Silicone hydrogel medical device for delivery of hyaluronic acid

Abstract

The present application is directed to a device that delivers high amounts of a biopolymer, hyaluronic acid (HA), during wear of a contact lens. The device is a delivery contact lens comprising HA as the main active ingredient and silicone-hydrogel contact lenses as a carrier. The device comprises: (a) a silicone hydrogel soft contact lens comprising 20-50% w / w of a hydrophobic silicone polymer domain, 20-50% w / w of a hydrophilic polymer domain, and 30 – 60 % w / w of water; (b) a boundary charge modifier in the amount of 0.5 – 3.0% w / w of the dried contact lens of (a), wherein the boundary charge modifier consists of a cationic head and a carbon chain length of 12 – 24, attached at the boundary surfaces of the water and the polymer domains of the contact lens; and (c) HA of molecular weight from 3 k to 3,500 k Dalton, adsorbed to the cationic head of the boundary charge modifier. The device provides a sufficient uptake and prolonged release of hyaluronic acid over 24 hours to meet with the intended efficacy and comfort of a daily disposable contact lens.

Description

[0001] SILICONE HYDROGEL MEDICAL DEVICE FOR DELIVERY OF HYALURONIC ACID

[0002] FIELD OF THE INVENTION

[0003] The present invention relates to medical devices, such as but not limited to soft contact lenses for the delivery of hyaluronic acid. The device comprises a prescribed amount of hyaluronic acid loaded by high-affinity adsorption to a silicone hydrogel soft contact lens comprising aqueous pores, one or more hydrophilic domains, and one or more hydrophobic domains, and a boundary charge modifier. When placed in contact with an ocular tear environment, the device releases substantially higher amounts of hyaluronic acid (at least 5 times greater) than a comparative, unmodified silicone hydrogel soft contact lens and with a controlled release kinetic profile over 24 hours for daily disposable use. This invention describes the preparation of a silicone hydrogel contact lens implanted with a boundary charge surface modifier for substantially increasing the lens’ affinity and uptake adsorption of hyaluronic acid and the subsequent sustained release of hyaluronic acid triggered by contacting with the tear fluid of physiological ionic strength.

[0004] BACKGROUND OF THE INVENTION

[0005] The topical application of hyaluronic acid (HA) eye drops can provide an elevated level of lubrication on the ocular surface and relieve symptoms of dry eye disease. Its ability to retain tear water and form a protective, hydrating layer on top of the muco-aqueous layer of the tear film helps to stabilize the tear film and improve overall eye comfort. Hyaluronic acid (HA) refers to sodium hyaluronate with molecular weight in the range of ~3 kDa to 4,000 kDa. It has been reported that hyaluronic acid above a molecular weight of 1,000 kDa is an anti-inflammatory mediator and more effective for relief of symptoms dry eye as shown by Medic et al. (2024). However, eye drops have diminished effectiveness due to rapid clearance from the ocular surface and hence require frequent instillation, which is inconvenient.

[0006] Drug delivery from contact lenses, compared with eyedrops, significantly improves a therapeutic drug’s retention on ocular surface, enhancing bioavailability and therapeutic efficacy.

[0007] Silicone-hydrogel contact lenses (SiHy) are known for their high oxygen permeability, which helps support eye health and comfort, especially during extended wear. However, the

[0008] - 1 -

[0009] 181727421.1 hydrophobic silicone domain in the SiHy contact lens, despite its tremendous benefits in oxygen permeation, does disturb tear homeostasis by adsorbing hydrophobic lipids and proteins, leading to more tear film breakups and evaporation loss. Wearers of silicone hydrogel lenses commonly experience issues such as dryness, discomfort, and lens-related irritation. Approximately 30 % to 50 % of contact lens wearers complain of dry eye symptoms and significantly increases the odds of dropping out of contact lens use. A silicone hydrogel lens with controlled delivery of sufficient hyaluronic acid would enhance the tear film stability, ocular lubrication, and overall eye comfort, potentially reducing most dry eye syndrome and improving contact lens retention rates.

[0010] There exists a need for a device and a method that improves both the uptake quantity and release kinetics of polysaccharides such as hyaluronic acid without impacting the lens physical and optical performances.

[0011] BRIEF DESCRIPTION OF THE DRAWING

[0012] FIG. 1 shows one embodiment of the present device. The incorporation of cationic charged ligand groups at the boundary surfaces of polymer-aqueous pores increases the populations of anionic HA macromolecules in proportion to charge densities and thus, enhances the delivery amounts and durations of HA release from the device.

[0013] FIG. 2 depicts the cumulative release of hyaluronic acid for Senofilcon A, and Otufilcon A silicone hydrogel lenses with and without the cetalkonium chloride (CKC) modification.

[0014] FIG. 3 depicts the cumulative release of hyaluronic acid for Senofilcon A, and Otufilcon A silicone hydrogel lenses with and without the CKC modification and the application of steam sterilization by autoclaving (aut).

[0015] FIG. 4 depicts the cumulative hyaluronic acid release and sphingosine soaking concentration for Otufilcon A silicone hydrogel lenses.

[0016] FIG. 5 further depicts the cumulative hyaluronic acid release as a function of sphingosine soaking concentration for Otufilcon A silicone hydrogel lenses.

