WO2021066747A1

WO2021066747A1 - Self-administrable and implantable polymeric micro-lance shaped device for controlled and targeted delivery

Info

Publication number: WO2021066747A1
Application number: PCT/SG2020/050552
Authority: WO
Inventors: Peng Chen; Aung THAN; Phan Khanh DUONG
Original assignee: Nanyang Technological University
Priority date: 2019-09-30
Filing date: 2020-09-30
Publication date: 2021-04-08

Abstract

Herein disclosed is a therapeutics delivery device implantable in a biological tissue, including a core which is biodegradable, wherein the core includes a first polymer composition, wherein the core has a height of at least 1 mm, and wherein the core has one end shaped as a sharp tip sufficient to penetrate the biological tissue, wherein the sharp tip includes a single bevel-angled shape, a double bevel-angled shape, a chiseled edge, a bullet-shape, or a frustum- shape. The core may be coated with one or more shells. In a preferred embodiment, therapeutics delivery device includes a poly(lactic acid-co-glycolic acid) (PLGA) core coated with a carboxymethyl cellulose (CMC) shell. Herein disclosed is also an applicator operable to implant the therapeutics delivery device, a method of producing the therapeutics delivery device, and applications of the therapeutics delivery device.

Description

SELF-ADMINISTRABLE AND IMPLANTABLE POLYMERIC MICRO-LANCE SHAPED DEVICE FOR CONTROLLED AND TARGETED DELIVERY

Cross-Reference to Related Application

[0001] This application claims the benefit of priority of Singapore Patent Application No. 10201909101Q, filed 30 September 2019, the content of it being hereby incorporated by reference in its entirety for all purposes.

Technical Field

[0002] The present disclosure relates to a therapeutics delivery device which is implantable and self-administrable. The present disclosure also relates to a method of producing the therapeutics delivery device and applications of the therapeutics delivery device. The therapeutics may include a drug, a cell, or a bioactive agent.

Background

[0003] Implantable drug delivery (IDD) devices offer several advantages over conventional drug delivery forms (e.g. oral or parenteral). First, implantable devices allow site- specific, localized drug administration where the drug is most needed. This may greatly enhance drug bioavailability and effects of the drug. It may also allow lower therapeutic dosage which not only minimizes potential side effects but also reduces the drug cost. Second, implantable devices allow for sustained (controlled) release of the bioactive agents. This may largely improve therapeutic efficacy as well as patient compliance, as the bioactive agents may be sustainably released for a longer period of time without the need of frequent and multiple applications. However, existing IDD devices, whether US-FDA approved ones, or the ones under consideration or testing in clinical trials, require a clinic visit, and one or more surgical or invasive procedures (e.g. surgical incision, large-bore needle insertion) to implant the system. As with any surgical procedures, pain, infections, bleeding, bruise (and reactions to the anesthetics if anesthesia is needed) may occur. Other potential side effects of conventional implants include, blistering, burning, coldness, discoloration of the skin, feeling of pressure, hives, infection, inflammation, itching, lumps, numbness, pain, rash, redness, scarring, soreness, stinging, swelling, tenderness, tingling, ulceration, and/or warmth at the insertion site.

[0004] For example, in one reported drug delivery device for subcutaneous implantation in animals, the device has a rod shaped polymeric inner matrix with an elongated body and two ends. The device is a few millimeters in terms of its length and diameter in size with no sharp-pointed microneedle. More importantly, the device undesirably requires surgery to be implanted and is not self-administrable.

[0005] In another reported example, a medicament-dispensing medical implant was described. The device was in the form of a stent, plug or a patch, fabricated from relatively non-inflammatory biogenic tissue or biopolymers, for implantation in the human body, for preventing restenosis following atherectomy. The device was a few millimeters in length and diameter with no sharp-pointed microneedle. Similarly, the device undesirably requires surgery to be implanted and is not self-administrable. [0006] In another example, an implantable drug delivery device that uses multiple reservoir elements covering with a shell and low-permeability barrier was reported. The shell may be breached by light irradiation to release drugs in the reservoir. The device is a few millimeters in length and diameter with no sharp-pointed microneedle. Similarly, the device undesirably requires surgery to be implanted and is not self- administrable.

[0007] In another example, a drug delivery system that provides for mixing various drugs, for using flow controllers to guide multiple drugs into a single or into multiple catheters, for controlling a dispensing of a fluid drug to an internal target site of a patient, has been reported. However, the system is apparently not for drug delivery implant.

[0008] Another implantable medical device for controlled drug delivery was reported. The device was elongated in shape, contained two or more discrete reservoirs, was a few millimeters in length and diameter with no sharp-pointed microneedle. The device undesirably needed the hollow bore of a needle, cannula, catheter, or trocar to inject the implantable medical device into a patient.

[0009] Another device for use in the treatment of osteonecrosis was reported. The device is cylindrically shaped, or in the form of multiple units, such as a plurality of beads. It needed surgery to implant the device and is not self-administrable. [0010] In another report, a layered polymeric monofilament fiber that includes side-by- side layers was reported, wherein a portion of each of the layers is exposed to the environment, for implantation in a patient. It is a cylindrical-shaped elongated fiber, several millimetres in length along its longitudinal axis with no sharp-pointed microneedle. It needed surgery to implant the fibers and is not self-administrable. [0011] A catheter and a stent were suggested for treating vascular conditions. The stent includes tubing having a wall defining a central lumen and a plurality of holes. However, this is not designed for a drug delivery implant.

[0012] Another example describes a method for performing non-invasive neurostimulation therapy of the brain via a nasal cavity, which is clearly not a drug delivery implant.

[0013] Another example describes a syringe needle-styled drug delivery system for implanting a drug-loaded rod into the vitreous of an eye. It includes a housing, with the first cannula that has an angulated end for facilitating penetration of tissue, and the second cannula to force the rod implant from the first cannula. The intravitreal implant is in a rod-and-cylindrical shape with no sharp-pointed end. It has a sharp tip of the first cannula from the applicator (similar to the conventional hyponeedle) to penetrate the tissue (not implant itself) in order to force the rod-shaped implant into the vitreous cavity by the second cannula. Nevertheless, it is not self-administrable.

[0014] Another example reports a drug delivery system comprising a plurality of rod shaped segments, for implanting it into the ocular region of the patients, for treating ocular conditions. The implant is a multi- segmented rod-shaped cylindrical shape. However, it needs surgery to implant the device and is not self-administrable.

[0015] Another example reports an implant into the suprachoroidal space for providing drugs to the eye. The implant has no sharp-pointed microneedle, needs surgical incision in the cornea to implant the device, and is not self-administrable.

