WO2019143293A1 - Self-implantable micro-drug-reservoirs for localized and controlled ocular drug delivery - Google Patents

Abstract

The present disclosure provides for a drug delivery device comprising a substrate having a surface comprising one or more microneedles, wherein each of the one or more microneedles comprises an apex shaped to penetrate a tissue layer, and each of the one or more microneedles is defined by a shell portion formed around a core portion, wherein the shell portion comprises a first biocompatible material and the core portion comprises a second biocompatible material which is different from the first biocompatible material. In a preferred embodiment, the first biocompatible material comprises methacrylated hyaluronic acid and the second biocompatible material comprises unmodified hyaluronic acid. Uses of the drug delivery device are also disclosed herein, preferably for ocular drug delivery. A method of fabricating a drug delivery device as described above is further disclosed herein.

Description

SELF-IMPLANTABLE MICRO-DRUG-RESERVOIRS FOR LOCALIZED AND CONTROLLED OCULAR DRUG DELIVERY

Cross-Reference To Related Application

[0001] This application claims the benefit of priority of Singapore Patent Application No. 10201800410Q, filed 16 January 2018, the content of it being hereby incorporated by reference in its entirety for all purposes.

Technical Field

[0002] The present disclosure relates to a drug delivery device and a method of fabricating such drug delivery device.

Background

[0003] Diseases affecting the cornea and other parts of eye threaten vision, which significantly impacts the quality of life, and with age-related eye diseases such as, cataract, glaucoma, ocular neovascular diseases including proliferative diabetic retinopathy, neovascular age-related macular degeneration, and corneal neovascularization (NV), such combinations become a leading cause of visual impairment and blindness. This becomes a major public health burden, which has probably turned into a global epidemic, as the increasing prevalence of such eye diseases is correlated to an upsurge in aging population, diabetes mellitus and prolonged use of contact lens. Moreover, eye trauma and chemical injury, which may be common in developing countries, can lead to corneal NV. There is thus a growing demand for treatment of such a devastating vision-threatening condition.

[0004] Conventionally, only a few therapeutic options are available. They include laser (e.g. photocoagulation) or surgical (e.g. diathermy, cautery, trabeculectomy) techniques that may be limited by the complexity and cost of the procedures as well as the risk of side effects (pain, bleeding, infection, perforation and other complications). Hence, treatment and management of ocular diseases remain an urgent unmet medical need.

[0005] Advances in understanding the molecular mechanisms of ocular diseases lead to development of potential therapeutic compounds, antibodies and proteins, some of which have already been approved by the US Food and Drug Administration (FDA). Nevertheless, efficient delivery of drugs into the eye remains challenging due to presence of multiple structural barriers (e.g. comeal epithelium and blood-retina barrier). For instance, the use of systemic route (e.g. intravenous, subcutaneous, oral) to administer such drugs requires the drugs to be in a large dose for achieving effective local drug concentration, and this may produce off-target adverse effects.

[0006] On the other hand, repetitive drug applications with high dosage may often be required for conventional topical administration (e.g. eye drops or ointments) due to extremely low bioavailability (less than 5% absorbed by eye) and fast clearance, and this may also lead to systemic side-effects (e.g. prolonged steroid eye drop usage causes ocular hypertension and systemic toxicity like uncontrolled hyperglycaemia).

[0007] In another instance, intraocular injection (e.g. intracameral and intravitreal injections) using conventional hypodermic needles to penetrate the surface barriers (cornea and sclera) suffers from poor patient compliance due to pain, need for frequent clinic visit, risk of infection, haemorrhage, and even permanent eye damage. Similar to topical eye drops, injecting drugs into ocular surface tissues (e.g. comeal intrastromal layer, sclera) suffers from poor drug retention due to back-flow of injected solution and subsequent tear wash-out. Furthermore, both conventional topical administration and local injection tend to produce burst release of drug that has a short effective duration, and this is not ideal for treating chronic progressive eye diseases, such as glaucoma.

[0008] Although contact lens-like hydrogels have been developed for improved topical delivery, such as providing for prolonged dmg residence time with minimal burst effect, their bioavailability remains poor.

[0009] In another attempt, while implanting intraocular drug reservoirs enables sustained release, it requires risky and painful surgical intervention. Hence, localized, long-lasting and efficient ocular drug delivery with good patient compliance remains an unmet medical need. [0010] In light of the above, there is a need to provide for a solution that ameliorates one or more of the abovementioned limitations. The solution of the present disclosure relates to a device for overcoming one or more of the issues mentioned above. The solution also relates to a method of forming such a device.

Summary

[0011] In a first aspect, there is provided for a drug delivery device comprising a substrate having a surface comprising one or more microneedles, wherein each of the one or more microneedles comprises:

an apex shaped to penetrate a tissue layer; and

each of the one or more microneedles is defined by a shell portion formed around a core portion, wherein the shell portion comprises a first biocompatible material and the core portion comprises a second biocompatible material which is different from the first biocompatible material.

[0012] In another aspect, there is provided for a drug delivery device as described in the above aspect for use in therapy.

[0013] In another aspect, there is provided for a drug delivery device as described in the above aspect for use in the treatment and/or prevention of a disease comprising an ocular disease, obesity, a metabolic disease, and/or a skin disease.

[0014] In another aspect, there is provided for a use of a drug delivery device as described in the above aspect in the manufacture of a drug delivery patch for the treatment and/or prevention of a disease comprising an ocular disease, obesity, a metabolic disease, and/or a skin disease.

[0015] In another aspect, there is provided for a method of treating and/or preventing a disease comprising an ocular disease, obesity, a metabolic disease, and/or a skin disease, the method comprising:

applying a drug delivery device as described in the above aspect to a tissue layer; and

removing the substrate of the drug delivery device from the one or more microneedles of the drug delivery device. [0016] In another aspect, there is provided for a method of fabricating a drug delivery device as described in the above aspect, the method comprising:

forming a shell portion comprised of a first biocompatible material for each of one or more microneedles in a mold which configures each of the one or more microneedles to have an apex shaped to penetrate a tissue layer;

forming a core portion comprised of a second biocompatible material on the shell portion of each of the one or more microneedles in the mold, wherein the second biomaterial is different from the first biocompatible material;

removing the one or more microneedles from the mold; and

attaching the one or more microneedles to a substrate to form the drug delivery device.

Brief Description of the Drawings

[0017] The drawings are not necessarily to scale, emphasis instead generally being placed upon illustrating the principles of the invention. In the following description, various embodiments of the present disclosure are described with reference to the following drawings, in which:

[0018] FIG. 1A is a schematic of the synthesis of methacarylated hyaluronic acid (MeHA) and the crosslinking through photo-activation. Briefly, N,N-dimethyl- formamide (133.3 mL) and methacrylic anhydride (4.7 mL) were added dropwise to HA solution (4.0 g of about 300 kDa HA in 200 mL deionized (DI) water), and adjusted to pH 8-9 with sodium hydroxide (NaOH). After continuous stirring for 1 day (4°C), the reaction solution was supplemented with sodium chloride (NaCl, 9.88 g) to precipitate MeHA in ethanol. MeHA precipitates were washed again with ethanol for 3 times, dissolved in DI water and dialyzed for 7 days. After lyophilisation, the purified MeHA was characterized by 'H NMR spectroscopy (Bruker Avance II 300MHz NMR). The degree of modification was determined by digital integration of the anomeric protons signals or methyl protons signals of HA and of the methacrylate proton signals at about 6.1, about 5.7, and about 1.9 ppm.

[0019] FIG. 1B shows a representative 'H NMR spectrum of HA. [0020] FIG. 1C shows a representative 'H NMR spectrum of MeHA. The degree of methacrylation was about 70% according to the ¹ H NMR map.

[0021] FIG. 1D is an illustration of ocular drug delivery using an eye-contact patch equipped with self-implantable micro-drug-reservoirs.

[0022] FIG. 2A is a schematic of the fabrication process of polymeric patch with an array of needle-shaped and double-layered micro-reservoirs.

[0023] FIG. 2B shows a bright-field image of stainless steel microneedles (MN) master-mold. The scale bar denotes 400 pm.

[0024] FIG. 2C shows a bright-field image of the corresponding double-layered MN obtained from the mold of FIG. 2B. The scale bar denotes 400 pm.

[0025] FIG. 2D shows a scanning electron microscopy (SEM) image of the stainless steel MN master-mold. The scale bar denotes 100 pm.

[0026] FIG. 2E shows a SEM image of the corresponding double-layered MN obtained from the mold of FIG. 2D. The scale bar denotes 100 pm.

[0027] FIG. 2F shows a SEM image of a stainless steel MN of the master-mold of

FIG. 2B. The scale bar denotes 10 pm.

[0028] FIG. 2G shows a SEM image of a corresponding double-layered MN obtained from the mold of FIG. 2D. The scale bar denotes 10 pm.

[0029] FIG. 2H shows the results of mechanical compression test (average from 4 measurements) of un-loaded and IgG(680)-loaded DL-MN (2 pg in 9 MNs).

[0030] FIG. 21 shows the bright-field images of DL-MN before (top) and after (below) compression.

[0031] FIG. 3A shows the characterization of double-layered MNs. Specifically, FIG. 3A shows a schematic of a polymeric patch with an array of DL-MN (the outer layer made of crosslinked MeHA and inner core made of HA).

[0032] FIG. 3B shows a SEM image of a polymeric patch with an array of DL-MN (the outer layer made of crosslinked MeHA and inner core made of HA). The scale bar denotes 100 pm.

[0033] FIG. 3C shows a representative confocal image of DL-MN loaded with IgG(680) in outer layer and IgG(488) in inner core. The scale bar denotes 100 pm.

[0034] FIG. 3D shows a representative confocal image of DL-MN with IgG(680) in outer layer only. The scale bar denotes 100 pm. [0035] FIG. 3E shows a representative confocal image of DL-MN with IgG(488) in inner core only. The scale bar denotes 100 pm.

[0036] FIG. 3F shows the in vitro fast and slow release profiles of DL-MNs in simulated tears or gelatin hydrogel (37°C). IgG(680) was loaded in outer layer and IgG(488) was loaded in inner core. Data represent mean ± SEM (n = 4 - 5).

[0037] FIG. 3G shows the in vitro release profiles of IgG(680) from HA-MN, MeHA- MN or DL-MN in PBS (37°C). IgG(680) was loaded in both layers of DL-MN. Data represent mean ± SEM (n = 4 - 5).

[0038] FIG. 3H shows the mechanical compression test of HA-MN, MeHA-MN and DL-MN. Data represent mean ± SEM (n = 4 - 5).

[0039] FIG. 31 shows the quantification of anti-VEGFR2 IgGs loaded in 3 x 3 array of DL-MN using an Easy-Titer IgG assay kit (ThermoFisher Scientific). 1 pg of IgG equally divided into the inner core and outer shell of MN.

[0040] FIG. 3J shows the protein staining on 12% polyacrylamide gel loaded with anti-VEGFR2 IgGs. Lane 1: freshly-prepared IgG. Lane 2: IgG collected from HA- MN. Lane 3: IgG collected from MeHA-MN. Lane M: molecular weight markers. Protein bands were stained with InstantBlue solution (Expedeon), and detected in a G:BOX Chemi XT4 imaging system (Syngene).

[0041] FIG. 3K shows the representative confocal images of immunostained VEGFR2 in primary human endothelial cells (HUVECs), using the freshly-prepared anti-VEGFR2 IgG, or anti-VEGFR2 IgG released from HA-MN / MeHA-MNs (being stored for 5 days) over 24 hrs. Hoechst 33342 (NucBlue Live ReadyProbes Reagent, Life Technologies) was used to stain the nuclei. The red fluorescence indicates the staining of VEGFR2 (Alexa Fluor 680). The scale bar denotes 10 pm.