[0017] FIG. 6 depicts the cumulative hyaluronic acid release as a function of CKC concentration for Otufilcon A silicone hydrogel lenses.

[0018] FIG. 7 depicts the cumulative hyaluronic acid release and the effect of application of steam sterilization by autoclaving (aut) for Otufilcon A silicone hydrogel lenses loaded with CKC.

[0019] - 2 -

[0020] 181727421.1 FIG. 8 depicts the cumulative hyaluronic acid release as a function of CKC loading at different ethanol concentrations for Otufilcon A silicone hydrogel lenses.

[0021] FIG. 9 depicts the cumulative hyaluronic acid release and the effect application of steam sterilization by autoclaving (aut) with sphingosine and phytosphingosine presence for Otufilcon A silicone hydrogel lenses.

[0022] FIG. 10 depicts the cumulative low molecular weight (3 kDa) hyaluronic acid release amount with CKC presence for Otufilcon A lenses.

[0023] FIG. 11 depicts the cumulative hyaluronic acid release and ionic strength of hyaluronic acid loading solution for Otufilcon A silicone hydrogel lenses loaded with CKC.

[0024] FIG. 12 depicts the cumulative hyaluronic acid release as a function of the pH of hyaluronic acid loading solution for Otufilcon A silicone hydrogel lenses loaded with CKC.

[0025] FIG. 13 depicts the cumulative hyaluronic acid release as a function of concentration of hyaluronic acid loading solution for Otufilcon A silicone hydrogel lenses loaded with CKC.

[0026] FIG. 14 depicts the cumulative hyaluronic acid release as a function pH of hyaluronic acid loading solution for Otufilcon A silicone hydrogel lenses that are unmodified (No CKC) or loaded with CKC.

[0027] DETAILED DESCRIPTION OF THE INVENTION

[0028] Definitions

[0029] A “drug” or “pharmaceutical compound” refers to a small molecule or macromolecule that has efficacy to cause a physiological change in a subject for a function of curing ocular diseases or improving health.

[0030] “Delivery contact lens” (DCL), as used herein, refers to a transparent, soft lens device that is in contact with the tear film, on top of the cornea, with the goal of delivering one or more drugs into the ocular environment.

[0031] “Silicone hydrogel contact lens” (SiHy) refers to a soft contact lens made primarily from silicone macromers which offer better oxygen permeability, longer wear time, and improved durability compared to traditional hydrogel lenses. These lenses have both hydrophobic and hydrophilic bulk domains which have implications in the lens performance, comfort, and biocompatibility.

[0032] A “charged drug” refers to a drug that is either cationic or anionic under normal physiological conditions.

[0033] - 3 -

[0034] 181727421.1 A “boundary charged double layer”, as used herein, consists of a pair of a negatively charged layer and a positively charged layer. This double layer is formed at the boundary interface of the water-polymer domains within a silicone hydrogel contact lens.

[0035] A “boundary charge modifier” is a compound implanted onto a water-polymer interface, either by physical dissolving or chemical bonding. It is responsible for creating a boundary charged double layer on the water-polymer domain interfaces.

[0036] Delivery Contact Lens Device

[0037] The present application is directed to a device that delivers high amounts of a biopolymer hyaluronic acid (HA) during wear of a contact lens. The device is a delivery contact lens (DCL) comprising HA as the main active ingredient and silicone-hydrogel contact lenses (SiHy) as a retainer. The device is capable of achieving a sufficient uptake and sustained release of hyaluronic acid over a 24 h period, aligning with the intended use timeframe of a daily disposable contact lens.

[0038] The device comprises: (a) a silicone hydrogel soft contact lens comprising 20-50% w / w of a hydrophobic silicone polymer domain, 20-50% w / w of a hydrophilic polymer domain, and 30 - 60 % w / w of water; (b) a boundary charge modifier in the amount of 0.5 - 3.0% w / w of the dried contact lens of (a), wherein the boundary charge modifier consists of a cationic head and a carbon chain length of 12 - 24, attached at the boundary surfaces of the water and the polymer domains of the contact lens; and (c) HA of molecular weight from 3 k to 3,500 k Dalton, adsorbed to the cationic head of the boundary charge modifier.

[0039] In the present device, the hydrophobic polymer domain components may be one or more materials selected from the group consisting of: [tris(trimethylsiloxy)silyl]propyl methacrylate (“TRIS”); 3-methacryloxy-2-hydroxypropoxy (propylbis(trimethylsilyloxy) methylsilane (“SIGMA”); fluorosiloxane macromer; mono-(3-methacryloxy-2- hydroxypropyloxy) propyl terminated, mono-butyl terminated polydimethylsiloxane; and mono-methacryloxypropyl terminated polydimethylsiloxane.

[0040] In the present device, the hydrophilic polymer domain components may be one or more materials selected from the group consisting of: 2-hydroxyethyl methacrylate, N,N- dimethylacrylamide, N-vinyl-2-pyrrolidone, N-vinyl-N-methylacetamide, 1 ,4-butanediol vinyl ether, 4,4-dimethyl-2-vinyl-2-oxazolin-5-one, methacrylic acid, 2- (methacryloyloxyethyl)-2-(trimethylammonioethyl) phosphate, ethylene glycol dimethacrylate, poly(N-vinyl pyrrolidone) (“PVP”), triethyleneglycol dimethacrylate, and ethoxyethyl methacrylamide.