[0016] Another example describes an implant inside the sclera for providing drugs to the eye. The implant has no sharp-pointed microneedle, needs surgical incision in the sclera to implant the device, and is not self-administrable.

[0017] Another implant was designed to be in the sclera for providing drugs to the eye. The device is a few millimeters in size, with no sharp-pointed microneedle, needs surgery to implant the device, and needs an adhesive or a sealant in order to fix to the sclera, needs at least one electrode and an opposite electrode which are fixed to the sclera, to disrupt the material in order to create openings and release the loaded agents. [0018] Another ocular implant device was contemplated for providing drugs to the vitreous cavity. The device is a few millimeters in size, with no sharp-pointed microneedle in shape, and needs surgery to implant the device.

[0019] Another report was on an implant device for intravitreal placement providing drugs to the vitreous cavity. The device is a few millimeters in size, with no sharp- pointed microneedle in shape, and needs surgery to implant the device.

[0020] A few examples describe a subconjunctival implant or an implant device for subconjunctival placement providing drugs to the eye. The device is a few millimeters in size, with no sharp-pointed microneedle in shape, and needs surgery to implant the device.

[0021] Another example describe an implant device for treating and/or preventing raised intraocular pressure, such as that associated with glaucoma or the use of corticosteroids, with carbonic anhydrase inhibitors. The device is a few millimeters in size, with no sharp-pointed microneedle in shape, and needs surgery to implant the device.

[0022] Another example contemplates a continuous release drug delivery implant which, among other mentioned places, can be mounted either on the outer surface of the eye or within the eye. The device is a few millimeters in size, with no sharp-pointed microneedle in shape, and needs surgery to implant the device.

[0023] Another example reports on an ocular implant device for providing drugs to the vitreous cavity. The device is a few millimeters in size, with no sharp-pointed microneedle in shape, and needs surgery to implant the device.

[0024] Another example relates to a bioadhesive ophthalmic inserts that are placed in the conjunctival sac. The device is a few millimeters in size, with no sharp-pointed microneedle in shape, and needs surgery to implant the device.

[0025] There is another example on a sustained-release implant for insertion into the vitreous of the eye. The device is a few millimeters in size, with no sharp-pointed microneedle in shape, and needs surgery to implant the device.

[0026] Another example reports an ocular drug delivery device placed in the cul-de- sac between the sclera and lower eyelid for administering the drug and acting as a reservoir. The device is a few millimeters in size, with no sharp-pointed microneedle in shape, and needs surgery to implant the device.

[0027] Another example contemplates a device for diagnosis of allergy. The device involves microneedle loaded with allergy antigen. However, ths device does not disclose its suitability for drug delivery.

[0028] Several other examples relates to device, system or methods that delivers a therapeutic agent (i.e. therapeutics) across the skin into the circulation, and/or withdraw fluid for analysis. For instance, the therapeutic agents may be delivered across the skin into the circulation to treat burn wound infection. In certain instances, device, system or methods are not related to drug delivery implant, or the drugs are only coated on microneedles. In one of the example, therapeutic agents are for delivery into the vagina mucosa, and thus used as an intravaginal device.

[0029] Another example reports a device to deliver insulin for diabetes patients using an integrated micro pump and microneedle array of closed-loop insulin delivery device. [0030] Several applications have also been reported, one of which concerns vaccination to a buccal mucosa using buccal microneedle patch. Another application is to deliver vaccine across the skin into the circulation, to use as transcutaneous immunization or vaccination

[0031] From the above, it may be observed that existing technology tends to be insufficient or does not provide a minimally-invasive and patient- friendly means to self- administer and implant the micro-drug-reservoirs for localized and controlled drug delivery into the deep tissue. So far, although there are a lot of technologies describing “implant or drug delivery devices” as a means for local drug delivery, almost all of these devices are large-sized, rod-shaped (no sharp-pointed lance shape), and needs surgical incision or large -bore hyponeedle injection to implant into the tissues. Although such implants may offer sustained drug release, unfortunately, invasive intervention (and clinic visit) amy still be required, and thus, the implantation is patient unfriendly because of pain, discomfort, irritation, and associated adverse effects such as bleeding, infection, scarring, etc.

[0032] There are a few technologies describing “microneedles (MNs) in drug delivery”. However, only “hollow or solid metal microneedles” were used to infuse drug solution or to dissolve coated-drug into the skin epidermis and dermis. Those drug delivery systems tend to be similar to the conventional drug injection except in using micrometre scale size needles. Unfortunately, the “hollow or solid metal microneedles” are not implantable. In addition, the drug release kinetics cannot be tailored using such hollow or solid metal microneedles, e.g. controlled release cannot be achieved.

[0033] Moreover, there seems to be a few conventional technologies using “dissolving microneedles (MNs) for controlled drug delivery”, but they are only able to deliver therapeutic agents into the skin epidermis and dermis, as those MNs are only about -600-800 pm in height. Therefore, those MNs tend to be insufficient for implanting directly into the deep tissue such as subcutaneous fat. In addition, due to the very short nature of MNs, the skin penetration ability of those dissolving MNs tends to be significantly inefficient, wherein only about 60-70% of MN length (-400-500 pm) may be inserted into the skin. Furthermore, those MNs embedded inside the skin may fail to perform as long-term drug depots because of the rapid turnover rate of epidermis and dermis (-28-40 days for human skin).

[0034] Therefore, in addition to the limitations mentioned above, existing IDD systems may not be patient friendly nor suitable for long-term home-based usages.

[0035] There is thus a need to provide for a solution that addresses one or more of the limitations mentioned above. The solution should at least provide for a safe, minimally- invasive, patient-friendly and self-administrable IDD strategy.

Summary

[0036] In a first aspect, there is provided a therapeutics delivery device implantable in a biological tissue, including: a core which is biodegradable, wherein the core includes a first polymer composition, wherein the core has a height of at least 1 mm, and wherein the core has one end shaped as a sharp tip sufficient to penetrate the biological tissue, wherein the sharp tip includes a single bevel-angled shape, a double bevel-angled shape, a chiseled edge, a bullet-shape, or a frustum- shape.

[0037] In various embodiments, the therapeutics delivery device may further include a shell which is biodegradable, wherein the shell includes a second polymer composition, wherein the first polymer composition and the second polymer composition have different degradation rate. There can be more than one shell.

[0038] In various embodiments, the core may be solid, hollow, or gutter-like.