[0042] FIG. 4A shows the in vitro anti-angiogenic activity of anti-VEGFR2 IgG in primary human endothelial cells (HUEVCs) (tube formation assay). Specifically, FIG. 4A shows anti-VEGFR2 IgG released from DL-MNs (being stored for 5 days) over 6 hrs (5 pg/ml), 24 hrs (10 pg/ml) or 120 hrs (50 pg/ml), which were used to treat the cells for about 18 hrs (with 10 ng VEGF). Freshly prepared IgG at different concentrations were also tested for comparison. Representative bright-field images (on the left) of tube formation in Matrigel and the statistics (rightmost bar graph) of tube length (% control) are shown. Data represent mean ± SEM (n = 3). [0043] FIG. 4B shows the in vitro biocompatibility of MNs in primary human corneal epithelial cells. Cells were exposed to different types of MNs for 2 days. Representative bright-field images (on the left) and the statistics of cell viability (% control) (using the AlamarBlue assay) (the rightmost bar graph) are shown. Data represents mean ± SEM (n = 3).

[0044] FIG. 4C shows the biphasic release profile of double-layered microneedles. Specifically, FIG. 4C shows the schematic release profile of DL-MN in agarose hydrogel.

[0045] FIG. 4D shows the merge of optical and fluorescence images of DL-MNs with the supporting patch (before insertion) and visualization of real-time release of IgG(680) and IgG(488) from DL-MN in agarose hydrogel (3 mins to 6 hrs). The scale bar denotes 200 pm.

[0046] FIG. 5A shows the average changes of fluorescence intensity (n = 4) in the region adjacent to DL-MNs due to released IgG(680) from the MeHA outer layer and IgG(488) from the HA inner core.

[0047] FIG. 5B shows the representative time-lapse confocal images and corresponding bright-field images of DL-MN in agarose hydrogel, showing the slow- release of IgG(680) from the outer layer of DL-MN. DL-MN becomes clear and transparent after 2 mins. The scale bar denotes 200 pm.

[0048] FIG. 5C shows the application of a polymeric patch containing 3 x 3 DL-MN array on the central region of porcine eye (about 30 seconds). Specifically, FIG. 5C shows a bright-field image before MN insertion.

[0049] FIG. 5D shows the application of a polymeric patch containing 3 x 3 DL-MN array on the central region of porcine eye (about 30 seconds). Specifically, FIG. 5D shows a bright-field image after MN insertion.

[0050] FIG. 5E shows a bright-field image of the cornea (cross-sectional view) upon MN insertion.

[0051] FIG. 5F shows a bright-field image of the cornea (cross-sectional view) after removal of the supporting patch.

[0052] FIG. 5G shows the hematoxylin and eosin stained section of the cornea showing the cavity caused by DL-MN penetration. The scale bar denotes 100 pm.

[0053] FIG. 5H shows the averaged insertion force of DL-MN into the cornea (n = 3). [0054] FIG. 51 shows a confocal image of embedded DL-MNs in cornea.

[0055] FIG. 5J shows the average fluorescence intensity changes (n = 4) in the region adjacent to DL-MNs.

[0056] FIG. 5K shows the confocal visualization of real-time release of IgG(680) and IgG(488) from DL-MN into the cornea.

[0057] FIG. 5L shows the transmittance of the fully-hydrated DL-MN, cornea and aqueous humor at different wavelengths (n = 3).

[0058] FIG. 6A shows the in vivo studies of self-implantable double-layered microneedles. Specifically, FIG. 6A shows a polymeric patch containing 3 x 3 DL- MNs applied on the central region of a mouse eye (about 30 seconds application duration).

[0059] FIG. 6B shows two bright-field images of the patch, before (left) and after (right) insertion into the mouse eye.

[0060] FIG. 6C shows the in vivo imaging of the mouse eyes, applied without (left) or with IgG(680)-loaded MN patch (right).

[0061] FIG. 6D shows two bright-field images of the eye treated with a IgG(680)- loaded MN patch, at day 0 (immediately after insertion) and at day 3.

[0062] FIG. 6E shows the red fluorescence spots in the cornea marking the penetration sites.

[0063] FIG. 6F shows the representative histological changes of mouse cornea, at day

0, day 1, day 3 and day 7 after MN patch application. The scale bar denotes 100 pm.

[0064] FIG. 6G shows the weight of the mouse at day 0 (2 hrs after insertion), day 1, day 3 or day 7 after patch application. Data represent mean ± SEM (n = 5).

[0065] FIG. 6H shows the food intake of the mouse at day 0 (2 hrs after insertion), day 1, day 3 or day 7 after patch application. Data represent mean ± SEM (n = 5).

[0066] FIG. 61 shows the grimace scale for pain assessment at day 0 (2 hrs after insertion). Data represent mean ± SEM (n = 5).

[0067] FIG. 6J shows the in vivo distribution of IgG in mice, treated without (control) or with systemic injection (intraperitoneal, I.P), eye drop (ED) (one-side only) or MN patch application (one side only) of IgG(680). The representative fluorescence images of dissected organs, 2 hrs after treatment, are shown. [0068] FIG. 6K shows the quantitative analysis (mean ± SEM; n = 5) of fluorescent intensities of eyes treated with MN patch at day 0 and day 3.

[0069] FIG. 7 A shows images of the DL-MN patch improving therapeutic efficacy of anti-VEGF therapy. Mouse eyes were treated differently 2 days after being inflicted with alkali-bum, and examined at day 7. Specifically, FIG. 7A shows the representative images of differently treated eyes. The dotted lines indicate the extent of neovascular outgrowth from the limbus.

[0070] FIG. 7B shows the quantifications of comeal neovascularization (mean ± SEM; n = 6 - 10). *p < 0.05, **p < 0.01 vs. control (control IgG eye drop, ED); ##p < 0.01 between indicated pairs.

[0071] FIG. 8A demonstrates the DL-MN patch for ocular delivery of DC 101 and diclofenac (Diclo) for synergistic therapeutic effect on comeal NV. Mouse eyes were treated differently 2 days after being inflicted with alkali-bum, and examined at day 7. Specifically, FIG. 8A illustrates dmg loadings in DL-MN and the representative images of differently treated eyes.

[0072] FIG. 8B shows the quantifications of comeal neovascularization (mean ± SEM, n = 5). The dotted lines indicate the extent of neovascular outgrowth from the limbus. *p < 0.05, **p < 0.01 vs. control; #p < 0.05, ##p < 0.01 between indicated pairs.

[0073] FIG. 8C demonstrates the DL-MN patch for ocular drug delivery. Specifically, FIG. 8C demonstrates DC 101 (0, 0.5, 1 or 2.5 pg in outer layer of MeHA) loaded in DL-MN. The mouse eyes were then treated differently 2 days after being inflicted with alkali-burn, and examined at day 7. Representative images of differently treated eyes and quantifications of comeal neovascularization (mean ± SEM; n = 4 - 5) are shown. The dotted lines indicate the extent of neovascular outgrowth from the limbus. Quantifications of macrophage accumulation (% of positive staining of a specific macrophage marker - F4/80) in differently treated eyes (mean ± SEM; n = 3 - 4) are also shown. *p < 0.05, **p < 0.01 vs. control.

[0074] FIG. 8D demonstrates the DL-MN patch for ocular drug delivery. Specifically, FIG. 8D demonstrates diclofenac (diclo) (0, 1, 2 or 5 pg in inner core of HA) loaded in DL-MN. The mouse eyes were then treated differently 2 days after being inflicted with alkali-burn, and examined at day 7. Representative images of differently treated eyes and quantifications of comeal neovascularization (mean ± SEM; n = 4 - 5) are shown. The dotted lines indicate the extent of neovascular outgrowth from the limbus. Quantifications of macrophage accumulation (% of positive staining of a specific macrophage marker - F4/80) in differently treated eyes (mean ± SEM; n = 3 - 4) are also shown. *p < 0.05, **p < 0.01 vs. control.

[0075] FIG. 9A shows the inflammation assessment of combinational therapy using double-layered microneedles. Mouse eyes were treated differently 2 days after being inflicted with alkali-burn, and examined at day 7. Specifically, FIG. 9 A shows the immunohistochemical staining of cornea with a specific macrophage marker - F4/80 surface antigen.

[0076] FIG. 9B shows the quantifications of macrophage accumulation (% of positive staining of F4/80; ED: eye drop).

[0077] FIG. 9C shows cytokine concentrations in collected tear film of differently treated eyes. Tear films were collected with MeHA-based MN-free eye patches (patches were put on the eyes for 1 min to absorb tear film).

[0078] FIG. 9D shows a SEM image of the patch soaked with PBS, and a confocal image of the patch with cy5-conjugated albumin absorbed from agarose hydrogel.

[0079] FIG. 9E shows the cytokine concentrations (IF6) measured with specific EFISA kits. Data represents mean ± SEM (n = 5). *p < 0.05, **p < 0.01 vs. control (no treatment on alkali-bum eyes); #p < 0.05, ##p < 0.01 between indicated pairs.

[0080] FIG. 9F shows the cytokine concentrations (VEGF) measured with specific EFISA kits. Data represents mean ± SEM (n = 5). *p < 0.05, **p < 0.01 vs. control (no treatment on alkali-bum eyes); #p < 0.05, ##p < 0.01 between indicated pairs.

[0081] FIG. 10A shows the representative confocal image of triple-layered MN, with the outer layer made of crosslinked MeHA containing IgG(488), middle layer made of HA (about 50 kDa) containing IgG(680), and inner core made of HA (less than 10 kDa) containing IgG(405). The scale bar denotes 200 pm.

[0082] FIG. 10B shows the time-lapse confocal images of real-time release from triple-layered MN in agarose hydrogel, showing rapid release of IgG(405), followed by slower release of IgG(680), and slowest release of IgG(488). The scale bar denotes 200 pm. [0083] FIG. 11A shows the prolonged release of HA-IgG conjugates loaded in microneedles. Specifically, FIG. 11A shows a schematic of HA and IgG conjugation, and loading into MeHA-MN.

[0084] FIG. 11B shows the UV-Vis spectra of HA, IgG(680) and HA-IgG(680) conjugate confirming the success of conjugation.

[0085] FIG. 11C shows the size distribution of HA, IgG(680) and HA-IgG(680) determined by dynamic light scattering (DLS) analysis, indicating that HA-IgG(680) forms larger nanoparticles.

[0086] FIG. 11D shows a representative confocal image of MeHA-MN loaded with HA-IgG(680) conjugates. The scale bar denotes 200 pm.

[0087] FIG. 11E shows the in vitro release profiles of HA-IgG(680) from MeHA-MN in PBS (0_/2 of about 1 week).

[0088] FIG. 11F shows the mechanical compression test of unloaded or HA-IgG(680) loaded MeHA-MN. Data represents the mean ± SEM (n = 4).

[0089] FIG. 12A shows layered microneedles for transdermal drug delivery. Specifically, FIG. 12A shows a bright-field image of the polymeric-patch equipped with an array of double-layered microneedles (DL-MN) (10x10 DL-MNs). The scale bar denotes 400 pm. FIG. 12A to FIG. 12C demonstrate DL-MNs can be well-pierced into the skin, creating about 300 pm deep cavities into the dermal layer at the insertion sites.

[0090] FIG. 12B shows representative confocal images of DL-MNs loaded with immunoglobulin G conjugated with Alexa Fluor: IgG(488) (green colour) in outer layer and IgG(680) (red colour) in inner core. The scale bar denotes 200 pm. FIG. 12A to FIG. 12C demonstrate DL-MNs can be well-pierced into the skin, creating about 300 pm deep cavities into the dermal layer at the insertion sites.

