soaked in 4 mL of 19 mg / mL, 27 mg / mL, or 40 mg / mL oleic acid ethanol. The duration of the soak was 24 hours at room temperature. After the loading step, the contact lenses were removed and excess fatty acid-ethanol solution was blotted off the lens surface, and the lenses were air dried overnight. The loading of OA was determined by weighing the dry lenses before and after the fatty acid loading period.

[0120] ​

[0121] In this example, the medical device is a HEMA daily ACUVUE contact lens, and the fatty acid PBCM is oleic acid. The daily ACUVUE is a conventional hydrogel commercial lens consisting primarily of poly-2-hydroxyethyl methacrylate (pHEMA), which is hydrophilic but more hydrophobic than aqueous pores (water content of 58%). Ethanol does not swell the HEMA lens, but methanol is a good solvent for the HEMA phase. The lenses were rinsed with PBS and air dried overnight before use. The dry lenses were soaked in 4 mL of 100-300 mg / mL oleic acid methanol. The duration of the soak was maintained at room temperature for 24 hours. After the soak, excess fatty acid-methanol solution was blotted off the lens surface, and the lenses were immersed in PBS for 1 hour to wash away the remaining methanol. After the wash, the lenses were dried overnight.

[0122] ​

[0123] In this example, the medical device is a HEMA daily ACUVUE contact lens, and the fatty acid PBCM is oleic acid. Methanol is a good swelling solvent, but is not suitable for processing because small amounts of residual can be harmful to the eye. Since the swelling of conventional hydrogel lenses in ethanol is negligible, we alternatively developed a water / ethanol solvent mixture that is capable of swelling the HEMA phase. The lenses were rinsed with PBS and then air dried before use. The dry lenses were soaked in 4 mL of 100 mg / mL oleic acid in a mixture of ethanol and deionized water (75 / 25 ethanol / water). After the loading step, the contact lenses were removed and excess fatty acid-ethanol solution was blotted off the lens surface. The lenses were washed in PBS for 1 hour and then air dried overnight. The loading of OA was determined by weighing the dry lenses before and after the fatty acid loading period.

[0124] ​

[0125] This example demonstrates the high precision of the oleic acid boundary charge control technology. We first demonstrated that the total KTF load in oleic acid loaded lenses can be precisely determined by the drug concentration during solution loading.

[0126] The lens brand is ACUVUE TruEye and loaded with oleic acid 2-4% (calculated as a function of dry lens weight change, which can fluctuate with drying conditions) using 15 mg / mL oleic acid ethanol by the same ethanol soak process described in Example 1. KTF was loaded according to the process in Example 2. ​ The linear correlation of KTF load range 70-375 pg / CL with solution loading concentration is shown.

[0127] In addition, we verified that the kinetic release profile is the same over a wide load range of 70-375 pg / CL KTF per contact lens (oleic acid load condition is 15 mg / ml in ethanol) with a fixed amount of oleic acid loading ​ ).

[0128] Our data in Example 10 also indicates that the same KTF release profile can be easily extended to more than 600 hours ​ ) with oleic acid loading adjusted to 6%, and further extended to 1500 hours (Example 5, ​ ) with oleic acid loading increased to 9.1%. By studying the optimal load combination of oleic acid and KTF drug %, we expect to be able to manufacture a silicone-hydrogel lens that precisely delivers 80% of the KTF effective load at a lower average dose in exactly one week.

[0129] Based on the consistent trend shown in ​ and ​ regarding the correlation of cumulative release rate percentage with oleic acid load in CL over a wide range, we expect to be able to specify the precise formulation of drug and oleic acid load weight in a commercial contact lens to deliver more than 80% of the KTF embedded amount (e.g., 80 pg) in the desired duration (e.g., 5 days) to achieve high precision drug delivery control.

[0130] ACUVUE Theravision by Johnson & Johnson is the first approved product (Japan) for delivering the drug ketotifen (KTF) in a daily disposable hydrogel lens. The in vitro experimental results of J&J etafilcon A-KTF release contact lenses show that almost 90% of the drug is released in less than 2 hours (see US Patent No. 9,962,376, column 1 line 58, ​). By comparing the results of Examples 6, 10 and 16 with the J&J clinical trial data disclosed in the '376 patent as described above, the oleic acid loaded contact lenses of the present application are capable of: (a) more precisely controlling the daily delivered dose of KTF within the wide range of 1-375 micrograms, (b) extending the duration of delivery of this precise dose from hours to weeks (600 hours), significantly longer than the burst release shown in the J&J clinical data of the '376 patent, (c) achieving the same precise controlled KTF delivery for both a HEMA daily disposable hydrogel lens (see Example 6) which was the only lens in the J&J clinical trial, and a more functional long-wear silicone-hydrogel lens (see Examples 10 and 16).