[0039] In another aspect, there is provided an applicator operable to implant the therapeutics delivery device described in various embodiments of the first aspect, including: a housing including a lancing module and one or more holder tubes, wherein each of the one or more holder tubes is configured for loading the therapeutics delivery device therein, and wherein the lancing module is operable to drive the therapeutics delivery device out of the one or more holder tubes to implant the delivery device described in various embodiments of the first aspect into the biological tissue.

[0040] In another aspect, there is provided a method of producing the therapeutics delivery device described in various embodiments of the first aspect, including: providing a mold to form the core, wherein the mold includes a depth of at least 1 mm and a closed end including a single bevel-angled shape, a double bevel-angled shape, a chiseled edge, a bullet-shape, or a frustum-shape; depositing a first polymer solution into the cavity, wherein the first polymer solution includes the first polymer composition; and drying the first polymer solution to form the core.

[0041] In various embodiments, the method may further include coating a second polymer solution on the core to form the shell, wherein the second polymer solution includes the second polymer composition.

[0042] In another aspect, there is provided a method of treating a medical condition or delivering a therapeutics, wherein the method includes: operating the applicator described in various embodiments of the above aspect to implant the therapeutics delivery device described in various embodiments of the first aspect in the biological tissue. Brief Description of the Drawings

[0043] The drawings are not necessarily to scale, emphasis instead generally being placed upon illustrating the principles of the present disclosure. In the following description, various embodiments of the present disclosure are described with reference to the following drawings, in which:

[0044] FIG. 1A shows a bright field image of salt micro-particles used in the core- shelled micro-lances (ML) of the present disclosure. Scale bar denotes 50 pm.

[0045] FIG. IB shows a bright field image of a salt-mixed PLGA-ML (50% NaCl). Scale bar denotes 50 pm.

[0046] FIG. 1C shows the average fluorescence intensity changes (FITC) (n = 3) of the real-time release of FITC from the ML’s shell in 4% agarose hydrogel.

[0047] FIG. ID is a scanning electron microscopy (SEM) image of salt-mixed PLGA MLs (50% NaCl) changes over time in PBS at 37°C. Scale bar denotes 10 pm.

[0048] FIG. IE is a SEM image of salt-mixed PLGA MLs (50% NaCl) changes over time in PBS at 37 °C. Scale bar denotes 10 pm.

[0049] FIG. IF is a scanning electron microscopy (SEM) image of salt-mixed PLGA MLs (50% NaCl) changes over time in PBS at 37°C. Scale bar denotes 10 pm.

[0050] FIG. 1G is a scanning electron microscopy (SEM) image of salt-mixed PLGA MLs (50% NaCl) changes over time in PBS at 37°C. Scale bar denotes 10 pm.

[0051] FIG. 1H is a schematic illustration of the fabrication and application of polymeric core-shelled micro-lances of the present disclosure.

[0052] FIG. 2A shows bright field and confocal images of salt-mixed PLGA-MLs (50% NaCl) without CMC coating, or MLs coated with CMC (core-shelled ML). The ML is loaded with red Cy5 dye (top), and the CMC shell is loaded with green FITC dye (bottom). Scale bar denotes 250 pm.

[0053] FIG. 2B shows confocal images of real-time releases of Cy5 and FITC from core-shelled ML in 4% agarose hydrogel. Scale bars denotes 250 pm.

[0054] FIG. 2C shows the in vitro release profiles (PBS, 37 °C; n = 4). From top to bottom: CL316,243 in the shell of core-shelled ML, in both core and shell of core- shelled ML, in the core of core-shelled ML, in salt-free PLGA-ML.

[0055] FIG. 2D is a plot of mechanical compression test of MLs (n = 4). A schematic illustration of compression test is shown in the inset. Each data point represents the mean value and each error bar indicates standard deviation (mean ± SD). n represents the number of experiments. [0056] FIG. 2E shows the in vitro and release profiles of salt-mixed PLGA- MLs using bright field images of salt-free PLGA-ML and salt-mixed PLGA-MLs (50% salt) after 50 days of incubation in PBS at 37 °C. Scale bars dnotes 250 pm.

[0057] FIG. 2F shows in vitro release profiles of Cy5 molecules from different salt- mixed PLGA-MLs (PBS at 37 °C). Data are expressed as fluorescence intensity in arbitrary unit (a.u.) and represent the mean and standard deviation (n = 4). 200 denotes 0%, NaCl, 202 denotes 10% NaCl, 204 denotes 20% NaCl and 206 denotes 50% salt NaCl.

[0058] FIG. 2G shows bright field images of different salt-mixed PLGA-MLs incubated in PBS at 37 °C, at day 0, day 5, day 15 and day 30.

[0059] FIG. 3A is a bright field image of porcine tissue (skin epidermis, Epi; dermis, D and subcutaneous white adipose tissue, sWAT) after lancing ML. Scale bar denotes 2 mm.

[0060] FIG. 3B shows a representative histological section of porcine tissue depicting ML insertion track (arrow) and ML implanted inside sWAT. Scale bar denotes 250 pm. It is to be noted that ML appears transparent because of the thin slice.

[0061] FIG. 3C shows bright field images of mouse skin, immediately (day 0) or 1 day (day 1) or 2 days (day 2) after ML application. Scale bars denotes 1 mm.

[0062] FIG. 3D shows the body weight assessment examined at day 0, day 1, day 4 or day 7. Data represents mean ± SD (n = 4).

[0063] FIG. 3E shows the food intake assessment examined at day 0, day 1, day 4 or day 7. Data represents mean ± SD (n = 4).

[0064] FIG. 3F shows the pain assessment examined at day 0, day 1, day 4 or day 7. Data represents mean ± SD (n = 4). [0065] FIG. 3G shows representative histological sections of mouse tissue showing ML penetration through the skin (left) and implantation inside sWAT (right). Scale bars denote 250 pm.

[0066] FIG. 3H depicts the in vivo imaging of the mice, at day 0, day 1 or day 7. [0067] FIG. 31 depicts in vivo distribution of Cy5 delivered through intraperitoneal (IP) injection or ML application. It is noted that fluorescence intensity is highly concentrated at the ML implantation site. The representative fluorescence images of dissected organs, at day 0 (2 h post-treatment) or day 7, are shown. [0068] FIG. 3J shows representative force-displacement curves of salt-free PLGA-ML and salt-mixed PLGA-ML (50% salt). Failure load is defined as the maximum load the ML can sustain without buckling or breaking (fracture). Young's modulus of salt-free PLGA- ML and salt-mixed PLGA-ML (50% salt) are -1.4 and -1.7 GPa, respectively. [0069] FIG. 3K shows bright field images of MLs after pressing with their respective failure loads (-1.5N), showing a typical buckling of salt- free PLGA-ML in the left image and breakage of salt-mixed PLGA-ML (50% NaCl salt) in the right image. [0070] FIG. 3L shows mechanical compression tests of different salt-mixed PLGA- MLs (0%, 10%, 50% and 75% NaCl salt).