[0091] FIG. 12C shows hematoxylin and eosin stained section of the porcine skin showing the cavity caused by DL-MN penetration. The scale bar denotes 100 pm. FIG. 12A to FIG. 12C demonstrate DL-MNs can be well-pierced into the skin, creating about 300 pm deep cavities into the dermal layer at the insertion sites. Detailed Description

[0092] The following detailed description refers to the accompanying drawings that show, by way of illustration, specific details and embodiments in which the invention may be practised. These embodiments are described in sufficient detail to enable those skilled in the art to practice the invention. Other embodiments may be utilized and changes may be made without departing from the scope of the invention. The various embodiments are not necessarily mutually exclusive, as some embodiments can be combined with one or more other embodiments to form new embodiments.

[0093] Features that are described in the context of an embodiment may correspondingly be applicable to the same or similar features in the other embodiments. Features that are described in the context of an embodiment may correspondingly be applicable to the other embodiments, even if not explicitly described in these other embodiments. Furthermore, additions and/or combinations and/or alternatives as described for a feature in the context of an embodiment may correspondingly be applicable to the same or similar feature in the other embodiments.

[0094] The present disclosure provides for a drug delivery device and its uses. The present disclosure also provides for a method of fabricating the drug delivery device. The present drug delivery device may comprise a flexible base with one or more structures each designed to release drugs at different rates. Each of the one or more structures may have a shell portion and a core portion, wherein the shell portion and/or the portion may include at least one drug, and the shell portion and the core portion are configured to release the drug at different rates.

[0095] The structures may be referred to as microneedles in the present disclosure, as each of the structures has an apex designed to penetrate a biological layer, e.g. a tissue layer or a biological membrane. This means that the apex may comprise a sharp end shaped to allow the one or more structures for penetrating the biological layer. The term“flexible” as used herein refers to a material that can be bent without getting damage, and reverts to its original form even after bending.

[0096] With the above in mind, details of the present drug delivery device, its uses and its method of production, and their various embodiments, are described as follow. [0097] In the present disclosure, there is provided for a drug delivery device comprising a substrate having a surface comprising one or more microneedles. Each of the one or more microneedles may comprise an apex shaped to penetrate a tissue layer, and each of the one or more microneedles may be defined by a shell portion formed around a core portion, wherein the shell portion may comprise a first biocompatible material and the core portion may comprise a second biocompatible material which is different from the first biocompatible material.

[0098] The term“biocompatible” or its grammatical variant, as used herein, refers to a material or substance capable of being in contact with living tissues or organisms without causing harm to the living tissue or the organism. A biocompatible material or substance, in the context of the present disclosure, encompasses a biodegradable material or substance that may be readily decomposed by biological means. Such biological means may include, without being limited to, dissolution by a biological fluid such as tissue fluid, or the degradability may be brought about by a living organism, such as microorganisms.

[0099] Advantageously, the apex allows for each of the one or more microneedles to easily penetrate a biological layer, such as a tissue layer or membrane layer, by simply pressing the drug delivery device onto the tissue layer. To achieve this, the apex is designed to be a sharp end that can penetrate the biological layer without exerting a significant amount of force. For instance, a human finger can be used to press the drug delivery device against the tissue layer, and the one or more microneedles remain implanted in the tissue layer even after the finger is removed. This allows for the drug to be delivered from the one or more microneedles to a diseased tissue layer or affected area. Advantageously, the present drug delivery is patient-friendly in the sense that no surgical procedures nor supervision from a medical practitioner is required for the drug to be administered, and due to the ease of penetration, pain is minimized or eliminated. As used herein, the term“apex” refers to a point or vertex where all the lateral surfaces or all the lateral edges meet, and such a point or vertex is positioned opposite to the base of a microneedle. For example, a cone is shaped to have a circular base at one end and a sharp end apex disposed opposite to the circular base. In another example, a square -based pyramid is shaped to have a square base at one end and a sharp end apex disposed opposite to the square base. [00100] In various embodiments, each of the one or more microneedles may have one end that forms the base and an opposing end (i.e. opposite to the base) that forms the apex. In various embodiments, each of the one or more microneedles may have a pyramidal or conical shape. Each of the one or more microneedles may also have a tubular shape having a flat base at one end and an apex with a sharp tip positioned at the opposing end (i.e. opposite to the flat base).

[00101] For the ease of penetration, each of the one or more microneedles may be structurally designed for improved mechanical strength to withstand the opposing force from a biological layer when the drug delivery device is inserted into the biological layer. In this regard, each of the one or more microneedles may have an aspect ratio of 1:1 to 10:1, 1:1 to 3:1, etc., according to various embodiments. The aspect ratio, as used herein, refers to a ratio of the height to a dimension of the base. Where the microneedle is of a conical shape, the dimension of the base refers to the diameter of the conical base. In certain embodiments, each of the one or more microneedles has an aspect ratio of 2:1. Where the microneedle is of a pyramidal shape with a rectangular or square base, the dimension of the base may refer to the width of the rectangular or square base.

[00102] The present drug delivery device is also advantageous as it provides for release of drugs at different rates. Each of the one or more microneedles may be defined by a shell portion formed around a core portion, and the shell portion may have a slower drug release rate compared to the core portion. As the microneedles are formed of two different portions, they may be referred to as dual-layered or bi-lay ered microneedles.

[00103] For different drug release profiles, the shell and core portions may be made of different biocompatible materials. For example, the shell portion may be formed of a first biocompatible material that dissolves slower in a body fluid than the second biocompatible material for the core portion. The use of such a first biocompatible material also results in a shell portion that can maintain the shape of each of the one or more microneedles, especially the structural integrity of the sharp end apex, such that the drug delivery device has the ease of penetration as described above. Such a first biocompatible material, however, need not be of a stiffness that is higher than the second biocompatible material forming the core portion. In various embodiments, the first biocompatible material may have a slower drug release rate compared to the second biocompatible material.

[00104] Meanwhile, the second biocompatible material forming the core portion may be a substance that dissolves faster in a body fluid compared to the first biocompatible material forming the shell portion. As the core portion is formed of such a second biocompatible material, it can even provide for burst release of the drug that is encapsulated in the core portion, hence providing a different drug release rate from the shell portion. The second biocompatible material, which may be faster-dissolving, is designed specifically for forming the core portion, as it may not be able to maintain the structural integrity of the sharp-pointed apex if such a second biocompatible material is solely used to form the entire one or more microneedles. Hence, by configuring the one or more microneedles to have the shell portion and the core portion, different biocompatible materials can be utilized to improve the structural integrity of the microneedles for ease of penetration and yet provide for different drug release profiles.

[00105] In other instances, the shell portion may be formed of a first biocompatible material that dissolves faster in a body fluid than the second biocompatible material for the core portion. In such embodiments, the first biocompatible material may have a faster drug release rate compared to the second biocompatible material. The present device is therefore versatile in that the shell portion and the core portion may be configured to provide different drug release rates. Said differently, the first biocompatible material and the second biocompatible material may be configured to have different drug release rates, according to various embodiments.

[00106] In various embodiments, each of the first biocompatible material and the second biocompatible material may comprise hyaluronic acid, poly(lactic-co-glycolic acid), a polysaccharide, polyvinyl alcohol, or a crosslinked derivative thereof. A crosslinked derivative may be used for the first biocompatible material or the second biocompatible material, depending on whether the shell or core portion requires a slower drug release rate. One example of a crosslinked derivative may be methacrylated hyaluronic acid. Other crosslinked derivatives may include biocompatible crosslinked polymers as an example. In the present disclosure, the crosslinked derivative may also be a crosslinked derivative of poly(lactic-co-glycolic acid), a polysaccharide, or polyvinyl alcohol. Crosslinked derivatives tend to dissolve slower in a body fluid compared to their non-crosslinked counterparts. Non-limiting examples of polysaccharide may include chitosan, pullulan, etc. In embodiments where the shell portion needs to have a slower drug release rate, the first biocompatible material may comprise or consist of methacrylated hyaluronic acid while the second biocompatible material may comprise or consist of hyaluronic acid. In embodiments where the core portion needs to have a slower drug release rate, the second biocompatible material may comprise or consist of methacrylated hyaluronic acid while the first biocompatible material may comprise or consist of hyaluronic acid. In some embodiments, the first biocompatible material or the second biocompatible material may comprise methacrylated hyaluronic acid.

[00107] To further illustrate on the advantages of the shell portion and core portion, an example is referred to. In this example, the core portion is formed from hyaluronic acid, which is considered a fast-dissolving substance in body fluids. As hyaluronic dissolves at a fast rate in body fluids, it may not be suitable for prolonged release of drugs. In addition, as it dissolves at a fast rate, the structural integrity of the one or more microneedles cannot be suitably maintained. To counter this, the shell portion may be formed of methacrylated hyaluronic acid, which forms a biocompatible crosslinked polymer that has a slower dissolution rate in body fluid, and this in turn helps to preserve the structural integrity of the one or more microneedles, such that when the microneedles contact the body fluid, the microneedles do not start dissolving but maintain its structural integrity for ease of penetration. At the same time, because it dissolves slower, a drug that is encapsulated in the shell portion takes a longer time to be released or diffused out, thereby providing a different (i.e. slower) drug release profile from the core portion. Advantageously, the present drug delivery device provides for different drug release profiles without sacrificing the ease of penetration.

[00108] In addition, the present delivery device may also be made to have multiple portions (i.e. multi-layered) instead of just having the shell and core portion. In this regard, each of the one or more microneedles may further comprise a middle layer disposed adjacent to the shell portion and the core portion. In other words, the middle layer is sandwiched between the shell portion and the core portion. [00109] In various embodiments, the middle layer may be configured to have a different drug release rate from the first biocompatible material and the second biocompatible material. In various embodiments, the middle layer may comprise hyaluronic acid, poly(lactic-co-glycolic acid), a polysaccharide, polyvinyl alcohol, or a crosslinked derivative thereof. As already mentioned above, the crosslinked derivative may be a crosslinked derivative of poly(lactic-co-glycolic acid), a polysaccharide, or polyvinyl alcohol. In some embodiments, the middle layer may comprise methacrylated hyaluronic acid. The presence of a middle layer helps to encapsulate a drug that may not be compatible for encapsulation in the shell portion and core portion, or may not be compatible for encapsulation with another drug in the shell portion and core portion.

[00110] In various embodiments, each of the one or more microneedles may have a base. The base of each of the one or more microneedles may be formed directly on the surface of the substrate according to various embodiments. This means there is no intervening layer or material between the base of a microneedle and the substrate. To release a drug from the core portion, the core portion may not be entirely surrounded by the shell portion. That is to say, after detaching the one or more microneedles from the substrate to be embedded in a tissue, the base of the core portion of each microneedle becomes exposed, and this allows for a drug to be released from the core portion through the exposed portion of the base. The microneedles may be detached from the substrate by peeling off the substrate from the tissue layer after the microneedles are embedded in the tissue.

[00111] For ease of detaching the microneedles from the substrate, the substrate may comprise a biomaterial which has a lower molecular weight compared to the first biocompatible material and the second biocompatible material. The term “biomaterial” as used herein refers to a biological or synthetic material which can be introduced to a body tissue as part of an implanted device without causing any adverse effects on the body tissue. By using a biomaterial of lower molecular weight than the first and second biocompatible materials, the biomaterial being smaller in molecular weight can dissolve faster when contacted with a body fluid. A non limiting example of such biomaterial may be hyaluronic acid having a molecular weight of 3 kDa. Advantageously, this helps to avoid using an amount of force that may entirely remove the penetrated microneedles from the tissue layer. The body fluid may be tears from the eye or any other tissue fluid.

[00112] In various embodiments, the shell portion, the core portion, and/or the middle layer may comprise at least one drug. The drug may comprise an immunoglobulin, an anti-vascular endothelial growth factor, a nonsteroidal anti-inflammatory drug, a peptide, or a nucleic acid. The peptide may be a protein. The drug may also comprise an antibody or may be a small molecule drug. Other therapeutic drugs may be used. The drug may be conjugated to the first biocompatible material forming the shell portion, the second biocompatible material forming the core portion, and/or the material forming the middle layer, for slower release of drug, which in turn provides for different drug release profile. With the ability to encapsulate and release drugs, the one or more microneedles may be referred to as“drug-reservoirs” in the present disclosure.