[0041] - 4 -

[0042] 181727421.1 The hydrophilic component raises the device’s water affinity, while the hydrophobic component enhances the device’s oxygen permeability (Dk / t > 100).

[0043] Non-limiting examples of silicone hydrogels that may be used in the present device include Balafilcon A, Comfilcon A, Delefilcon A, Enfilcon A, Fanfilcon A, Galyfilcon A, Kalifilcon A, Lehfilcon A, Lotrafilcon A, Lotrafilcon B, Narafilcon A, Olifilcon A, Olifilcon B, Samfilcon A, Senofilcon A, Senofilcon C, Sifilcon A, Somofilcon A, Stenfilcon A, Otufilcon A, and Verofilcon A.

[0044] Based on experimental results, the inventors set a general guideline for choosing an appropriate boundary charge modifier for a DCL in the delivery of HA. First, the aqueous solubility of the implanted boundary charge modifier must be low (<1.0 - 2.0 mg / mE at 25 °C) to minimize its leaching during a device’s storage time in packaging solution or when in placement in the eye for use. Second, the boundary charge modifier is preferably to be metabolically inert, or not cytotoxic to minimize impacts from possible accidental release in the storage solution or during extended use.

[0045] In one embodiment, the boundary charged modifier is a surface-active agent imposing electrostatic interactions. One preferred surface-active agent is an alkyl amine, or a cationic quaternary ammonium salt, all having a carbon chain length of 8 - 24, preferably 12 - 18.

[0046] Another preferred surface- active agent as a boundary charge modifier is a sphingolipid. Sphingolipids are a class of lipids that contain a sphingoid base and are key components of cell membranes. They play crucial roles in maintaining membrane structure, cellular signaling, and apoptosis. Sphingolipids include ceramides, sphingosine, phytosphingosine, sphingomyelins, and glycosphingolipids, each contributing to diverse biological functions across different cell types.

[0047] In one embodiment, the boundary charged modifier is a quaternary ammonium salt, a sphingolipid, or a cationic lipid

[0048] In another embodiment, the boundary charged modifier is cetalkonium chloride, sphingosine, or phytosphingosine.

[0049] HA of medical or pharmaceutical grade manufactured under GMP standards is preferred. In one embodiment, the HA has a molecular weight of 3-50 kDa, 50 to 1,000 kDa, 1,000 to 3,500 kDa, or a combination thereof, most preferably in the range of 500 to 2,500 kDa. The molecular weight of HA can be designed to provide a broad range of preferred daily dosage as well as kinetic release profiles.

[0050] In ophthalmology, high MW hyaluronic acid (over 500 - 1000 kDa) is more effective in artificial tears and other eye drops due to their unique viscoelastic and hydrating powers. In

[0051] - 5 -

[0052] 181727421.1 aqueous medium, HA has a high viscosity and numerous glucosamine groups that facilitate their anchoring into the mucin layer underneath topical tear film, which not only is highly effective in retaining tear water (by reducing water molecules’ chemical potential) but also provides a lubricated boundary for comfortably blinking and additionally protecting of the cornea. These are some of the known benefits of HA’s relieving the discomfort associated with dry eye disease. Due to its longer residence time in mucin-tear layer and complex supramolecular structure, HA can reduce ocular inflammation, which is a root cause of the multifactorial pathophysiology of dry eye disease.

[0053] The present device incorporates cationic boundary charge modifiers at the polymer- aqueous pore interfaces and forms charged double layers at the polymer-pore interfacial zone (see FIG. 1). Each of the boundary charged double layers is formed from (i) the charge of a head group of a boundary charge modifier, and (ii) the charge of HA or a counter ion to the charged head group, 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 use life of the device. HA is composed of repeating units of P-l,4-glucuronic acid and P-l,3-N-acetyl-D-glucosamine. The carboxyl group on the glucuronic acid residues ionizes in physiological conditions, resulting in a negative charge on the molecule. The anionic nature of HA allows it to interact with and stabilize cationic entities such as calcium, magnesium and certain peptide ions.

[0054] The cationic boundary charge modifiers enrich adsorption of high molecular HA to the interfacial layers and consequently enhanced the HA delivery amount and duration 5 to 50 times more than the control without boundary charge modifiers.

[0055] The total amount of high molecular weight HA released from a DCL is between a desired therapeutical range of 5 and 50 pg per lens daily. The composition and morphological feature (such as the hydrophobic / hydrophilic ratio, weight % of boundary charge modifier, and weight % of water pores) can be tailored in each application based on the required specifications of the contact fluid wetting, oxygen exchange rate, cell adhesion and growth, as well as in the desired HA release kinetic profile.