[0071] FIG. 3M shows the statistics of the failure loads of different salt-mixed PLGA- MLs (D) and MLs loaded with different salt%. (mean ± SD; n = 4).

[0072] FIG. 3N shows mechanical compression tests of the same salt-mixed PLGA- MLs (50% salt) without or with carboxymethyl cellulose (CMC) coating layer.

[0073] FIG. 30 shows the statistics of the failure loads of a salt-mixed PLGA-MLs loaded with different % of CL316,243 (E). (mean ± SD; n = 4).

[0074] FIG. 3P shows the representative insertion force of salt-mixed PLGA-ML (50% salt) into porcine skin. The insertion force for ML upon its penetration into the porcine skin is -0.1 N as indicated by a small dip in the force-displacement curve followed by a steeper slope.

[0075] FIG. 4A depicts effective anti-obesity treatment using diet-induced obese mice treated with MLs without drug (control), or with CL316,243 (CL, 1 mg/kg/week) via IP injection or ML application (CL316,243 loaded in both core and shell), once every 1 week (1 wk) or 2 weeks (2 wk) for 6 weeks. FIG. 4A shows representative images of differently treated mice at week 6.

[0076] FIG. 4B shows the relative body weight increases at week 6. Each data point represents mean ± SD from four to five mice in each group.

[0077] FIG. 4C shows the relative body weight increases over the course of treatment (C). Each data point represents mean ± SD from four to five mice in each group.

[0078] FIG. 4D shows the statistical analyses (mean ± SD, n = 4) of IgWAT, EpiWAT and interscapular BAT.

[0079] FIG. 4E shows representative images of IgWAT, EpiWAT and interscapular BAT. [0080] FIG. 4F shows intraperitoneal glucose tolerance test at week 5 (mean ± SD, n = 4).

[0081] FIG. 4G shows infrared thermographic images of the representative mice at week 6.

[0082] FIG. 4H shows the average body surface temperature of differently treated obese mice and non-obese mice (normal diet; dashed rectangle), measured from the indicated the areas (mean ± SD, n = 4). Statistical comparison among groups was performed using one-way ANOVA. * p < 0.05, ** p < 0.01 vs. control (no treatment). [0083] FIG. 41 shows a prototype image of a lancing applicator of the present disclosure, which includes a stainless- steel spring house within a chamber, and a shaft connected with a stainless- steel rod which is fitted into a tube holder. The spring loading force is -1.5 kgf.

[0084] FIG. 4J shows time-lapse images of a salt-mixed PLGA-ML (50% salt) lancing from the applicator, recorded with 240 fps. The lancing speed is determined to be ~1 m/s.

[0085] FIG. 4K shows a bright field image salt-mixed PLGA-MLs (50% salt) embedded within 4% agarose hydrogel, showing a typical ML’s trajectory path and insertion depth. The trajectory angle is -22° deviated from the insertion line (white dash line), and the insertion depth is 2-3 mm between the gel surface and the blunted end of ML. Scale bar denotes 2 mm.

[0086] FIG. 4L shows a bright field image of a salt-mixed PLGA-ML, showing its bevel angle (ø) of ~ 75°. Scale bar denotes 250 pm.

[0087] FIG. 4M is an illustrative diagram of the proposed ML’s trajectory mechanism: The asymmetry of the resistive forces from the agarose gel acting on ML’s tip causes a moment on ML’s tip, causing the whole ML to bend as it is advanced into the agarose gel. The net resultant force (F) experienced by ML is proportional to the sine of the bevel angle (sin ø).

[0088] FIG. 4N is an illustrative diagram of the proposed ML’s trajectory mechanism: The asymmetry of the resistive forces from the agarose gel acting on ML’s tip causes a moment on ML’s tip, causing the whole ML to bend as it is advanced into the agarose gel. [0089] FIG. 5A illustrates that present ML approach improves browning and metabolic effects of CL316,243 in obese mice. Diet-induced obese mice were treated with ML without drug (control), or with CL316,243 (CL, ~1 mg/kg/week) via IP injection or core-shelled ML application (loaded in both core and shell), once every 1 or 2 weeks for 6 weeks. FIG. 5A shows representative hematoxylin and eosin staining images of IgWATs isolated from differently treated mice. Asterisks * indicate white adipocytes while the circle indicates a brown-like adipocyte. Scale bar denotes 50 pm.

[0090] FIG. 5B illustrates that present ML approach improves browning and metabolic effects of CL316,243 in obese mice. Diet-induced obese mice were treated with ML without drug (control), or with CL316,243 (CL, ~1 mg/kg/week) via IP injection or core-shelled ML application (loaded in both core and shell), once every 1 or 2 weeks for 6 weeks. FIG. 5B shows representative immunoblots of UCP1, PRDM16, PGCla, aP2 and actin in IgWAT.

[0091] FIG. 5C illustrates that present ML approach improves browning and metabolic effects of CL316,243 in obese mice. Diet-induced obese mice were treated with ML without drug (control), or with CL316,243 (CL, ~1 mg/kg/week) via IP injection or core-shelled ML application (loaded in both core and shell), once every 1 or 2 weeks for 6 weeks. FIG. 5C shows respective statistics of UCP1, PRDM16, PGCla, aP2 and actin in IgWAT (mean ± SD, n = 4; normalized to actin density).

[0092] FIG. 5D illustrates that present ML approach improves browning and metabolic effects of CL316,243 in obese mice. Diet-induced obese mice were treated with ML without drug (control), or with CL316,243 (CL, ~1 mg/kg/week) via IP injection or core-shelled ML application (loaded in both core and shell), once every 1 or 2 weeks for 6 weeks. FIG. 5D shows quantification of serum levels of cholesterol (mean ± SD, n = 4). Statistical comparison between groups was performed using one-way ANOVA. * p < 0.05 vs. control.

[0093] FIG. 5E illustrates that present ML approach improves browning and metabolic effects of CL316,243 in obese mice. Diet-induced obese mice were treated with ML without drug (control), or with CL316,243 (CL, ~1 mg/kg/week) via IP injection or core-shelled ML application (loaded in both core and shell), once every 1 or 2 weeks for 6 weeks. FIG. 5E shows quantification of serum levels of triglyceride (mean ± SD, n = 4). Statistical comparison between groups was performed using one-way ANOVA. * p < 0.05 vs. control.