[00113] The present disclosure also provides for a drug delivery device as described above for use in therapy. The present disclosure also provides for a drug delivery device as described above for use in the treatment and/or prevention of a disease comprising an ocular disease, obesity, a metabolic disease, and/or a skin disease. The present disclosure also provides for use of a drug delivery device as described above in the manufacture of a drug delivery patch for the treatment and/or prevention of a disease comprising an ocular disease, obesity, a metabolic disease, and/or a skin disease. The drug delivery patch may be an eye patch or a skin patch. The present disclosure also provides for a method of treating and/or preventing a disease comprising an ocular disease, obesity, a metabolic disease, and/or a skin disease, the method comprising applying a drug delivery device as described above to a tissue layer, and removing the substrate of the drug delivery device from the one or more microneedles of the drug delivery device. Embodiments and advantages described in the context of the present drug delivery device are analogously valid for uses of the drug delivery device as described herein above, and vice versa.

[00114] The present disclosure further provides for a method of fabricating a drug delivery device as described above. The method may comprise forming a shell portion comprised of a first biocompatible material for each of one or more microneedles in a mold which configures each of the one or more microneedles to have an apex shaped to penetrate a tissue layer, forming a core portion comprised of a second biocompatible material on the shell portion of each of the one or more microneedles in the mold, wherein the second biocompatible material is different from the first biocompatible material, removing the one or more microneedles from the mold, and attaching the one or more microneedles to a substrate to form the drug delivery device.

[00115] Embodiments and advantages described in the context of the present drug delivery device and its uses are analogously valid for the present method of fabricating the present drug delivery device described herein, and vice versa.

[00116] As mentioned above, the present method involves forming a shell portion for each of the one or more microneedles. Advantages of the shell portion have already been described above. As described above, for different drug release profiles, the shell and core portions may be made of different biocompatible materials. Specifically, the shell portion may be formed of a first biocompatible material that, when crosslinked, dissolves slower in a body fluid than the second biocompatible material for the core portion. The use of such a first biocompatible material also results in a shell portion that can maintain the shape of each of the one or more microneedles, especially the structural integrity of the sharp end apex, such that the drug delivery device has the ease of penetration as described above. Such a first biocompatible material, however, need not be of a stiffness that is higher than the second biocompatible material forming the core portion. The shell portion may comprise a first biocompatible material as already described above.

[00117] In various embodiments, forming the shell portion may comprise adding the first biocompatible material to the mold. The first biocompatible material may be allowed to set in the mold and dried to form the shell portion. In various embodiments, forming the shell portion may comprise drying the first biocompatible material in the mold. The drying can be carried out by any suitable means as long as the shell portion can be formed in the mold.

[00118] In various embodiments, the present method may further comprise subjecting the shell portion to a crosslinking agent after removing the one or more microneedles from the mold to form crosslinkages in the first biocompatible material. The crosslinking agent may comprise ultraviolet light. Forming crosslinkages allows for the shell portion to dissolve slower, thereby providing slower drug release compared to a biocompatible material that is not crosslinked.

[00119] After the shell portion is formed in the mold, the second biocompatible material for forming the core portion may be added on top of the shell portion in the mold. The second biocompatible material forming the core portion may be a substance that dissolves faster in a body fluid compared to the first biocompatible material forming the shell portion. As the core portion is formed of such a second biocompatible material, it can even provide for burst release of the drug that is encapsulated in the core portion, hence providing a different drug release rate from the shell portion. As the core portion is formed of a faster-dissolving biocompatible material, it may not be able to maintain the structural integrity of the sharp-pointed apex if such a second biocompatible material is solely used to form the one or more microneedles. Hence, by configuring the one or more microneedles to have such a shell portion and a core portion, different biocompatible materials can be relied on to improve the structural integrity of these microneedles for ease of penetration and yet provide for different drug release profiles. According to various embodiments, forming the core portion may comprise adding an aqueous solution to the mold, wherein the aqueous solution may comprise the second biocompatible material. According to various embodiments, forming the core portion may comprise drying the aqueous solution in the mold.

[00120] In other embodiments, the shell portion may be configured to provide for a faster drug release rate compared to the core portion, and this has already been mentioned above. In such embodiments, the first biocompatible material may have a faster drug release rate compared to the second biocompatible material. The present device is therefore versatile in that the shell portion and the core portion may be configured to provide different drug release rates. Said differently, the first biocompatible material and the second biocompatible material may be configured to have different drug release rates, according to various embodiments.

[00121] In various embodiments, each of the first biocompatible material and the second biocompatible material may comprise hyaluronic acid, poly(lactic-co-glycolic acid), a polysaccharide, polyvinyl alcohol, or a crosslinked derivative thereof. Forming crosslinkages to obtain the crosslinked derivatives for use as the first biocompatible material or the second biocompatible material depends on whether the shell or core portion requires a slower drug release rate. One example of a crosslinked derivative may be methacrylated hyaluronic acid. As already mentioned above, the crosslinked derivative may also be a crosslinked derivative of poly(lactic-co-glycolic acid), a polysaccharide, or polyvinyl alcohol. Other crosslinked derivatives may include biocompatible crosslinked polymers as an example. Forming crosslinkages to form crosslinked derivatives tend to cause slower dissolution of a biocompatible material in a body fluid compared to their non-crosslinked counterparts. Non-limiting examples of polysaccharide may include chitosan, pullulan, etc. In embodiments where the shell portion needs to have a slower drug release rate, the first biocompatible material may comprise or consist of methacrylated hyaluronic acid while the second biocompatible material may comprise or consist of hyaluronic acid. In embodiments where the core portion needs to have a slower drug release rate, the second biocompatible material may comprise or consist of methacrylated hyaluronic acid while the first biocompatible material may comprise or consist of hyaluronic acid. In some embodiments, the first biocompatible material or the second biocompatible material may comprise methacrylated hyaluronic acid.

[00122] Forming the shell portion and the core portion may take place in the mold. The mold may be one designed to have cavities that are shaped for each of the one or more microneedles to be formed with an apex. The advantage of forming the one or more microneedles through such a mold provides for ease of penetration of the drug delivery device due to formation of the apex. Advantages of the apex have already been described above.

[00123] As mentioned above, the apex allows for each of the one or more microneedles to easily penetrate a biological layer, such as a tissue layer or membrane layer, by simply pressing the drug delivery device onto the tissue layer. To achieve this, the apex is formed to be a sharp end that can penetrate the biological layer without exerting a significant amount of force. For instance, a human finger can be used to press the drug delivery device against the tissue layer, and the one or more microneedles remain implanted in the tissue layer even after the finger is removed. This allows for the drug to be delivered from the one or more microneedles to a diseased tissue layer or affected area. Advantageously, the present drug delivery is patient-friendly in the sense that no surgical procedures nor supervision from a medical practitioner is required for the drug to be administered, and due to the ease of penetration, pain is minimized or eliminated.

[00124] In various embodiments, the present method may further comprise forming a middle layer on the shell portion before forming the core portion on the middle layer. Advantages of forming a middle layer has already been described above. The forming of a middle layer helps to create another portion for encapsulating a drug that may not be compatible for encapsulation in the shell portion and core portion, or may not be compatible for encapsulation with another drug in the shell portion and core portion. The middle layer may also provide for a different drug release rate from the shell portion and/or the core portion.

[00125] The middle layer may be of the same biocompatible material as the core portion. The middle layer may be of a different biocompatible material from the first biocompatible material and the second biocompatible material so as to have the middle layer configured to have a different drug release rate from the shell portion (i.e. first biocompatible material) and the core portion (i.e. second biocompatible material). The middle layer may comprise hyaluronic acid, poly(lactic-co-glycolic acid), a polysaccharide, polyvinyl alcohol, or a crosslinked derivative thereof, according to various embodiments. As already mentioned above, the crosslinked derivative may be a crosslinked derivative of poly(lactic-co-glycolic acid), a polysaccharide, or polyvinyl alcohol. In some embodiments, the middle layer may comprise methacrylated hyaluronic acid.

[00126] In various embodiments, the substrate may be formed of a biomaterial comprising a lower molecular weight compared to the first biocompatible material and the second biocompatbile material, thereby allowing for easier detachment of the one or more microneedles from the substrate as the lower molecular weight biomaterial dissolves faster. This helps to avoid using an amount of force that may entirely remove the penetrated microneedles from the tissue layer. The body fluid may be tears from the eye or any other tissue fluid.

[00127] In the present method, forming the shell portion may further comprise adding at least one drug to be encapsulated in the shell portion. In the present method, forming the core portion may further comprise adding at least one drug to be encapsulated in the core portion. In the present method, forming the middle layer may further comprise adding at least one drug to be encapsulated in the middle layer. Embodiments regarding the drug have already been described above.

[00128] In the context of the present disclosure, the word“substantially” does not exclude“completely” e.g. a composition which is“substantially free” from Y may be completely free from Y. Where necessary, the word“substantially” may be omitted from the definition of the invention.

[00129] In the context of various embodiments, the articles“a”,“an” and“the” as used with regard to a feature or element include a reference to one or more of the features or elements.

[00130] In the context of various embodiments, the term“about” or“approximately” as applied to a numeric value encompasses the exact value and a reasonable variance.

[00131] As used herein, the term“and/or” includes any and all combinations of one or more of the associated listed items.

[00132] Unless specified otherwise, the terms "comprising" and "comprise", and grammatical variants thereof, are intended to represent "open" or "inclusive" language such that they include recited elements but also permit inclusion of additional, unrecited elements.

[00133] While the methods described above are illustrated and described as a series of steps or events, it will be appreciated that any ordering of such steps or events are not to be interpreted in a limiting sense. For example, some steps may occur in different orders and/or concurrently with other steps or events apart from those illustrated and/or described herein. In addition, not all illustrated steps may be required to implement one or more aspects or embodiments described herein. Also, one or more of the steps depicted herein may be carried out in one or more separate acts and/or phases.

Examples

[00134] The present disclosure provides for a strategy that utilizes an array of detachable microneedles. The microneedles may be disposed on an eye patch. This strategy can fulfill the clinical need for safe and effective ocular drug delivery for treatment of eye diseases and injuries, which can be challenging for conventional solutions due to presence of ocular barriers.

[00135] The microneedles can penetrate the ocular surface tissue and serve as implanted micro-reservoirs for controlled drug delivery. The biphasic drug release kinetics (i.e. both fast and sustained releases) enabled by the double-layered micro reservoirs significantly enhance therapeutic efficacy. Using corneal neovascularization as an example of a disease model, it is demonstrated that delivery of a neutralizing antibody to vascular endothelial growth factor receptor type 2 (DC101) via an eye patch comprising the microneedles produces about 90% reduction of the comeal neovascular area. Furthermore, quick release of an anti-inflammatory compound (diclofenac) followed by a sustained release of DC 101 provides synergistic therapeutic outcome.

[00136] A swellable eye patch without microneedles is also compared to assess the treatment effectiveness.

[00137] Through an eye patch application, drug delivery becomes easy and minimally invasive to ensure good patient compliance. The present intraocular drug delivery strategy allows for effective home-based treatment of many eye diseases.

[00138] The flexible polymeric eye patch may be equipped with an array of biodegradable and detachable microneedles for localized, highly efficient, and controlled ocular drug delivery. Microneedles can penetrate the ocular barriers (epithelial and stromal layers of cornea) with minimal invasiveness and be self- implanted as drug reservoirs for controlled drug release. The microneedles may have double-layer microneedles for biphasic release kinetics and packaging of multiple drugs for synergistic therapy. As a proof-of-concept demonstration, the superior effectiveness of such eye patch in the treatment of corneal neovascularization (NV) as compared to topical eye drop and fast drug-release approaches are shown in the examples below. A swellable eye patch without microneedles is also used to collect eye fluid for monitoring the therapeutic effectiveness based on biomarker detection. The present drug delivery device could be paradigm- shifting for long term home- based treatment and management of various eye diseases. [00139] Details of the present drug delivery device, its uses, and method of fabricating such a drug delivery device, are discussed, by way of non-limiting examples, as set forth below.