[0056] The present device is effective to deliver HA to the ocular environment with the following advantages in the following aspects: (i) The burst release in the first hour can be reduced to less than 30 - 50% of the total delivery due to the higher partition of the HA molecules in the boundary7charged double layer; ii) The total capacity' of HA loading per device can be precisely tuned by adjusting a device’s aqueous porosity, total surface area, surface charge density, and interfacial double charge layer thickness so that the HA uptake

[0057] - 6 -

[0058] 181727421.1 capacity and kinetics are precise and consistent within the intended device duration time; iii) Due to the strength of the electrostatic interactions between the boundary charge modifier with HA, a fraction of loaded HA is expected to remain adsorbed at the double layer for the intended duration of the device. Being adsorbed in a contact lens, this fraction of HA would help to maintain a high level of moisture on the lens surface, keeping the users’ eyes hydrated, reducing dryness of the eye and preventing hydrophobic deposits in lens fouling. The moisture-retaining ability of HA would also help to maintain prolonged comfort since it would prevent the lenses from drying out, especially in dry environments or during prolonged wear.

[0059] The device of the present invention further provides high loading and controlled delivery of HA over 24 hours and efficient exchange of oxygen and carbon dioxide in wearing due to the silicone content in DCL. The timing of HA delivery is synchronized with the typical wear duration of the contact lens. HA release is triggered by the ionic strength of the tear fluid, and the internal hydration of the silicone hydrogel lens helps to reduce tear evaporation and minimize hydrophobic deposits.

[0060] In one embodiment, the device is stored in a packaging solution having an ionic strength <10 mM and an osmolarity of 200 - 400 mOsm / 1. In one embodiment, the packaging solution has a pH between 5.0 and 8.0.

[0061] The packaging solution composition can be tuned by adjusting the ionic strength and desired osmolarity, for example, with respect to the concentrations of boric acid, glycerol, polyethylene glycol, and povidone.

[0062] The present device provides anti-inflammatory efficacy and relief from dryness and irritation of the eye.

[0063] Method for Preparing DCL

[0064] The present invention also provides a method for preparing the present DCL device. The method comprises the steps of: (a) soaking a contact lens in a first aqueous solution comprising a boundary charge modifier for a first period of time to bind the boundary charge modifier into the contact lens, (b) removing the contact lens from the first aqueous solution and rinsing the contact lens to remove unbound boundary charged modifier, and (c) immersing the contact lens into a second aqueous solution comprising the hyaluronic acid for a second period of time to load the hyaluronic acid into the contact lens, to obtain the device.

[0065] In the present method, the boundary charge modifier is embedded at the polymer-pore boundary surfaces in the lens matrix by an aqueous solution soaking. The boundary charge

[0066] - 7 -

[0067] 181727421.1 modifier is first dissolved in an aqueous solution, and then the lens is soaked in a volume reservoir of this aqueous solution for a determined period to ensure sufficient uptake. After removing the first aqueous solution from the surface of the contact lens loaded with the boundary charge modifier, the contact lens is then placed in an aqueous solution of HA to create a boundary charged double layer which consists of an immobilized layer of the charges of the head groups of the boundary modifier paired with the other half of charges from the HA and other counter ions. The boundary charge modifier is a molecule having a charged head group with its hydrophobic tail implanted and thus, immobilized, at the polymer domain’s pore boundary surface throughout the use life of the device attributed to its sparse aqueous solubility.

[0068] In one embodiment, the first aqueous solution comprises a mixture of alcohol and water. The boundary charge modifier is embedded into the polymers of the soft contact lens matrix by a swelling process by soaking the boundary charge modifier in the aqueous solution comprising an alcohol, such as ethanol, 1 -propanol, or 2-propanol.

[0069] The desired sparse solubility of boundary modifiers inherently imposes a hurdle of loading them by all aqueous solutions. The hydrophobic nature of boundary charged modifiers requires dissolution with a less hydrophilic cosolvent such as ethanol. Ethanol is the least toxic compared to many other organic solvents. Ethanol is highly aqueous miscible, and most favorably, a good solvent that can swell the silicone phase to substantially accelerate the loading rate. Further, alcohols such as ethanol are commonly used in contact lens manufacturing for demolding steps and then routinely washed off by water to complete the processing. Other alcohol options that are commonly used are 1 -propanol and 2-propanol. To avoid a significant degree of swelling, which could alter the lens morphology, an ethanol- water mixture is most efficient in shortening the uptake cycle time as well as preserving lens’ optical performances. By soaking a SiHy in an ethanol-water solution containing the boundary charge modifier, the modifier molecules can permeate within the lens matrix and redistribute themselves primarily at the water-polymer interfaces where the implantation of their hydrophobic tails into the polymer and the exposure of head-groups electrostatic charges into aqueous pores accounted for the most energetically favorable interfacial structure.

[0070] In one embodiment, when the first aqueous solution comprises an alcohol and water, Step (b) can be replaced by (bl) removing the contact lens from the ethanolic solution, and (b2) soaking the contact lens in water to remove the remaining ethanol from the contact lens.

[0071] In step (c), the contact lens already loaded with the boundary charge modifier is immersed into a second aqueous solution comprising HA.

[0072] - 8 -

[0073] 181727421.1 The ionic strength has a significant impact on the formation of the boundary charged double layer. The interfacial double charged layer thickness is inversely proportional to the square root of the pore fluid's ionic strength. It is important to maintain the pore and the storage solution fluid at a low ionic strength between 0.1-50 mM, preferably 1-50 mM, most preferably <10 mM, (for example, 5 mM), to allow higher partition of HA anions (due to a thicker interfacial zone) and thus, minimizing premature leaking of HA in storage. The low osmolarity of a low-ionic-content storage solution, for example, can be addressed by tuning with nonionic polyethylene glycol, glycerol, or other ocular isotonic additives.