[0094] FIG. 5F illustrates that present ML approach improves browning and metabolic effects of CL316,243 in obese mice. Diet-induced obese mice were treated with ML without drug (control), or with CL316,243 (CL, ~1 mg/kg/week) via IP injection or core-shelled ML application (loaded in both core and shell), once every 1 or 2 weeks for 6 weeks. FIG. 5F shows quantification of serum levels of free fatty acids (mean ± SD, n = 4). Statistical comparison between groups was performed using one-way ANOVA. * p < 0.05 vs. control.

[0095] FIG. 5G illustrates that present ML approach improves browning and metabolic effects of CL316,243 in obese mice. Diet-induced obese mice were treated with ML without drug (control), or with CL316,243 (CL, ~1 mg/kg/week) via IP injection or core-shelled ML application (loaded in both core and shell), once every 1 or 2 weeks for 6 weeks. FIG. 5G shows quantification of serum levels of insulin free fatty acids (mean ± SD, n = 4). Statistical comparison between groups was performed using one way ANOVA. * p < 0.05 vs. control.

[0096] FIG. 5H shows the in vitro biocompatibility of CMC-coated salt-mixed PLGA- MLs (50% salt) tested on human white adipocytes. Briefly, human pre-adipose cells (Zen-Bio Inc., USA; SP-F-2) collected from subcutaneous WAT of non-diabetic human subjects (BMI 25 - 29.9) were grown till confluence in DMEM/ F-12 containing 10% FBS. The cells (defined as day 0) were then treated for 7 days with serum- free differentiation medium (DMEM/F-12 containing 0.5 pg/ml insulin, 0.5 mM IBMX, 100 nM dexamethasone, 100 nM rosiglitazone, 33 mM biotin, 17 pM pantothenate, 10 pg/ml transferrin and 1 nM triiodothyronine). The cells were then cultured for another 7 days with serum-free growth medium (DMEM/F-12 containing insulin and dexamethasone). At day 14, white adipocytes cultured in 96 wells were exposed to ML for 2 weeks. Representative bright field images of adipocytes with lipid droplets are shown. Oil Red O staining confirmed the formation of lipid droplets in white adipocytes (inset).

[0097] FIG. 51 shows the statistics of cell viability (alamarBlue assay) (%control; mean ± SD, n = 5).

[0098] FIG. 5J shows i/7 vitro bioactivity of mirabegron released from salt-mixed PLGA-MLs. Drug released from MLs or freshly prepared drug (~1 pM, lpg/2.5 ml) was used to treat white adipocytes for ~7 days before conducting the immunoblot experiment. The representative immunoblot images of uncoupling protein 1 (UCP1) and actin are shown.

[0099] FIG. 5K shows the statistics (mean ± SD, n = 3; normalized to actin density) are shown. Student’s t-test: ***p <0.001 vs. no treatment.

[00100] FIG. 6A shows the results of diet-induced obese mice were treated ML without drug (control), or with CL316,243 and rosiglitazone (combined dosage of 1 mg/kg/week) via subcutaneous (SC) injection or ML application, once every 2 weeks for 6 weeks. MLs loaded with rosiglitazone in core (Ro-c) and CL316,243 in shell (CL s) or MLs loaded with rosiglitazone in shell (Ro-s) and CL316,243 in core (CL-c) were used in the experiments. The relative body weight increases of differently treated mice. Data represents mean ± SD (n=4 or 5).

[00101] FIG. 6B shows the statistical analyses (mean ± SD, n = 4) for IgWAT, EpiWAT and interscapular BAT. [00102] FIG. 6C shows the representative images of IgWAT, EpiWAT and interscapular BAT.

[00103] FIG. 6D shows the representative hematoxylin and eosin staining images of IgWATs. Scale bar denotes 50 pm.

[00104] FIG. 6E shows the quantification of serum levels of cholesterol (mean ± SD, n = 4).

[00105] FIG. 6F shows the quantification of serum levels of triglyceride (mean ± SD, n = 4).

[00106] FIG. 6G shows the quantification of serum levels of free fatty acids (mean ± SD, n = 4). [00107] FIG. 6H shows the quantification of serum levels of insulin (mean ± SD, n =

4).

[00108] FIG. 61 shows the intraperitoneal glucose tolerance test at week5 (mean ± SD, n = 4). Glucose levels from non-obese mice (normal diet) and mice treated with ML (CL-c+CL-s) (CL316,243 in both core and shell) are used for comparison. Statistical comparison between groups was performed using one-way ANOVA. * p < 0.05 vs. control. [00109] FIG. 6J shows the intraperitoneal glucose tolerance test at week5 (mean ± SD, n = 4). Glucose levels from non-obese mice (normal diet) and mice treated with ML (CL-c+CL-s) (CL316,243 in both core and shell) are used for comparison. Statistical comparison between groups was performed using one-way ANOVA. * p < 0.05 vs. control.

[00110] FIG. 6K shows a bright field image, and their corresponding histological and fluorescence images of mouse IgWATs collected after 1 week. Scale bar denotes 250 pm.

[00111] FIG. 6L shows a bright field image, and their corresponding histological and fluorescence images of mouse IgWATs collected after 1 week. Scale bar denotes 250 pm.

[00112] FIG. 6M shows a bright field image, and their corresponding histological and fluorescence images of mouse IgWATs collected after 1 week. Scale bar denotes 250 pm. [00113] FIG. 6N shows a bright field image, and their corresponding histological and fluorescence images of mouse IgWATs collected after 6 weeks post-application of Cy5- loaded salt- mixed PLGA-MLs (50% salt). Scale bar denotes 250 pm.

[00114] FIG. 60 shows a bright field image, and their corresponding histological and fluorescence images of mouse IgWATs collected after 6 weeks post-application of Cy5- loaded salt- mixed PLGA-MLs (50% salt). Scale bar denotes 250 pm.

[00115] FIG. 6P shows a bright field image, and their corresponding histological and fluorescence images of mouse IgWATs collected after 6 weeks post-application of Cy5- loaded salt- mixed PLGA-MLs (50% salt). Scale bar denotes 250 pm.

[00116] FIG. 7A shows the average daily food intake of obese mice treated without (Control) or with CL316,243 (1 mg/kg/week) via IP injection or ML application (CL316,243 loaded in both core and shell), once every 1 week (1 wk) or 2 weeks (2 wk) for 6 weeks. Sample size of n = 4.