[00140] Example 1: Materials

[00141] Sodium hyaluronate with different molecular weights (MW) (less than 10 kDa, about 50 kDa, about 300 KDa) were purchased from Freda Biochem Co., Ltd. (China). Methacrylic anhydride (MA, 94%) and N,N-Dimethylformamide (DMF) were purchased from Sigma-Aldrich (USA). Immunoglobulins conjugated with Alexa Fluor 405 [IgG(405)], IgG(488), IgG(680), IgG(unconjugated), alamarBlue cell viability reagent, Medium 200, low serum growth supplements (LSGS), Geltrex LDEV-free reduced growth factor basement membrane matrix, bovine serum albumin (BSA), phosphate buffered saline (PBS), penicillin-streptomycin, and trypsin-EDTA were obtained from ThermoFisher Scientific (USA). Vascular endothelial growth factor receptor 2 antibody (VEGFR2; GTX14094; GTX10972) was obtained from GeneTex (USA). All reagents were of analytical grade and used without further purification.

[00142] Example 2: Synthesis and Characterization of Methacrylated Hyaluronic Acid (MeHA)

[00143] MeHA was synthesized by the following steps. Briefly, hyaluronic acid (HA) aqueous solution (4.0 g of HA, MW about 300 kDa, dissolved in 200 mL deionized water (DI)) was continuously stirred overnight (at 4°C), before 133.3 mL DMF and 4.76 mL MA were added dropwise to the HA solution. The reaction solution was then regulated to pH 8-9 with 1 M NaOH, followed by continuous stirring for another 18 hrs (at 4°C). Subsequently, the reaction solution was supplemented with 9.88 g NaCl to reach a 0.5 M NaCl concentration. After precipitation was done in ethanol, the MeHA precipitates were washed again with ethanol for 3 times before being dissolved in DI water, and dialyzed for 7 days. After lyophilisation, the purified MeHA was characterized by 'H NMR spectroscopy (Bruker Avance II 300MHz NMR). The degree of modification was determined by digital integration of the anomeric protons signals or methyl protons signals of HA and of the methacrylate proton signals at about 6.1, about 5.7, and about 1.9 ppm.

[00144] Example 3: Fabrication of Polymeric Patches with Micro-Drug-Reservoirs [00145] Microneedles (MN) patches were prepared via a micromolding method. Briefly, polydimethyl-siloxane (PDMS, Sylgard 184, Dow Coming USA) micromolds were created by pouring PDMS solution into the custom-designed stainless steel master-molds (Micropoint Technologies, Singapore), which was followed by degassing (10 mins in vacuum oven) and curing (70°C for 2 hrs). To fabricate eye patches with double-layered MN, MeHA aqueous solution (50 mg/mL MeHA together with 0.5 mg/mL Irgacure 2959 in DI water) was cast into the plasma-treated PDMS micromolds through centrifugation (3220 g, 5 mins). After air drying at room temperature in a fume hood (about 12 hrs), un-modified HA solution (about 50 kDa, 200 mg/mL) was applied and centrifuged (805 g, 5 mins) to fill the MN cavities. After air drying again for about 12 hrs, HA solution (less than 10 kDa, 500 mg/mL) was added to produce a robust supporting substrate, and dried overnight. Finally, MN patches were gently peeled off from the micromolds, and exposed to low-intensity ultraviolet light for 3 minutes (360 nm, about 2 mW/cm²).

[00146] Example 4: Characterization of MN Patches

[00147] MN patches were examined using a field-emission scanning electron microscope (FESEM) (JSM-6700, JEOL), and digital microscope (Leica DVM6). MN patches loaded with different IgGs, such as IgG(405), IgG(488) and IgG(680), were visualized with a confocal laser scanning microscope (LSM800, Carl Zeiss). The mechanical property of MNs was tested using an Instron 5543 Tensile Tester (Instron). A vertical force was applied to the MNs using a flat-headed stainless steel cylindrical probe (at a constant speed of 0.5 mm/min). The force was continuously recorded until a displacement of 450 pm was reached. Microneedle insertion test was performed. Briefly, MNs (3 x 3 MN patch, loaded with or without 2 pg IgG) were mounted onto the cylindrical probe, and pressed perpendicular to the isolated porcine cornea at a rate of 5 mm/min until a pre-set maximum load of 4 N was reached. Force exerted on the cornea by the MN as a function of its displacement into cornea was recorded. The insertion force was estimated when the force against the cornea showed discontinuity followed by a steep slope. The transmittance of the MNs (loaded with or without 2 pg IgG), isolated porcine cornea and aqueous humor (in PBS) was measured in a spectrophotometer (Shimadzu UV-1800). [00148] To evaluate the in vitro biocompatibility of MNs, primary human corneal epithelial cells (Merck Millipore, SCCE016) grown with the EpiGRO ocular complete media (Chemicon, Merck) were exposed to different types of MNs for 2 days, before analysing cell morphology using an inverted microscope (1X71, Olympus, equipped with a digital camera OlympusE330) and tested for cytotoxicity using an almarBlue cell viability assay. The absorbance of the incubated media containing 10% alamarBlue (about 3 hrs) was measured at 570 nm using a plate reader (SpectraMax M5, Molecular Devices).

[00149] The protein bands of IgGs being released from MNs (on 12% polyacrylamide gel) were stained with InstantBlue solution (Expedeon) and imaged with a G:BOX Chemi XT4 imaging system (Syngene, USA). The ability of released IgGs to bind IgG-specific proteins on human umbilical vein endothelial cells (HUVEC; Sigma- Aldrich, 200P-05N) was confirmed by immunofluorescence staining. Briefly, cells were fixed with 4% formaldehyde solution (15 mins) before washing with PBS (3 times, 5 mins each). After being blocked with 1% BSA in PBST (PBS with 0.1% tween20) (1 hr), the cells were incubated overnight in PBST containing 1% BSA and anti-VEGFR2 IgG (free IgG or IgG released from MNs). After washing with PBS, the cells were then incubated with the secondary antibody tagged with Alexa Fluor 680 (2 hrs), and after washing again, the cells were imaged using a confocal microscope.

[00150] To evaluate the in vitro bioactivity of IgGs loaded in MNs, HUVECs exposed to 10 ng/mL VEGF (SantaCruz Biotechnology) were treated with different doses of anti-VEGFR2 IgG (free IgG or IgG released from MNs) for about 18 hrs. Tube formation of HUVECs grown on the Geltrex matrix (ThermoFisher) were then recorded using an inverted microscope, and tube lengths were measured using the ImageJ (NIH.gov).

[00151] To determine the in vitro insertion capability, MN patches (3 x 3 MN array) were applied briefly to the porcine cornea (about 30 seconds), before the supporting base was removed. The corneas were then excised, washed with PBS, and analysed for the presence of fluorescence spots produced from IgG(680) loaded in MNs using confocal microscopy.

[00152] For histological analyses, the corneal tissues were fixed with 4% formaldehyde solution (24 hrs), and cryoprotected with 30% sucrose solution (24 hrs), before embedding in FSC22 Frozen Section Media (Leica Microsystem) for cryo sectioning (5 mhi thick) (CM 1950 cryostat, Leica Microsystems), and haematoxylin and eosin staining (Sigma- Aldrich).

[00153] Example 5: Evaluation of In Vitro and In Vivo Drug Release Profiles

[00154] To evaluate the release profiles of MNs, MNs were immersed in 0.5 mL of simulated tear fluid (NaCl 0.68 g, NaHCOs 0.22 g, KC1 0.14 g, CaCl₂-2H₂0 0.008 g, in 100 mL DI water, pH 7.4), gelatin hydrogel (15% w/v gelatin type b from Sigma- Aldrich, in DI water, pH 7.4) or PBS (pH 7.4), and placed in an incubator shaker (50 rpm, 37°C). IgG molecules (conjugated with Alexa Fluor dyes) released from MNs were measured using a fluorescence spectrometer (SpectraMax M5, Molecular Devices). The real-time visualization of IgG releases from double-layered MNs (DL- MNs) in agarose hydrogel (1.4% w/v, in DI water, pH 7.4) or porcine cornea were analysed using a confocal microscopy.

[00155] The in vivo fluorescence imaging of IgG release from MNs was conducted on mice (C57BL/6J, 7-8 weeks old male). Specifically, under anesthetized condition, a MN patch with 3 x 3 array of IgG(680)-loaded double-layered MNs was gently applied on one cornea (one eye only) for 30 seconds. The mice were then imaged immediately (day 0) or at day 3 using an in vivo imaging system (IVIS Spectrum, Perkin Elmer). In some experiments, MN treated mice were euthanized before collecting their eye balls and incising the cornea to identify the presence of fluorescence spots produced by MNs using confocal microscopy.

[00156] The in vivo bio-distribution of IgG(680) released from MNs was also analysed. Briefly, mice were divided into 4 groups, with treatment of 10 pg IgG(680) through intraperitoneal injection, topical eye-drop instillation or intra-corneal delivery using MN patch, or without treatment as control. After 2 hrs of treatment, the mice were euthanized and their eyes and major organs (liver, heart, kidney and lung) were dissected and visualized by IVIS imaging system.

[00157] Example 6: In Vivo Studies of MN Patches

[00158] All the animal experiments were approved by the Nanyang Technological University, Institutional Animal Care and Use Committee (NTU-IACUC) under protocol ARF-A0350. The mice were housed in light and temperature controlled facility (l2-hr light/ l2-hr dark cycle, 2l°C), and allowed free access to water and normal diet. Double-layered MN patches (3 x 3 MN array) were applied onto the central cornea area of anesthetized mice for 30 seconds. After removing the supporting base, corneas were imaged by a bright-field microscope. After receiving MN insertion, mice were immediately returned to their cages, allowing recovery from anaesthesia. After 10 mins, each mouse behaviour was monitored to access their pain (Grimace scales: orbital tightening, nose bulge, check bulge, and ear position). The body weight and food intake were also recorded (at day 1, 3 and 7). In some tested mice groups, mice were euthanized immediately after MN insertion or later (at day 1, 3 and 7), and corneas were collected to examine histological changes.

[00159] Example 7: Ocular Bum Mouse Model for Ocular Delivery of MN Patches

[00160] Chemical-bum injury to the mouse eye was also studied. Briefly, mice were anesthetized first before a sterilized Whatman filter paper (2 mm) soaked with 1 M NaOH solution was placed on the mice eyes for 30 seconds (at day 0). Eyes were then extensively flushed with sterilized PBS solution (about 10 mL) using a syringe. At day 2, corneal neovascular outgrowths were imaged by a microscope. Mice were then randomly divided into 6 groups, treated only once with eye-drop instillation of either non-specific control IgG or anti-VEGFR2-IgG (DC 101) (10 pg in 20 pL PBS), application of HA-only MN loaded with either control IgG or DC101 (1 pg), or application of DL-MN loaded with either control IgG or DC101 (1 pg, about 0.5 pg each in both layers). For testing the ocular delivery of two drugs, mice were randomly divided into different groups, treated only once with DClOl-loaded DL-MN (0.5 pg in MeHA layer), diclofenac-loaded DL-MN (1 pg in HA core), 2-drug-loaded DL- MN (DC101 in MeHA and diclofenac in HA core), or eye-drop instillation of both drugs. For all the interventions, anaesthesia was maintained throughout the procedure with 2% isoflurane. After 5-6 days, corneas were imaged and analysed using ImageJ. Comeal NV was quantified. Briefly, the vessel lengths (VL) were measured from the limbus to their leading edges, and sectoral circumference was expressed as clock hours (CH, 1 being 30 degrees of arc), and the vessel area (VA) was calculated using the formula, VA = 0.2p x VL x CH. In some experiments, mice were euthanized and corneas were collected for immunohistochemistry analyses of macrophage infiltration in corneal stromal layer. F4/80 as a major macrophage marker was stained with a specific antibody (R&D systems). Tear films from differently treated mice were collected by simply placing MeHA patch without MNs on their ocular surface for 1 min, before centrifugation of the patch at 16,000 g for 5 mins. Subsequently, interleukin-6 (IL6) and VEGF concentrations were determined by IL6 and VEGF ELISA kits (Invitrogen), respectively.