[0074] When placing the device in the eye, the thickness change of the pore boundary layer in response to ionic strength change (e.g. from 1 - 10 mM in storage to the physiological 150 mM of common tears) automatically triggers a release of HA in the eye.

[0075] In one embodiment, the contact lens loaded with the boundary charge modifier is then loaded into an aqueous solution of HA, containing an extremely low ionic strength ~0 mM, for example, 0-5 mM or 0-10 mM to facilitate uptake of HA and the formation of the boundary charged double layer.

[0076] The inventors found that, beyond ionic strength, the ionic composition and particularly the pH of the HA loading solution has a significant impact on the electrostatic attraction within the boundary charged double layer. When the pH of a fluid medium is raised or lowered by one unit, the boundary charge modifier’s dissociation into cations is changed by a factor of ~10 according to the equilibrium equation of pKa. Such a change in the modifier’s dissociation rate will proportionally affect the surface charge density and local pH at the interface and consequently on the HA’s electrochemical potential as well as its release kinetics from the device.

[0077] In one embodiment, the contact lens loaded with the boundary charge modifier is then loaded into an aqueous solution of HA under a low ionic strength of 0.1 - 20 mM, or 1 - 10 mM. This aqueous solution has a pH of 4.5 to 5.5, or 5.0 to 5.2 that may be lowered from physiological pH range, 7.2 - 7.4 by using 0.1M HC1. At the pH range of 5.0 - 5.2, the boundary charge modifier’s dissociation is expected to change by a factor of 10 - 100, which will significantly increase the HA retention and therefore its uptake in the device. HA is generally stable over a wide pH range, but extreme pH conditions (both very acidic and very basic) can lead to degradation of the HA polymer. At pH 5.0-5.2, however, HA is still a stable molecule and thus, would stabilize interfacial adsorbed HA against acidity induced degradation.

[0078] In one embodiment, the present method further comprises a step (d) after the step (c): - 9 -

[0079] 181727421.1 (d) autoclaving the device of (c) at 100-150 °C for 10-30 minutes.

[0080] In one embodiment, the present method further comprises a step (e) after the step (d):

[0081] (e) sterilizing the autoclaved device in a packaging solution.

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

[0083] EXAMPLES EXAMPLE 1. Incorporation of boundary surface charge modifier CKC by an aqueous loading method.

[0084] In this example, the cationic boundary charge modifier is loaded into the silicone polymer domain of a silicone-hydrogel contact lens using an all aqueous solution of CKC.

[0085] The contact lenses are conventional (commercial) silicone hydrogel, Otufilcon A and Senofilcon A. These lenses consist of a HEMA hydrophilic polymer and silicone hydrophobic polymer. The Otufilcon A lenses have a 44% water content, and the Senofilcon A lenses have a 38 % water content in aqueous pores. The silicone hydrogel commercial lenses were rinsed with deionized water and then air-dried overnight before use. The lenses were soaked individually in 10 mL of 0.05 or 0. 10 mg / mL solution of CKC in deionized water. The soaking duration was 7 days at room temperature. Following the loading step, the contact lenses were taken out and rinsed with 5 ml of deionized water to remove unabsorbed CKC solution from the loading step. HA loading was carried out by immersing each CKC-loaded lens individually in heat-resistant glass vials holding 2.2 mL of a 0.5 mg / g HA solution in deionized water at room temperature for 1-day. The molecular weight of the HA used was 2,200 kDa. For lenses that were sterilized, these lenses were immersed in the HA solution and then autoclaved at 121°C and 15.8 psi pressure for 15 minutes. Following the autoclaving process, lenses were left in the HA solution at room temperature for 7 days before further evaluation.

[0086] - 10 -

[0087] 181727421.1 EXAMPLE 2. Incorporation of boundary surface charge modifier sphingosine by an aqueous loading method.

[0088] In this example, the cationic boundary charge modifier is loaded into the silicone polymer domain of a silicone-hydrogel contact lens using an aqueous solution of sphingosine with ascorbic acid.

[0089] The contact lenses are commercial silicone hydrogel, Otufilcon A silicone hydrogel. These lenses consist of a HEMA hydrophilic polymer and silicone hydrophobic polymer, having a 44% water content in aqueous pores. The silicone hydrogel commercial lenses were rinsed with deionized water and then air-dried overnight before use. The lenses were loaded with sphingosine using two different methods. In the first method, lenses were soaked individually in 5 mL of 0.5 or 1.0 mg / mL sphingosine solution in deionized water that has 5.0 mg / mL ascorbic acid. The soaking duration was 1 day at room temperature. In the second method, lenses were soaked individually in 5 mL of 0.3 mg / mL sphingosine solution in deionized water that has 17.5 mg / mL polyvinylpyrrolidone (PVP) and the soaking duration was also 1 day at room temperature. Following the sphingosine loading step, the unabsorbed boundary charge modifier on the lens after its removal from the first aqueous solution was carried out by placing the lens in deionized water. HA loading was done by immersing each sphingosine-loaded lens in separate heat-resistant glass vials holding 2.2 mL of a 0.1 mg / g HA solution in deionized water at room temperature for 1-day. The molecular weight of the HA used was 2,200 kDa.