[00117] FIG. 7B shows the average daily food intake of obese mice treated without (Control) or with CL316,243 +rosiglitazone (combined dosage of 1 mg/kg/week) via subcutaneous (SC) injection or ML application, once every 2 weeks for 6 weeks. MLs differently loaded with rosiglitazone in core/shell (Ro-c/s) and/or CL316,243 in core/shell (CL-c/s) were used (Sample size of n = 4). [00118] FIG. 7C shows the average body surface temperatures (mean ± SD, n = 4). [00119] FIG. 7D depicts the adipose insulin resistance index for the fasting serum insulin (IR = insulin x FFA) (mean ± SD, n = 4). Statistical comparison among groups was performed using one-way ANOVA. * p < 0.05, ** p < 0.01 vs. Control (no treatment).

[00120] FIG. 7E depicts the adipose insulin resistance index for fatty acids product (IR = insulin x FFA) (mean ± SD, n = 4). Statistical comparison among groups was performed using one-way ANOVA. * p < 0.05, ** p < 0.01 vs. Control (no treatment). [00121] FIG. 8A demonstrates for the in vivo study of present core-shelled MLs for diet-induced obese mice that were treated without (Control) or with CL316,243

(lmg/kg per week) via ML application once every 2 weeks for 6 weeks. CL316,243 were loaded in either ML’s core or shell (CL-c or CL-s). FIG. 8A shows relative body weight increases over the course of treatment (means ± SD, n = 4).

[00122] FIG. 8B demonstrates for the in vivo study of present core-shelled MLs for diet-induced obese mice that were treated without (Control) or with CL316,243

(lmg/kg per week) via ML application once every 2 weeks for 6 weeks. CL316,243 were loaded in either ML’s core or shell (CL-c or CL-s). FIG. 8B shows the statistical analyses (mean ± SD, n = 4) of IgWAT, EpiWAT and interscapular BAT (relative to the respective body weight). [00123] FIG. 8C demonstrates for the in vivo study of present core-shelled MLs for diet- induced obese mice that were treated without (Control) or with CL316,243

(lmg/kg per week) via ML application once every 2 weeks for 6 weeks. CL316,243 were loaded in either ML’s core or shell (CL-c or CL-s). FIG. 8C shows representative hematoxylin and eosin staining images of IgWATs. Scale bar denotes 50 pm. [00124] FIG. 8D demonstrates for the in vivo study of present core-shelled MLs for diet- induced obese mice that were treated without (Control) or with CL316,243

(lmg/kg per week) via ML application once every 2 weeks for 6 weeks. CL316,243 were loaded in either ML’s core or shell (CL-c or CL-s). FIG. 8D shows quantification of serum levels of cholesterol (mean ± SD, n = 4). [00125] FIG. 8E demonstrates for the in vivo study of present core-shelled MLs for diet- induced obese mice that were treated without (Control) or with CL316,243

(lmg/kg per week) via ML application once every 2 weeks for 6 weeks. CL316,243 were loaded in either ML’s core or shell (CL-c or CL-s). FIG. 8D shows quantification of serum levels of triglyceride (mean ± SD, n = 4).

[00126] FIG. 8F demonstrates for the in vivo study of present core-shelled MLs for diet- induced obese mice that were treated without (Control) or with CL316,243

(lmg/kg per week) via ML application once every 2 weeks for 6 weeks. CL316,243 were loaded in either ML’s core or shell (CL-c or CL-s). FIG. 8D shows quantification of serum levels of free fatty acids (mean ± SD, n = 4).

[00127] FIG. 8G demonstrates for the in vivo study of present core-shelled MLs for diet- induced obese mice that were treated without (Control) or with CL316,243

(lmg/kg per week) via ML application once every 2 weeks for 6 weeks. CL316,243 were loaded in either ML’s core or shell (CL-c or CL-s). FIG. 8D shows quantification of serum levels of insulin (mean ± SD, n = 4).

[00128] FIG. 8H shows intraperitoneal glucose tolerance test at week5. Data shows mean ± SD (n = 4). Statistical comparison among groups was performed using one-way ANOVA. * p < 0.05, ** p < 0.01 vs. Control (no treatment).

[00129] FIG. 9 depicts core-shelled ML coated with crosslinked methacrylated- hyaluronic acids (MeHA). Confocal images of real-time releases of Cy5 (red, in ML core) and FITC (green, in ML shell) fluorescence dyes from salt-mixed PLGA-ML (50% salt) and crosslinked-MeHA shell, respectively, in 4% agarose hydrogel. Scale bars = 250 pm. MeHA was synthesized by functionalizing HA (-300 kDa) with methacrylic anhydride. After dip-coating with MeHA solution (50 mg/ml MeHA together with 0.5 mg/ml Irgacure 2959 in deionized water) and drying, the core-shelled ML was exposed to low-intensity ultraviolet light (5 min) to crosslink MeHA. As shown, crosslinked-MEHA can be successfully coated on ML (-250 pm diameter, -4 mm length, with a sharp-pointed tip of -30 pm). FITC dye (green) loaded in crosslinked-MEHA shell attained much slower release (-2 - 3 days) when compared to the fast-release profile of carboxymethyl cellulose (CMC) shell (within minutes) (also see FIG. 2A to 2C).

[00130] FIG. 10 shows, in the left image, a representative bright field image of hollow micro-lances (MLs) made of Poly (D, L-lactide-co-glycolide) (PLGA, 50:50, Mw 54,000-69,000, loaded with Cy5 fluorescence dye). Scale bar denotes 1 mm. The right image is a plot of the mechanical compression test of a PLGA-based hollow micro- lance. As shown in the left image, the fabricated PLGA-based hollow MLs are long cylindrical in shape with the diameter of -250 pm and length of -4 mm. The ML has a sharp-pointed tip at one end, with the diameter of -30 pm and a single bevel angle of -75°, and a flat-blunted one at the other end. The central hollow channel has the diameter of -100 pm. PLGA-based hollow MLs could attain the mechanical force of more than 1 N per needle (stiffness of -3.5 N/mm) without breaking or deformation (Supplementary Figurelb). It indicates that those hollow MLs are strong enough for skin penetration, as the force required to penetrate the human skin is -150 - 300 mN/needle for sharp-tipped microneedles (-40 pm tip diameter).

[00131] FIG. 11 demonstrates hyaluronic acids-based MLs for fast drug release kinetic. Fast-dissolving low-molecular weight hyaluronic acids (HA) (miniHA, 3 - 10 kDa molecular weight; 1000 mg/ml) was used to make ML by thermal pressing molding method. HA-ML can be successfully fabricated (-250 pm diameter, -4 mm length, with a sharp-pointed tip of -30 pm). Those HA-MLs are strong enough to penetrate porcine skin and provide very fast releasing kinetic (within minutes).