[00161] Example 8: Statistical Analysis

[00162] Quantitative data were represented as mean ± SEM. Statistical analysis was performed using one-way analysis of variance (ANOVA) followed by Tukey’s post hoc test. A p value of <0.05 was considered to be statistically significant.

[00163] Example 9: Results - Fabrication of Eye Patch with Double-Layered Microneedles

[00164] Hyaluronic acid (HA) is a non-sulphated glycosaminoglycan that is distributed abundantly throughout the human body in the connective tissues as well as vitreous eye fluid. As a natural biopolymer with unique viscoelastic property and transparency, HA has been widely used in ophthalmology, particularly in artificial tear solution as a lubricant for dry eyes. However, because of their fast dissolving nature, HA-MNs cannot maintain its sharp-pointed structural integrity and mechanical strength during penetration into a wet surface like cornea. In addition, HA-MN can only afford burst release of drugs. In comparison, crosslinked methacrylated HA (MeHA), which is synthesized by functionalizing HA with methacrylic anhydride (FIG. 1A) is more resistive to dissolution and offers slow release of drugs, but the stiffness of MeHA-MNs is inferior to HA MNs.

[00165] Combining the merits of HA and MeHA, an eye-contact patch equipped with double-layered MNs (DL-MN) for controlled ocular drug delivery is presently developed (FIG. 1D), using a simple micromolding method (FIG. 2A). The MNs have a HA inner core and a MeHA outer layer. Because the highly dissolvable HA is covered by MeHA, the MNs are able to penetrate the wet cornea surface. Briefly, a small amount of MeHA aqueous solution, with or without therapeutic compounds, was centrifuged into the reverse MN structures in the female polydimethylsiloxane (PDMS) mold. Hollow MN structures were formed after drying in ambient overnight as hydrophilic and viscous MeHA polymers tend to stick onto the hydrophilic surface of MN cavities in plasma-treated PDMS mold. Subsequently, unmodified HA solution, with or without therapeutic compounds, was filled in the remaining cavities and air-dried to form solid MNs. Finally, pure HA solution was introduced into the PDMS mold on top of the MN array to make the supporting substrate. After drying, the MN patch was peeled off from the mold and subjected to a brief exposure to ultraviolet light to crosslink MeHA outer layer of MNs.

[00166] As revealed by the scanning electron microscopy (SEM) and optical microscopy (FIG. 2B to 2G and 21), the fabricated patch consists of an array of pyramidal- shaped MNs with tip diameter of about 10 pm, height of about 500 pm, base width of about 250 pm, and inter-needle spacing of about 400 pm. The final MN dimensions are smaller than the stainless steel templates (300 pm bases with 600 pm height), due to the shrinkage of PDMS and HA/MeHA during the fabrication process. The MN design is based on the findings that pyramidal-pointed tips (compared to conical one) with the aspect ratio of 2:1 (height to base diameter) have better tissue penetration. The eye patches (about 2 x 2 mm) with a 3 x 3 MN array were used for mice whose cornea size is about 3 mm in diameter (FIG. 2B to 2G).

[00167] Immunoglobulins (IgG) labelled with Alexa Fluor 680 or Alexa Fluor 488 as the model therapeutic compounds were separately loaded in the different layers of MNs (FIG. 3A and 3B). The confocal fluorescence imaging confirms that red IgG (680) and green IgG (488) can be separately encapsulated in the outer and inner layers of MN, respectively, while the substrate is free of IgG molecules (FIG. 3C to 3E). Upon insertion of MNs into the tissue simply by thumb pressing on the supporting substrate, the tissue fluid is drawn into the MNs and quickly dissolves the interfacial HA layer between the MNs and substrate, thereby causing detachment of MNs (FIG. 1D). The embedded DL-MNs serve as the micro-reservoirs for localized and sustained drug release. The inner HA core being exposed to the tissue fluid undergoes quick dissolution and discharge of the drugs, whereas the outer MeHA layer dissolves slowly letting the drug molecules slowly seep through the cosslinked polymer matrix.

[00168] IgG recovered from 3 x 3 DL-MNs dissolved in phosphate buffer saline (PBS) was 0.92 ± 0.21 pg, correlating well with the nominal loading amount (1 pg) (FIG. 31). In vitro stability of encapsulated IgGs was tested by evaluating their molecular weight on polyacrylamide gel electrophoresis after storage of MN patch at 4°C for 1 week. Majority of IgGs (82.11 ± 11.2%) released from either HA or crosslinked MeHA matrix was intact as evidenced by the expected band at about 150 kDa (FIG. 3J). The bioactivity of IgGs (specific to vascular endothelial growth factor receptor 2, VEGFR2) released from MNs (being stored for 5 days) over 24 hrs was confirmed by immunofluorescence staining demonstrating the capability of binding and recognizing VEGFR2 on endothelial cells (FIG. 3K). The bioactivity of IgGs released from MNs (being stored for 5 days) over 6 hrs, 24 hrs or 120 hrs was further confirmed by their inhibitory effect on endothelial cell tube formation (FIG. 4A). Furthermore, in vitro biocompatibility of DL-MNs was proven by the well-preserved morphology and viability of corneal epithelial cells with the presence of either un modified HA or crosslinked MeHA (FIG. 4B).

[00169] Example 10: Characterization of Double-Layered Micro-Drug-Reservoirs

[00170] In vitro drug release kinetics was examined by monitoring the release profiles of different IgG molecules encapsulated in the two compartments of DL- MNs. As shown in FIG. 3F, IgG(488) loaded in the fast dissolving inner HA layer can be quickly released in both artificial tear fluid (which mimics tear) and gelatin hydrogel (which mimics comeal stromal tissue). Specifically, more than 80% was released within 5 mins in the former and 30 mins in the latter. In contrast, prolonged release profile was observed for IgG(680) loaded in the crosslinked MeHA outer layer (ti_/2 of about 2 days in tear fluid and about 3 days in gelatin hydrogel) because IgG molecules can only slowly diffuse through the interwoven meshwork of MeHA. In comparison to the fast release from HA-MNs or slow release from MeHA-MNs, the biphasic release profile was realized when a single drug molecule, IgG(680), was loaded in both compartments of DL-MNs (FIG. 3G).

[00171] To predict in vivo drug release profile, DL-MNs were embedded within the agarose hydrogel (which mimics corneal tissue in which water content is about 80%) and were continuously monitored under confocal microscope (FIG. 4C). Immediately after insertion, DL-MNs were quickly detached from the supporting substrate (less than 60 seconds) into the hydrogel because fluid was quickly drawn into MN- substrate junction. As the supporting substrate is made of highly-dissolvable low molecular weight HA (3-10 kDa), HA molecules at the MN-substrate junction dissolve rapidly as the fluid inside and at the surface are quickly drawn into the hydrophilic HA matrix. The patches were then removed allowing MNs embedded into the hydrogel. The mechanic perturbation caused by the penetration and substrate removal process also facilitates MN detachment. The fast-dissolving HA inner-core released the loaded green IgG(488) within 10 mins, while the crosslinked MeHA outer-shell gradually swelled and slowly discharged the encapsulated red IgG(680) into the hydrogel (FIG. 4D, FIG. 5A). The real-time release profile of double-layered micro-implants in hydrogel was captured by fluorescence microscopy as shown in FIG. 4D. Taken together, the data demonstrated the biphasic release kinetics of DL- MNs, i.e. a burst phase followed by a slow discharge over several days. Note that MNs in hydrogel were barely visible under bright-field imaging (FIG. 5B), suggesting that they are essentially transparent and hence suitable for use in corneal tissue.

[00172] The mechanical strength of MNs was assessed by compression test. Consistent with other studies, the mechanical strength of HA-MNs (about 0.4 N per needle) is strong enough for skin penetration (FIG. 3H). In contrast, the mechanical strength of crosslinked MeHA-MNs (about 0.15 N per needle) is much weaker. Because MeHA is highly viscose, only a lower polymer concentration can be used to fabricate MNs (50 mg/mL for MeHA-MNs vs. 200 mg/mL for HA-MNs). This compromises the mechanical strength of MeHA. The mechanical strength of DL-MNs is similar to that of HA-MNs (about 0.4 N per needle), indicating that the mechanical property of DL-MN is dictated by the inner HA core. In addition, drug loading into DL-MNs (2 pg) does not compromise the mechanical properties of DL-MNs (FIG. 2H). When an 18G hypodermic needle (outer diameter of 1.27 mm) was used, the force required to penetrate human cornea may be about 0.5 N per needle. Considering the sharp tip of a MN (about 10 pm), the force needed to penetrate the cornea can be much lower, thereby greatly lowering the amount of pain associated with administration of the MN.

[00173] To further confirm the insertion capability and drug releases of DL-MNs, MN patches were applied on the porcine cornea simply by thumb pressing (about 1.9 N, about 30 seconds) (FIG. 5C). Porcine eye has been commonly used as a good model for studying cornea, as its anatomical structures, water content, and thickness (about 0.9 mm) are similar to that of human cornea (about 0.6 mm). As demonstrated in FIG. 5E to 5G, MNs were well-penetrated and embedded into the porcine cornea. The subsequent histological study shows cavities into the comeal stromal layer about 150 pm deep which is about one third of MN height (FIG. 5G). This is because of compressive deformation of MN (FIG. 21), and the elastic property of the cornea provided by the densely-packed intertwining collagen fibrils. Such elastic deformations make polymeric DL-MNs advantageous over metallic MNs which may cause comeal puncture. The insertion force into the cornea was estimated (FIG. 5H). The force required to penetrate is about 0.05 N per needle, as indicated by the transition point where resistance from the tissue sharply increases. The success rate of MN embedment is about 85% as evidenced by visual inspection of the removed substrate (FIG. 5E and 5D) and the fluorescence marks left behind (FIG. 51). The fluorescence spots in the cornea from the embedded MNs were further analysed by confocal microscopy (FIG. 5K). Similar to the results shown in FIG. 4D, HA inner- core quickly discharged its cargo IgG(488) while the crosslinked MeHA outer-shell slowly released IgG(680) into the comeal tissue. It was also observed that the transmittance in the visible range of fully hydrated DL-MN is about 73-86%, which is comparable to that of cornea and aqueous humour (FIG. 5L). Taken together, these experiments demonstrated that DL-MNs are strong enough to penetrate into the cornea, easy to detach from the supporting patch after insertion, transparent inside the cornea, and capable of biphasic release kinetics.

[00174] Example 11: In Vivo Studies of Biosafety and Ocular Dmg Delivery

[00175] The capability of insertion and drug releases of DL-MNs were further investigated in vivo. DL-MN patch loaded with IgG(680) was gently pressed onto the cornea of mouse eye for about 30 seconds (FIG. 6A). MNs were quickly detached from the supporting substrate and implanted into the cornea. MN patch before and after being applied onto the cornea is shown in FIG. 6B and the detachment of MNs is evidenced. The treated eyes were flushed with PBS, and then examined with the fluorescence and bright- field imaging. In vivo fluorescence images indicated that fluorescence intensity of IgG(680) was strong only at the MN insertion spots, but absent in the control eye (FIG. 6C), suggesting the successful implantation of MNs into the cornea. FIG. 6D shows a representative bright-field image of mouse eye immediately after MN insertion (day 0). The MN applied corneas were further analysed under confocal microscope. It was found that IgG(680)-loaded MNs were embedded within the cornea (about 90% success rate), as evidenced by visual inspection of the removed substrate (FIG. 6B) and the fluorescence marks left behind (FIG. 6E). Subsequent histological examination revealed the small penetration cavities (about 100 pm) inside the comeal stromal layer, similar to the observation in porcine cornea (FIG. 5G).