[0090] The drug release experiments in the following examples were normally carried out by soaking the HA loaded lenses in 2.2 mL of release medium (IX PBS at 7.4 pH consisting of NaCl: 137 mM, KC1: 2.7 mM, Na2HPO4: 10 mM, KH2PO4: 1.8 mM). During the release experiments, the full 2.2mL was withdrawn at pre-determined time intervals and immediately exchanged for 2.2 mL of fresh PBS medium. The HA concentration in the sample aliquots was measured using a competitive hyaluronan enzyme-linked immunosorbent (ELISA) assay kit (Echelon Bioscience, Salt Lake City, UT). The assay kit had a detection range between 50 to 1600 ng / mL of HA, and some samples were diluted with PBS to prevent signal saturation.

[0091] EXAMPLE 3. Release Kinetics of HA From Silicone Hydrogel Contact Lenses Loaded with CKC or sphingosine using the aqueous loading method.

[0092] FIG 2. shows the cumulative release with time of HA from Otufilcon A and Senofilcon A lenses. Otufilcon A and Senofilcon A control lenses released approximately 1.2 pg of HA over 24 h, while Otufilcon A with CKC and Senofilcon A with CKC released ~20

[0093] - 11 -

[0094] 181727421.1 pig of HA and 32 pig of HA over 24 h, respectively. This means that the presence of CKC can increase the amount of HA released by at least 15 -fold depending on the lens type. The effect of steam sterilization using an autoclave on the cumulative release is shown in FIG 3. For the CKC modified lenses the amount of HA at 24 h decreased from 20 to 7 pg and from 32 to 9 pg for Otufilcon A and Senofilcon A, respectively, which is a decrease by approximately 3- fold. However, the amount of HA released of autoclaved lenses is still significantly higher than the control unmodified lenses which release only 1.2 pg. The difference in HA release between Otufilcon A and Senofilcon A is attributed to the different silicone composition of each lens which affects the water content and hence the porosity. FIG 4 and FIG 5 show the release kinetics of HA from Otufilcon A lenses loaded with sphingosine (SPH). Otufilcon A control lenses released less than 1.2 pg of HA over 24 h, while lenses loaded with sphingosine release from 12 pg to 20 pg of HA over 24 h. This shows that the presence of sphingosine can increase the amount of HA released from Otufilcon A lenses by at least 10- fold depending on the sphingosine loading concentration. The results of this example show that silicone contact lens modified with cationic boundary charge modifiers such as CKC and sphingosine can significantly release more HA compared to unmodified lenses.

[0095] EXAMPLE 4. Incorporation of boundary surface charge modifier CKC by ethanol / water loading method.

[0096] The cationic boundary charge modifier is loaded into the silicone polymer domain of a silicone-hydrogel contact lens using a 35:65 or a 40:60 (by weight) ethanol: aler mixture. The contact lenses are commercial silicone hydrogel, Otufilcon A silicone hydrogel. These lenses consist of a HEMA hydrophilic polymer and silicone hydrophobic polymer and have a 44% water content. For this experiment, lenses were rinsed with deionized water and then airdried before use. Three lenses were soaked in 2.5 g of 1.4 or 2.7 mg / g of CKC in 35:65 or 40:60 w / w ethanol: water. The soaking duration was 1 day at room temperature. Following the loading step, the ethanol / water solvent was removed, and the swollen lenses were washed in 10 mL of deionized water. This washing cycle was repeated in at least three cycles and in the last cycle the lenses were removed and stored in 5 mL of deionized water for subsequent use. HA loading was carried out by immersing each CKC-loaded lens individually in heat- resistant glass vials holding 2.2 mL of a 1.0 mg / g HA solution in deionized water at room temperature for 7-days. For lenses that were sterilized, lenses were immersed in the HA solution and then autoclaved at 121°C and 15.8 psi for 15 minutes. Following the autoclaving process, lenses were left in the HA solution at room temperature for 7-days. HA solutions

[0097] - 12 -

[0098] 181727421.1 were prepared using either a HA with molecular weight of 2,200 kDa or a HA with a molecular weight of 3.0 kDa.

[0099] EXAMPLE 5. Ethanol / water loading of sphingosine or phytosphingosine into silicone hydrogel commercial contact lenses.