[00132] FIG. 12A demonstrates fast-dissolving polymeric MLs. Bright field images of MLs, made of polyvinyl alcohol (PVA; 9 kDa) and propylene glycol (PG) by thermal pressing molding method are shown. As shown, fast-dissolving polymeric ML can be successfully fabricated (-250 pm diameter, -4 mm length, with a sharp-pointed tip of -30 pm). PVA/PG-MLs are strong enough to penetrate porcine skin (FIG. 12C) and provide fast releasing kinetic (within minutes).

[00133] FIG. 12B demonstrates fast-dissolving polymeric MLs. Bright field images of MLs, made of polyvinylpyrrolidone (PVP), 10 kDa molecular weight by thermal pressing molding method are shown. As shown, fast-dissolving polymeric ML can be successfully fabricated (-250 pm diameter, -4 mm length, with a sharp-pointed tip of -30 pm). PVP-MLs are strong enough to penetrate porcine skin (FIG. 12D) and provide fast releasing kinetic (within minutes).

[00134] FIG. 12C is a bright field image of porcine tissue (skin - epidermis and dermis, and subcutaneous white adipose tissue, sWAT) immediately after lancing a ML made of polyvinyl alcohol (PVA) and propylene glycol (PG).

[00135] FIG. 12D is a bright field image of porcine tissue (skin and sWAT) immediately after lancing a ML made of polyvinylpyrrolidone (PVP). [00136] FIG. 13 A shows PLGA-based gutter-like MLs acting as a cell delivery device. Specifically, FIG. 13A shows the bright field and corresponding image of the gutter like ML loaded with cells. FIG. 13 A shows the ML. Cells were stained with NucBlue™ Live ReadyProbes™ Reagent (Hoechst 33342, a nucleic acid stain) (ThermoFisher Scientific). PLGA-based gutter-like MLs were surface-modified with poly-L-lysine (PLL) coating before cell loading. Scale bar denotes 50 pm.

[00137] FIG. 13B shows PLGA-based gutter-like micro-lances acting as a cell delivery device. Specifically, FIG. 13B shows the bright field and corresponding image of the gutter-like ML loaded with cells. FIG. 13B shows the ML tip. Cells were stained with NucBlue™ Live ReadyProbes™ Reagent (Hoechst 33342, a nucleic acid stain) (ThermoFisher Scientific). PLGA-based gutter- like MLs were surface-modified with poly-L-lysine (PLL) coating before cell loading. Scale bar denotes 50 pm.

[00138] FIG. 13C shows PLGA-based gutter-like micro-lances acting as a cell delivery device. Specifically, FIG. 13C shows the bright field and corresponding image of cell- loaded gutter-like ML inserted in a 4% agarose hydrogel. FIG. 13C shows ML. Cells were stained with NucBlue™ Live ReadyProbes™ Reagent (Hoechst 33342)

(ThermoFisher Scientific). PLGA-based gutter-like micro-lances were surface- modified with poly-L-lysine (PLL) coating before cell loading. Scale bar denotes 50 pm.

[00139] FIG. 13D shows PLGA-based gutter-like micro-lances acting as a cell delivery device. Specifically, FIG. 13D shows the bright field and corresponding image of cell- loaded gutter-like ML inserted in a 4% agarose hydrogel. FIG. 13D shows ML tip. Cells were stained with NucBlue™ Live ReadyProbes™ Reagent (Hoechst 33342)

[00140] FIG. 14 illustrates different types of MLs, such as a solid ML (having a solid core), a hollow ML (having a hollow core), and a gutter-like ML (a gutter-like core). Detailed Description

[00141] The following detailed description refers to the accompanying drawings that show, by way of illustration, specific details and embodiments in which the present disclosure may be practised.

[00142] 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.

[00143] The present disclosure relates to a self-administrable and implantable polymeric micro-implant for controlled and localized therapeutics delivery into the tissues safely and effectively. The therapeutics may include or may be a drug, a cell, any bioactive agent, or a mixture thereof. The micro-implant may be termed herein a therapeutics delivery device or micro-lance (ML), as the micro-implant is shaped like a lance having a sharp-pointed end for penetrating biological tissue. The sharp-pointed lance-shaped micro-implant (micro-lance) can be easily and comfortably injected into the skin or other tissues weekly or regularly by the patient at home without pain and need of skills (“patient-friendly”). In addition, with the help of a lancing-device or applicator, the micro-lance can quickly be pierced into the skin or tissue (within a fraction of second) and embedded in the deep tissue (e.g. subcutaneous tissue, sclera) in a minimally-invasive and painless manner (i.e. comfort and convenient). Advantageously, there is no need to visit a clinic, or medical practitioner, for invasive methods (e.g. surgical incision, hyponeedle or trocar or cannula insertion, etc.) to implant the micro-lances. The embedded micro-lance is operable as long-acting micro reservoirs in tissues for “controlled and targeted therapeutics delivery”. As the micro lances are biodegradable, there is advantageously no need for them to be manually removed from the biological tissues or body.

[00144] The therapeutics delivery device may be configured as a polymeric micro lance having a core and optionally one or more shells forming an outer layer coating the core. For instances, the therapeutics delivery device may be configured as a polymeric core-shell micro-lance, with the outer coating layer of fast-dissolving polymer and the inner core of slow-dissolving polymer, in order to achieve several advantages: (1) fast-dissolving polymer (e.g. CMC) is used to make the coating-layer of micro-lance (shell) for quick delivery of bioactive agents within minutes while the core of micro-lance made up of slow-dissolving polymer (e.g. PLGA) is for slow and sustained release of bioactive agents over several weeks to months (up to 3-6 months), (2) the sustained drug release kinetic (several weeks to months) engineered through the polymeric core, e.g. mixture of different concentration of porogens, combination or mixture of different polymers, etc., (3) a PLGA-based micro-lance core is a non limiting example that can provide a strong mechanical strength to pierce the skin and implant into the underlying subcutaneous tissue, (4) the micro-lance implanted into the subcutaneous fat or tissue can then serve as the long-acting micro-dmg-reservoir which slowly discharges its cargo. The achieved controlled drug release, including the fast- release from outer coating-shell and sustained-release from core is superior to delivery devices that has only fast or slow releasing drug delivery platform, and in terms of treatment efficacy and efficiency. Due to the continuously drug releasing capability of the present implanted micro-lances, low therapeutic dose and prolonged dosing interval can be realized. Lowering the therapeutic dose not only produces lesser side-effects but also reduces the cost. Lengthening the dosing interval greatly enhances the patient compliance.