[00176] The comeal anatomical structures after MN insertion were monitored over 1 week (FIG. 6F). No puncture on the cornea was observed in all experiments (FIG. 6F), suggesting that DL-MNs are strong enough to penetrate into the stromal layer, but not too stiff to spear throughout the whole cornea. The small insertion marks were only observed immediately after insertion in all mice (day 0) and almost disappeared after 24 hrs (FIG. 6F), suggesting the closure of epithelium through the spontaneous repairing process. At day 3 and day 7, DL-MN treated comeal epithelium restored its stmctural integrity and appeared as normal as untreated cornea in all mice (FIG. 6F). There were no significant differences of body weight and food intake between the control and test groups (FIG. 6G and 6H). There were no visible indications of corneal opacity, inflammation or haemorrhage in any MN-treated eyes (FIG. 6D). No signs of pain in the test group were observed based on the grimace scale pain assessment (FIG. 61). All these observations indicate that MN insertion and implantation into cornea is minimally invasive without causing obvious adverse effects on the eyes, general health state and behaviour.

[00177] As shown in FIG. 6J, systemic injection (peritoneal injection) of 10 pg IgG(680) led to strong fluorescence in all major organs (liver, kidney, heart, lung, etc.), but not the eyes. Local instillation of IgG(680) containing eye-drop also produced fluorescence signal in liver and lung, in addition to the treated eye. This is consistent with the notion that topical instillation results in systemic dmg distribution away from the eye via the highly vascularized conjunctiva. In contrast, intra-comeal delivery of IgG(680) at the same dose using MN patch only introduced fluorescence signal in the applied eye (much stronger than that caused by eye-drop). In addition, fluorescence signal in the eye was still observable for over 3 days (FIG. 6K), showing that the intra-comeal micro-implants act as drug depots for localized and sustained ocular drug delivery.

[00178] Example 12: Double-Layered Micro-Drug-Reservoirs Improving Efficacy of Anti-VEGF Therapy [00179] Eye trauma, including chemical injury and infection, can trigger corneal neovascularization (NV), and may cause comeal opacity, visual impairment and even blindness. Studies have shown that vascular endothelial growth factor (VEGF) is a key mediator of corneal NV. VEGF promotes blood vessel formation mainly via VEGF receptor type 2 (VEGFR2). Recently, anti- VEGF therapies (e.g. ranibizumab) have become a standard care for vaso-proliferative diseases, including those in the eye. However, for angiogenic ocular diseases like comeal NV, frequent high-dose topical application tends to be required to overcome ocular barrier (e.g. comeal epithelium), which is accompanied with adverse effects (e.g. subconjunctival haemorrhage).

[00180] On the other hand, conventional intraocular injection may cause infection, bleeding and retinal detachment. This example demonstrates advantages of the present DL-MN eye -patches for improving therapeutic efficacy of anti- VEGF drug (FIG. 7A and 7B).

[00181] Corneal NV was induced by the well-established alkali burn-injury model (day 0). At day 2 when neovascular growth started to appear from the limbus (the border between cornea and sclera), mice corneas were treated once with non-specific control IgG or anti-VEGFR2 IgG (DC 101) delivered through topical eye-drop or MN patch application (FIG. 7A and 7B). DC 101 was used because of its proven effectiveness against angiogenesis in murine models. Similar to the untreated eyes, eyes treated with control IgG eye-drop showed substantial comeal NV (1.48 ± 0.45 mm² vs. 1.50 ± 0.24 mm²) (day 7). Similarly, topical delivery of DC 101 via eye-drop (with a high dose of 10 pg in 10 pl) had no significant effect on comeal NV as compared to the untreated eyes (1.24 ± 0.21 mm²). In contrast, eyes treated with DC101 (about 1 pg) delivered through fast-dissolving HA-only MNs led to about 44% reduction in neovascular area (0.84 ± 0.43 mm²) (FIG. 7A and 7B). This is consistent with the previous studies that about 40% reduction of comeal NV can be expected using rapid ocular delivery approaches (e.g. subconjunctival injection). Even though the efficacy of highly targeted ocular dmg delivery using HA-only MN is much better than topical application, such rapid drug delivery could not efficiently suppress the continuous outgrowth of immature blood vessels because of the fast dmg clearance by natural fluid circulation in the eye. The same problem makes conventional intraocular injection ineffective for many eye diseases.

[00182] In comparison, it was found that DC 101 delivered through the present DL- MN (about 1 pg equally divided into the inner core and outer shell of MN) offered much improved therapeutic effect with 90% reduction of neovascular area (0.12 ± 0.17 mm²) (FIG. 7A and 7B). Both the vessel extension (from the limbus) and circumference (in clock hours) were significantly reduced by DClOl-loaded DL-MN patches, owing to the biphasic release profile of DC 101. DL-MNs not only provide initial bolus dose to quickly reach the therapeutic level at the early onset of the disease, but also sustain drug release to maintain therapeutic effect for a much longer period. Eyes treated with control IgG using either HA-only MN or DL-MN showed no significant changes compared with the untreated ones. In summary, the desirable therapeutic outcome observed in DL-MN treated group is attributable to both highly targeted and controlled drug delivery.

[00183] Example 13: Combinational Therapy Using Double-Layered Micro-Drug- Reservoirs for Synergistic Effect

[00184] Ocular delivery of multiple drugs at different stages of the disease progression can offer a more effective treatment outcome due to their synergistic effects. The initial inflammatory response is a factor that triggers ocular neovascularization (e.g. corneal NV, uveitis-related ocular NV). Under chronic inflammatory condition, inflammatory cells (e.g. macrophages) produce a large number of pro-inflammatory cytokines (e.g. interleukin 6, IL6) and angiogenic growth factors (particularly VEGF), which creates a vicious circle of persistent inflammation and neovascularization. To tackle this issue, DL-MNs were loaded with two drugs, nonsteroidal anti-inflammatory drug (1 pg diclofenac) in its fast-dissolving HA core and anti-VEGFR2 drug (0.5 pg DC101) in slow-dissolving crosslinked MeHA shell. As shown in FIG. 8A and 8B, ocular delivery of either only DC 101 in MeHA layer or only diclofenac in HA core using DL-MN patch exerted inhibition on neovascular area (0.52 ± 0.20 mm² and 0.63+ 0.25 mm², respectively). The therapy combining both drugs was much more effective (0.16+ 0.24 mm²). It is further shown in FIG. 8C and 8D that that even the high dosage of either one of the drugs, i.e. diclofenac (2 pg or 5 pg in HA) without DC101, or DC101 (1 pg or 2.5 pg in MeHA) without diclofenac, was not able to attain the therapeutic outcomes as good as that of a combinational delivery (1 pg Diclofenac in HA plus 0.5 pg DC101 in MeHA). This experiment further confirmed the synergistic combination of these two types of drugs. In comparison, topical instillation of same-dosage of both DC 101 and diclofenac produced no significant therapeutic effect (FIG. 8A and 8B).

[00185] The corneal inflammation was further analysed by immunofluorescence staining (FIG. 9A and 9B), and demonstrated that the cornea treated with diclofenac alone in HA core showed significantly fewer infiltrating F4/80+ macrophages as compared to the untreated cornea. Although DC 101 alone (loaded in MeHA shell) suppressed corneal inflammation to a lesser extent, DL-MNs-loaded with both diclofenac and DC 101 showed most significant suppression on the number of infiltrating macrophages. Taken together, these observations show that (i) fast release of diclofenac alone mainly suppresses inflammation with limited anti-angiogenic effects, (ii) conversely, the slow release of DC101 is poorly effective in inhibiting inflammation despite its strong anti-angiogenic effect, (iii) a quick release of diclofenac followed by a sustained release of DC 101 led to a much better treatment outcome.

[00186] Tear fluid can accurately reflect the dynamic changes of ocular surface tissue (e.g. cornea, sclera). As MeHA based patch is highly swellable, it was used to collect mouse tear film for analysis of the concentration of inflammatory and angiogenic cytokines (e.g. IL6, VEGF) in tear film after treatment (FIG. 9C). As shown in FIG. 9D, the pore size of the fully swelled MeHA-patch was 2-5 pm, suggesting that large proteins can be easily absorbed. The absorption of Cy5-conjugated albumin from agarose hydrogel confirmed the suitability of the patch to collect biomarkers in tear film. Both IL6 and VEGF levels in tear film of burn-induced comeal NV were significantly higher than those in normal tear film (FIG. 9E and 9F), and consistent with the results of macrophage infiltration shown in FIG. 9A and 9B, DL-MN delivery of DC 101 and diclofenac produced most significant effect on reducing those cytokine levels in tear film.

[00187] Example 14: Summarized Discussion of the above Examples

[00188] Effective ocular drug delivery for the treatment of vision threatening diseases

(such as glaucoma, neovascular complications, etc.) remains challenging due to the presence of various anatomical and physiological barriers. With the above examples, it has been demonstrated that the present eye patch with micro-drug-reservoirs self implantable into the ocular surface tissue for controlled drug release is advantageous for overcoming such a challenge. The flexible patch can be readily applied by gentle and brief thumb pressing on the ocular surface, which is as easy as wearing a disposable contact lens without causing discomfort or requiring high skills. As the micro-drug-reservoirs may be comprised of multiple compartments, they allow the release of the same drug with biphasic kinetics or sequential release of different drugs for synergistic therapy. Compared to conventional topical eye drop application which requires repeated high-dose instillation and often associated with systemic side- effects, and painful risky intraocular injections in clinics, the demonstrated eye patches offer a unique opportunity for patients to conveniently and effectively manage their eye disorders at home.

[00189] Here, comeal neovascularization (NV) in mice was used as the disease model to demonstrate the effectiveness of the eye patch. It has been demonstrated that intra corneal delivery of DC 101 (a monoclonal antibody that blocks VEGFR2) using micro-implants can achieve about 90% reduction of neovascular area with a single treatment of 1 pg dosage. In comparison, eye drop application of DC 101 even at a much higher dosage (10 pg) failed to show significant therapeutic effect. Systemic intraperitoneal injection of 1 mg DC 101 (every second day for 1 week) only led to marginal effect (about 20% reduction of neovascular area). This is expected as the major challenge faced by ocular drug delivery is the limited ability of drug molecules to penetrate the ocular tissue efficiently, due to the presence of ocular barriers. It has been reported that only 0.1% of bevacizumab (a monoclonal antibody against VEGF) in eye drop can reach the aqueous humour. It has been shown that eye drop application of 10 pg axitinib (another anti- VEGF agent; daily for 10 days) failed to show significant effect on corneal NV of mice. Clinical studies in human suggested that repeated high-dosage of topical administration is required for the treatment of corneal NV disease. However, only about 41% or 48% reduction of neovascular area can be achieved using ranibizumab eye drop (about 0.2 mg, 4 times daily for 3 weeks) or bevacizumab eye drop (about 0.1 mg, 5 times daily for about 3 months), respectively. Considering the present mouse experiments and that human cornea is about 20 times larger than mice cornea, single treatment with about 20 mg should be effective for human comeal NV based on the present approach. As the human cornea (about 600 pm thick, about 11 mm in diameter) is thicker and larger than mice cornea (about 150 pm, about 2.3 mm), a larger patch (e.g. 10 x 10 MN array) with longer MNs (e.g. 800 pm) may be suitable for human. With the ability to overcome the ocular surface barriers for localized delivery with high bioavailability, and achieve multi-phasic release kinetics or synergistic therapy from sequential release of multiple drugs, the present microneedle approach can realize low effective dosage and application frequency. This is needed to relieve the patient’s burden and enhance patient compliance.