[0100] The cationic boundary charge modifier is loaded into the silicone polymer domain of a silicone-hydrogel contact lens using a 51:49 and a 56:44 (by weight) ethanol :water mixture for sphingosine and phytosphingosine, respectively. The contact lenses are a commercial silicone hydrogel type made with Otufilcon A. These lenses consist of a HEMA hydrophilic polymer and silicone hydrophobic polymer and have a 44% water content. Lenses were rinsed with deionized water and then air-dried before use. The sphingosine (27.1 mg) was fully dissolved in 6.25 g of ethanol followed by addition of 6 g of water. Three lenses were soaked in 3.0 g of 2.2 mg / g sphingosine in 51:49 ethanol: water mixture for 24 hours. For phytosphingosine, phytosphingosine (27.1 mg) was fully dissolved in 6.9 g of ethanol followed by addition of 6 g of water. Three lenses were soaked in 3.0 g of 2.0 mg / g phytosphingosine in 56:44 ethanol: water for 24 hours. Following this step, the ethanolic solution was removed and lenses were washed in 10 mL deionized water in three cycles. After the last cycle, lenses were then stored in 5 mL of deionized water for subsequent use. HA loading was carried out by immersing each sphingosine or phytosphingosine-loaded lens in separate heat-resistant glass vials containing 2.2 mL of a 0.5 mg / g HA solution in deionized water at room temperature for at least 5-days. For lenses that were sterilized, lenses were immersed in the HA solution and then autoclaved at 121°C and 15.8 psi for 15 minutes. Following the autoclaving process, lenses were left in the HA solution at room temperature for at least 5 days.

[0101] EXAMPLE 6. Release kinetics of HA from silicone hydrogel contact lenses loaded with CKC or sphingosine using the ethanol / water loading method.

[0102] FIG 6 shows the impact of CKC concentration on the release kinetics of HA from Otufilcon A lenses. As shown, sterilized Otufilcon A lenses loaded with 1.4 and 2.7 mg / g CKC released 5.0 and 8.5 pg of HA over 24 hrs, respectively, while the sterilized Otufilcon A control released only 1.0 pg of HA. This is an increase of at least 5-fold in the total amount of HA released.

[0103] The effect is even greater when the CKC-loaded lenses are non-sterilized as shown in FIG 7. For instance, non-autoclaved Otufilcon A lenses loaded with CKC released about 30.0

[0104] - 13 -

[0105] 181727421.1 ig over 24 h which is at least 3-folds greater than what the autoclaved lenses released (i.e., 8.5 pig). We also tested the impact of the ethanolic CKC loading solution composition. Lenses were soaked in a CKC solution (2.7 mg / g) in either 40:60 ethanokwater or 35:65 ethanol: waler for CKC uptake and then loaded with HA as described before.

[0106] FIG 8 shows the cumulative HA release with time, and it is seen that sterilized CKC- loaded lenses with the 35:65 and 40:60 ethanokwater solution released 15 pg and 8.5 pg of HA over 24 h, respectively. Because silicone hydrogel lenses swell when immersed in ethanol, the difference in the amounts of HA released could be attributed to the differences in the degree of lens swelling caused by the different ethanokwater solutions, which also lead to differences in the amount of CKC loaded for each case.

[0107] FIG 9 shows the HA release kinetics from Otufilcon A lenses loaded with sphingosine (SPH) or phytosphingosine (PHS) using an ethanokwater as the loading solvent. It is shown that Otufilcon A lenses with SPH released about 7.5 pg and 5.0 pg HA over 24 hrs for nonautoclaved and autoclaved lenses, respectively, while Otufilcon A lenses with PHS released about 12 pg and 2.5 pg HA over 24 hrs for non-autoclaved and autoclaved lenses, respectively. The amounts of HA released from all cases are still significantly higher than for the control Otufilcon A lenses which released only 1.0 pg over the entire 24 h period. SPH and PHS are both long-chain amino alcohols with significant roles in cell membrane structure and cellular signaling. The main differences between these two compounds are the presence of a double bond in SPH, and the additional hydroxyl group in PHS. As a result, these compounds are expected to have different physicochemical properties that will make them have different impact on the HA uptake and release.

[0108] FIG 10 shows the correlation of cumulative low molecular weight (3.0 kDa) hyaluronic acid release amount with CKC presence for Otufilcon A lenses. Lenses loaded with CKC were soaked in a 2.7 mg / g CKC in 40:60 ethanol: water for CKC uptake and then loaded with HA as described before. A total of three lenses were soaked in 2.5 g of ethanol: water mixture with CKC for 1-day as explained in Example 4. As shown in FIG 10, sterilized Otufilcon A control lenses without CKC release less than 2.0 pg over 72 h, while sterilized Otufilcon A lenses loaded with CKC released approximately 200 pg HA after 72 h. This signifies that the presence of CKC increased the amount of HA released by approximately 100-fold.

[0109] - 14 -

[0110] 181727421.1 EXAMPLE 7. Effect of the ionic strength of loading medium on release kinetics of HA from Silicone Hydrogel Contact Lenses Loaded CKC

[0111] FIG 11 shows the HA release kinetics from sterilized CKC-loaded Otufilcon A lenses that were soaked in HA solution in phosphate buffer saline (PBS) at different ionic strengths: 0, 33, and 165 mM. For all cases, we used the same CKC loading concentration of 2.7 mg / g CKC using a 40:60 ethanokwater mixture as the loading solvent. Lenses loaded with HA solutions at 33 and 165 mM released approximately 2.0 and 1.5 pg of HA over 24 h, respectively. This is only moderately higher than the amount of HA released from unmodified lenses without CKC which releases only 1.2 pg over 24 h. On the other hand, lenses loaded with a HA solution with 0 mM ionic strength release almost 30 pg over 24 hour period, which is greater than the amount of HA released from the 33 mM ionic strength solution by 15-fold.