[0131] The present application, together with its objects and the particular parts thereof, can now be understood from the following detailed description taken in connection with the accompanying drawings, in which: CLAIM

Claims

1. A contact lens comprising a nanocomposite and a drug, wherein the nanocomposite comprises: (a) a hydrophilic polymer domain formed from 2-hydroxyethyl methacrylate (HEMA), N,N-dimethylacrylamide (DMA), N-vinyl-2-pyrrolidone (NVP), 4,4-dimethyl-2-vinyl-2-oxazolin-5-one, methacrylic acid (MAA), N-(hydroxymethyl)acrylamide, N-[3-(dimethylamino)propyl]methacrylamide, or ethyleneglycol dimethacrylate monomers, (b) a hydrophobic polymer domain formed from 3-[tris(trimethylsilyloxy)silyl]propyl methacrylate (TRIS), 3-methacryloyloxy-2-hydroxypropyloxy(propyl bis(trimethylsilyl)oxy)methylsilane (SIGMA), dimethylsiloxane, monomethylsiloxane, fluorosiloxane, or methyl methacrylate (MMA) monomers, (c) water pores, and (d) a boundary charged double layer; wherein the boundary charged double layer comprises: (i) a charge of a head group of a boundary charge modifier, and (ii) a charge of the drug that is opposite to the charge of the head group, when the drug is positively charged, the boundary charged modifier is oleic acid, linoleic acid, a-linolenic acid, or myristic acid; when the drug is negatively charged, the boundary charged modifier is sphingosine, wherein at least 80% of the drug in the boundary charged double layer is partitioned at a boundary interface of the hydrophilic polymer domain and the water pores when the drug is hydrophilic, and at least 80% of the drug in the boundary charged double layer is partitioned at a boundary interface of the hydrophobic polymer domain and the water pores when the drug is hydrophobic.

2. A contact lens comprising a nanocomposite and a drug, wherein the nanocomposite comprises a hydrophilic polymer domain, water pores, and a boundary charged double layer, wherein the hydrophilic polymer domain is formed from 2-hydroxyethyl methacrylate (HEMA), N,N-dimethylacrylamide (DMA), N-vinyl-2-pyrrolidone (NVP), 4,4-dimethyl-2-vinyl-2-oxazolin-5-one, methacrylic acid (MAA), N-(hydroxymethyl)acrylamide, N-[3-(dimethylamino)propyl]methacrylamide, or ethyleneglycol dimethacrylate monomers, wherein the boundary charged double layer comprises: (i) a charge of a head group of a boundary charge modifier, and (ii) a charge of the drug that is opposite to the charge of the head group, when the drug is positively charged, the boundary charged modifier is oleic acid, linoleic acid, a-linolenic acid, or myristic acid; when the drug is negatively charged, the boundary charged modifier is sphingosine, wherein at least 80% of the drug in the boundary charged double layer is partitioned at a boundary interface of the hydrophilic polymer domain and the water pores.

3. The contact lens of claim 1 or 2, wherein the boundary charge modifier is a molecule having a charged head group and a hydrophobic tail and is immobilized at the boundary charged double layer throughout the useful life of the contact lens.

4. The contact lens of claim 1, wherein the water pores are 10 to 40 weight percent and the size of each of the polymeric domains is between 10 to 100 nanometers.

5. A method for making the contact lens of claim 1, comprising the steps of: immersing a starting nanocomposite comprising hydrophilic polymeric domains, hydrophobic polymeric domains, and water pores in a solution comprising a boundary charge modifier dissolved in a good solvent that swells the hydrophilic polymer or the hydrophobic polymer; and carrying the boundary charge modifier into the hydrophilic polymeric domains or the hydrophobic polymeric domains by a swelling process.

6. A method for making the contact lens of claim 2, comprising the steps of: immersing a starting nanocomposite comprising hydrophilic polymeric domains and water pores in a solution comprising a boundary charge modifier dissolved in a good solvent that swells the hydrophilic polymer; and carrying the boundary charge modifier into the hydrophilic polymeric domains by a swelling process.

7. The method of claim 5, wherein the hydrophobic polymer is silicone and the good solvent is ethanol.

8. The method of claim 5 or 6, wherein the hydrophilic polymer is HEMA and the good solvent is methanol.

9. The method of claim 5 or 6, wherein the hydrophilic polymer is HEMA and the good solvent is a mixture of water and ethanol.

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