[00190] Although double-layered MNs were used in the present experiments, triple compartmentalization can be fabricated (FIG. 10A and 10B). Here, crosslinked MeHA was chosen as the polymeric matrix for sustained drug release. Prolonged release for 5-6 days was achieved by simply loading the drug molecules in MeHA matrix. Even slower release can be achieved by conjugating the drug molecules with HA molecules to form nanoparticulate conjugates. Two weeks of sustained release can be attained when HA-IgG conjugates are loaded in MeHA matrix (FIG. 11A to 11F). Apart from HA, other biodegradable and biocompatible polymers may also be utilized, for example, FDA-approved synthetic polymer, poly(lactic-co-glycolic acid) (PLGA). PLGA has been used in conventional intraocular implants for sustained ocular drug delivery (e.g. ozurdex, a dexamethasone-loaded PLGA-based intravitreal implant). By choosing or mixing different PLGA molecules, mechanical properties of microneedle and drug release kinetics (several days to months) can be readily tailored.

[00191] The present microneedle approach can be applied for other eye diseases as well, for example, delivery of b-adrenergic receptor blockers or prostaglandin analogues for glaucoma, corticosteroids (e.g., prednisolone) for anterior uveitis, fluconazole for fungal keratitis. It may also be used for intra-comeal delivery of riboflavin to patients with keratoconus without the need of comeal epithelial scraping and debridement, thereby avoiding post-operative pain, infection and permanent damage usually associated with the traditional surgical methods. In summary, the microneedle eye patch demonstrated herein comprises implantable micro-drug- reservoirs for localized, controlled, and efficient ocular dmg delivery in a convenient, safe and painless manner, and provides a cost-effective home-based solution for many ocular diseases.

[00192] Example 15: Commercial and Potential Applications

[00193] The present device provides for a minimally-invasive self-implantable micro- drug-reservoirs that enable controlled release of therapeutic molecules, and is also suitable for transdermal drug delivery system. The device has been developed for prevention and treatment of various diseases, apart from ocular diseases. Non-limiting examples of such other diseases include obesity, diabetic mellitus and other metabolic diseases, skin infection and other skin diseases.

[00194] The trans-dermally delivered drugs could be any drugs or compounds that can be used in obesity, diabetic mellitus and other metabolic diseases, skin infection and other skin diseases, etc.

[00195] The layered microneedle may also include other biocompatible and biodegradable polymers, and non-limting examples include poly(lactic-co-glycolic acid) and its derivatives, chitosan and its derivatives, etc.

[00196] In summary, a self-implantable, biodegradable, and multi-layered (multi- compartmented) micro-drug-reservoirs for controlled ocular delivery of anti- angiogenic agents (or other drugs) have been developed herein. The sharp-pointed pyramidal- shaped (or otherwise shaped) microneedle arrays are tethered on a rapidly dissolvable and flexible polymeric patch, which can be easily and comfortably applied on the ocular surface daily or regularly by the patient at home without pain and need of skills (“patient- friendly”). After thumb-pressing for a short period of time (e.g. less than 2 minutes), the micro-drug-reservoirs may be detached from the patch substrate and be embedded in the ocular tissue serving as the reservoir-based drug delivery (“short administration time”). As the entire patch does not need to remain attach, eye discomfort and irritation are minimized (“comfort and convenient”). No surgical method is needed to implant the microneedles.

[00197] The biocompatible biopolymer, e.g. naturally occurring hyaluronic acids (HA), may be selected as the polymeric carrier, which is highly-biocompatible, biodegradable and inexpensive. [00198] The fabrication process is also simple and inexpensive, only loading and low speed centrifugation are involved without any destructive processes (pressurization, heating, etc.).

[00199] The bi-layered or multi-layered microneedles, with the outer layer of crosslinked methacrylated-HA (MeHA) and one or more inner layers of un-modified HA, were developed as one example to achieve the following advantages:

[00200] (1) Unmodified HA is used to make fast-dissolving layer of microneedles for quick delivery of anti-angiogenic agents (within minutes) while crosslinked MeHA is used for slow and sustained release of drugs over a few days.

[00201] (2) The implanted microneedles are protected from complete dissolution in ocular tissues with the help of intertwining meshwork of crosslinked MeHA filaments through which loaded macromolecules could diffuse into the surrounding tissue.

[00202] (3) The outer layer of crosslinked MeHA matrix maintains the microneedles’ sharp-pointed structural integrity during insertion into a wet surface like cornea (unlike rapidly-dis solvable microneedles which quickly dissolve in the aqueous layer covering the whole eyes).

[00203] (4) The filling of low molecular weight HA makes the microneedles strong enough to penetrate into the ocular tissue (unlike MeHA which results in microneedles with low mechanical strength).

[00204] (5) The drug release kinetics can be tailored by engineering the polymeric microneedles, e.g. crosslinking degree, loading the larger drug-polymer conjugates.

[00205] The achieved controlled drug release, which comprises a fast-release from un-modified HA and a sustained-release from outer crosslinked MeHA, is found to be superior to fast-releasing drug delivery platform, in terms of treatment efficacy and efficiency. The present drug delivery device allows for lowering the therapeutic dose, which not only produces lesser side effects but reduces the cost as well. The flexible patches as disclosed herein are equipped with micro-drug-reservoirs amenable for effective, patient-friendly, and convenient home -based treatment and management. Such a technology platform could be paradigm- shifting to combat not only angiogenic eye diseases but also other eye diseases.

[00206] While the invention has been particularly shown and described with reference to specific embodiments, it should be understood by those skilled in the art that various changes in form and detail may be made therein without departing from the spirit and scope of the invention as defined by the appended claims. The scope of the invention is thus indicated by the appended claims and all changes which come within the meaning and range of equivalency of the claims are therefore intended to be embraced.

Claims

1. A drug delivery device comprising a substrate having a surface comprising one or more microneedles, wherein each of the one or more microneedles comprises: an apex shaped to penetrate a tissue layer; and

each of the one or more microneedles is defined by a shell portion formed around a core portion, wherein the shell portion comprises a first biocompatible material and the core portion comprises a second biocompatible material which is different from the first biocompatible material.

2. The drug delivery device of claim 1, wherein the first biocompatible material and the second biocompatible material are configured to have different drug release rates.

3. The drug delivery device of claim 1 or 2, wherein the first biocompatible material has a slower drug release rate compared to the second biocompatible material.

4. The drug delivery device of any one of claims 1 to 3, wherein each of the first biocompatible material and the second biocompatible material comprises hyaluronic acid, poly(lactic-co-glycolic acid), a polysaccharide, polyvinyl alcohol, or a crosslinked derivative thereof.

5. The drug delivery device of any one of claims 1 to 4, wherein the first biocompatible material comprises methacrylated hyaluronic acid.

6. The drug delivery device of any one of claims 1 to 5, wherein each of the one or more microneedles has a pyramidal or conical shape.

7. The drug delivery device of any one of claims 1 to 6, wherein each of the one or more microneedles has an aspect ratio of 1:1 to 10:1.

8. The drug delivery device of any one of claims 1 to 7, wherein the substrate comprises a biomaterial which has a lower molecular weight compared to the first biocompatible material and the second biocompatible material.

9. The drug delivery device of any one of claims 1 to 8, wherein each of the one or more microneedles further comprises a middle layer disposed adjacent to the shell portion and the core portion.

10. The drug delivery device of any one of claims 1 to 9, wherein the middle layer is configured to have a different drug release rate from the first biocompatible material and the second biocompatible material.

11. The drug delivery device of claim 9 or 10, wherein the middle layer comprises hyaluronic acid, poly(lactic-co-glycolic acid), a polysaccharide, polyvinyl alcohol, or a crosslinked derivative thereof.

12. The drug delivery device of any one of claims 9 to 11, wherein the middle layer comprises methacrylated hyaluronic acid.

13. The drug delivery device of any one of claims 1 to 12, wherein the shell portion, the core portion, and/or the middle layer comprises at least one drug.

14. The drug delivery device of claim 13, wherein the drug comprises an immunoglobulin, an anti-vascular endothelial growth factor, a nonsteroidal anti inflammatory drug, a peptide, or a nucleic acid.

15. A drug delivery device of any one of claims 1 to 14 for use in therapy.

16. A drug delivery device of any one of claims 1 to 14 for use in the treatment and/or prevention of a disease comprising an ocular disease, obesity, a metabolic disease, and/or a skin disease.

17. Use of a drug delivery device of any one of claims 1 to 14 in the manufacture of a drug delivery patch for the treatment and/or prevention of a disease comprising an ocular disease, obesity, a metabolic disease, and/or a skin disease.

18. A method of treating and/or preventing a disease comprising an ocular disease, obesity, a metabolic disease, and/or a skin disease, the method comprising:

applying a drug delivery device of any one of claims 1 to 14 to a tissue layer; and

removing the substrate of the drug delivery device from the one or more microneedles of the drug delivery device.

19. A method of fabricating a drug delivery device according to any one of claims 1 to 14, the method comprising:

forming a shell portion comprised of a first biocompatible material for each of one or more microneedles in a mold which configures each of the one or more microneedles to have an apex shaped to penetrate a tissue layer;

forming a core portion comprised of a second biocompatible material on the shell portion of each of the one or more microneedles in the mold, wherein the second biomaterial is different from the first biocompatible material;

removing the one or more microneedles from the mold; and

attaching the one or more microneedles to a substrate to form the drug delivery device.

20. The method of claim 19, wherein forming the shell portion comprises adding the first biocompatible material to the mold.

21. The method of claim 19 or 20, wherein forming the shell portion comprises drying the first biocompatible material in the mold.

22. The method of any one of claims 19 to 21, further comprising contacting the shell portion with a crosslinking agent after removing the one or more microneedles from the mold to form crosslinkages in the first biocompatible material.

23. The method of claim 22, wherein the crosslinking agent comprises ultraviolet light.

24. The method of any one of claims 19 to 23, wherein forming the core portion comprises adding an aqueous solution to the mold, wherein the aqueous solution comprises the second biocompatible material.

25. The method of claim 24, wherein forming the core portion comprises drying the aqueous solution in the mold.

26. The method of any one of claims 19 to 25, wherein the first biocompatible material and the second biocompatible material are configured to have different drug release rates.

27. The method of any one of claims 19 to 26, wherein each of the first biocompatible material and the second biocompatible material comprises hyaluronic acid, poly(lactic-co-glycolic acid), a polysaccharide, polyvinyl alcohol, or a crosslinked derivative thereof.

28. The method of any one of claims 19 to 27, wherein the first biocompatible material comprises methacrylated hyaluronic acid.

29. The method of any one of claims 19 to 28, further comprising forming a middle layer on the shell portion before forming the core portion on the middle layer.

30. The method of claim 29, wherein the middle layer is configured to have a different drug release rate from the first biocompatible material and the second biocompatible material.

31. The method of claim 29 or 30, wherein the middle layer comprises hyaluronic acid, poly(lactic-co-glycolic acid), a polysaccharide, polyvinyl alcohol, or a crosslinked derivative thereof.

32. The method of any one of claims 29 to 31, wherein the middle layer comprises methacrylated hyaluronic acid.

33. The method of any one of claims 19 to 32, wherein forming the shell portion further comprises adding at least one drug to be encapsulated in the shell portion.

34. The method of any one of claims 19 to 33, wherein forming the core portion further comprises adding at least one drug to be encapsulated in the core portion.

35. The method of any one of claims 29 to 34, wherein forming the middle layer further comprises adding at least one drug to be encapsulated in the middle layer.