[0112] EXAMPLE 8. Effect of the pH of loading medium on release kinetics of HA from Silicone Hydrogel Contact Lenses Loaded CKC

[0113] FIG. 12 shows the HA release kinetics from sterilized CKC-loaded Otufilcon A lenses that were soaked in HA solution at different pHs: 4.33, 6.29, and 7.96. The HA aqueous solution composition has 10 mM Na2HPO4 and 1.8 mM KH2PO4 in deionized water. To modify the pH, we use a few pL of either a 0.1 M HC1 or a 0.1 M KOH solution in deionized water. CKC-loaded lenses loaded with HA solutions at pH of 7.96, 6.29, and 4.33 release approximately 4, 11, and 25 pg of HA over 24 h, respectively. This signifies an increase of HA release of nearly 3-fold when the pH changes from 7.96 to 6.29, and an increase of at least 2-fold when the pH changes from 6.29 to 4.33.

[0114] FIG. 13 shows the effect of HA loading concentration on the HA release amounts from sterilized CKC-loaded Otufilcon A lenses. Three different HA loading concentrations were used: 0.3, 1.0, and 1.8 mg / g HA in an aqueous solution containing 1 mg / g boric acid, 2 mg / g glycerol, and 5 mg / g PVP. The pH of the HA loading solution was 5.1. CKC-loaded lenses loaded with HA concentrations at 0.3, 1.0, and 1.8 mg / g HA release 11.0, 35.0, and 49.0 pg of HA over 24 h, respectively. There is approximately a 4.0-fold increase for HA released when the HA concentration is increased from 0.3 to 1.0 mg / mL. There is also a significant increase for HA release when the HA concentration is increased from 1.0 to 1.8 mg / mL.

[0115] FIG. 14 shows the HA release kinetics from sterilized CKC-loaded Otufilcon A lenses that were soaked in HA solution at either pH 5.12 or 7.26. The HA aqueous solution

[0116] - 15 -

[0117] 181727421.1 composition has 1 mg / g boric acid, 2 mg / g glycerol, and 5 mg / g PVP in deionized water. To change the pH, we use a few pL of either a 0.1 M HC1 or a 0.1 M KOH solution in deionized water. CKC-loaded lenses loaded with HA solutions at pH of 5.12 and 7.26 release 38.0 and 21.0 g of HA over 24 h, respectively, while the blank lenses loaded with HA solutions at pH of 5.12 and 7.26 release about 1.2 pg of HA over 24 h for both cases. This signifies that decreasing the pH of the HA loading solution from 7.26 to 5.12 increases the total amount of HA release by almost a 2-fold factor. It is important to note that blank lenses loaded with HA solution at pH 7.26 and 5.12 released less than 1.5 pg of HA over a 24-hour period. As a result, the presence of CKC in the contact lens is crucial for achieving a significantly higher release of HA when the pH of the HA solution is decreased from 7.26 to 5.12.

[0118] - 16 -

[0119] 181727421.1

Claims

WHAT IS CLAIMED IS:

1. A device for sustaining ocular delivery of hyaluronic acid, comprising: a. a silicone hydrogel soft contact lens comprising 20-50% w / w of a hydrophobic silicone polymer domain, 20-50% w / w of a hydrophilic polymer domain, and 30 - 60 % w / w of water; b. a boundary charge modifier in the amount of 0.5 - 3.0% w / w of the dried contact lens of (a), wherein the boundary charge modifier consists of a cationic head and a carbon chain length of 12 - 24, attached at the boundary surfaces of the water and the polymer domains of the contact lens; and c. hyaluronic acid of molecular weight from 3 k to 3,500 k Dalton, adsorbed to the cationic head of the boundary charge modifier at the pore-polymer interface.

2. The device according to claim 1, wherein the boundary charged modifier is a quaternary ammonium salt or a cationic lipid.

3. The device according to claim 2, wherein the boundary charged modifier is an alkyl amine having a carbon chain length of 8 - 24, a cationic quaternary ammonium salt having a carbon chain length of 8 - 24, or a sphingolipid .

4. The device according to claim 1, wherein the hyaluronic acid has a molecular weight of 3 kDato 3,500 kDa, or a combination thereof.

5. The device according to claim 1, wherein the device is stored in a packaging solution having an ionic strength <10 mM and an osmolarity of 200 - 400 mOsm / L.

6. The device of claim 5, wherein the packaging solution has a pH between 5.0 and 8.0.

7. A method of preparing the device of claim 1, comprising the steps of:(a) soaking the contact lens in a first aqueous solution containing the boundary charge modifier for a first period of time to load the boundary charge modifier into the contact lens,(b) removing the first aqueous solution from the contact lens and washing the contact lens with water,- 17 -181727421.1(c) immersing the contact lens into a second aqueous solution comprising the hyaluronic acid for a second period of time to load the hyaluronic acid into the contact lens, to obtain the device.

8. The method of claim 7, wherein the first aqueous solution comprises 20 % -70 % w / w alcohol in water..

9. The method of claim 7, further comprising a step (d) after the step (c):(d) autoclaving the device of (c) at 100-150 °C for 10-30 minutes.

10. The method of claim 8, further comprising a step (e) after the step (d):(e) sterilizing the autoclaved device in a packaging solution.- 18 -181727421.1

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