[00145] The therapeutics delivery device, wherein the therapeutics may contain a cell, may be configured as a polymeric hollow micro-lance, wherein a hollow channel lay inside the center of the core (hollow core). The polymeric micro-lance may also have a gutter-like core wherein a central hollow channel is laterally open through the lateral wall of the core. The present polymeric hollow micro-lance or gutter-like micro-lance for cell delivery or replacement therapy (e.g. islet cells, stem cells, neurons, melanocytes, dendritic cells, T lymphocytes, or T-cells) etc., provides several advantages: (1) the central cavity or channel of the micro-lance is designed to create local micro-environment of transplanted cells, and thus, cells can be largely insulated from the hostile host environment, and at the same time, easily accessible to the nutrient and gas exchange, (2) the lateral wall of the micro-lance can behave as a drug- containing material that can improve cell survival, and enhance cell function, (3) the lateral wall of the micro-lance can also be capable of releasing the immunosuppressant drug, and thus, through sustained local immunosuppression, the host may be considered as the foreign transplanted cells as self, achieving long term toleration of transplanted cells, (4) co-delivery of both therapeutic cells and bio-active agents is possible, with appropriate cells and drug ratio, to allow for synergistic therapeutic effect. As an example, co-transplantation of insulin (loading in polymeric material) and beta-islet cells (loading in hydrogel) for the management of diabetes patients (or glucagon and alpha-islet cells for hypoglycaemia) is possible with the present therapeutics delivery device. The lateral wall of the micro-lance can be porous (e.g. pores each having 0.5 - 5 pm diameter pore size) to allow better nutrient and gas exchange between inner transplanted cell environment and outer recipient tissue environment.

[00146] Moreover, as the fabrication process of the present therapeutics delivery device is based on a straightforward moulding method, the micro-lance fabrication is simple, inexpensive, and suitable for mass production.

[00147] A biocompatible biopolymer, for example PLGA, can be chosen as the polymeric carrier for the forming the micro-lance, which is highly biocompatible, biodegradable, and inexpensive, rendering safe and cost-effective long-term usage of the present therapeutics delivery device. The polymeric carrier may include more than one polymer species or each species may have a range of molecular weights.

[00148] In addition, the present disclosure also provides for a lancing device or applicator operable to achieve several advantages: (1) a strong spring is included in the applicator to offer fast lancing speed and short lancing duration (millisecond range) for efficient and pain-free micro-lance injection into the tissue, (2) by inspiring the buckling prevention- strategy in nature, a rigid column (a stainless-steel tube) is included in the applicator to laterally support the micro-lance during injection into the tissue. The lateral support greatly enhances the critical buckling load of micro-lance, which in turn reduce or resist buckling or breaking during insertion. Briefly, the custom- made high-speed lancing device includes of a spring and a shaft connected with a stainless-steel rod which is well-fitted into a stainless- steel tube, into which a micro lance can be seamlessly loaded. Once the spring is relieved, the force transferred through the rod can quickly push the micro-lance towards the tissue. In addition to the usage of a spring for rapid and pain-free ML injection, other methods such as pneumatic pressure or pyro-drive produced by a high-pressure compressed air or gas can also be applied for lancing (i.e. driving) the micro-lance out of the applicator and into a biological tissue to be implanted.

[00149] Details of various embodiments of the present therapeutics delivery device, present applicator, method of producing the therapeutics delivery device, their applications, and advantages associated with the various embodiments are now described below. Where advantages of the various embodiments have been described in the examples section, they shall not be iterated for brevity.

[00150] In the present disclosure, there is provided a therapeutics delivery device implantable in a biological tissue. The therapeutics delivery device may include a core which is biodegradable, wherein the core includes a first polymer composition, wherein the core has a height of at least 1 mm (e.g. 1 - 30 mm, 1 - 10 mm, etc.), and wherein the core has one end shaped as a sharp tip sufficient to penetrate the biological tissue, wherein the sharp tip includes a single bevel-angled shape, a double bevel-angled shape, a chiseled edge, a bullet-shape, or a frustum- shape. The therapeutics delivery device herein exchangeably refers to a micro-lance described in the present disclosure. In various embodiments, multiple micro-lances can be used and implanted simultaneously. For example, the micro-lance may be configured as an array having a plurality of such micro-lances. The micro-lance may have a solid core (solid micro lance) or hollow core (hollow micro-lance). The solid core and hollow core are capable of housing a therapeutics. The core may be coated with one or more shells.

[00151] In various embodiments, the therapeutics delivery device may further include a shell which is biodegradable, wherein the shell includes a second polymer composition, wherein the first polymer composition and the second polymer composition have different degradation rate. That is to say, there is provided a therapeutics delivery device implantable in a biological tissue. The therapeutics delivery device may include a core and a shell both of which are biodegradable. The core may include a first polymer composition and the shell may include a second polymer composition, wherein the first polymer composition and the second polymer composition have different degradation rate. The core may be coated with one or more of such shells. Where multiple shells are present, each of the shells may have a different degradation rate. The core may have a height of at least 1 mm (e.g. 1 - 30 mm, 1 - 10 mm, etc.), and the core may have one end shaped as a sharp tip sufficient to penetrate the biological tissue, wherein the sharp tip comprises a single bevel-angled shape, a double bevel-angled shape, a chiseled edge, a bullet-shape, or a frustum-shape. A frustum shape refers to a shape that has one end broader than the other end, wherein the cross-section from one end tapers down (decreases in size) toward the other end. Non limiting examples include fmsto-conical shape and a pyramidal frustum shape.

[00152] In various embodiments, the core may be a solid core, a hollow core, or a gutter-like core.

[00153] In various embodiments, the first polymer composition can contain one or more polymers. The first polymer composition may include, without being limited to, poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly(lactic acid-co-glycolic acid) (PLGA), polyanhydride, polyorthoester, polyetherester, polycaprolactone (PCL), polyesteramide, poly(butyric acid), poly(valeric acid), polyvinylpyrrolidone (PVP), polyvinyl alcohol (PVA), polyethylene glycol (PEG), a block copolymer of PEG-PLA, PEG-PLA-PEG, PLA-PEG-PLA, PEG-PLGA, PEG-PLGA-PEG, PLGA-PEG-PLGA, PEG-PCL, PEG-PCL-PEG, PCL-PEG-PCL, a copolymer of ethylene glycol-propylene glycol-ethylene glycol (PEG-PPG-PEG), hyaluronic acid, a derivative thereof, or a mixture thereof. The polymers for the first polymer composition may have a variety of molecular weights. The poly(lactic acid-co-glycolic acid) may include poly(D-lactic- co-glycolic acid), poly(L-lactic-co-glycolic acid), and/or poly(D,L-lactic-co-glycolic acid).

[00154] In various embodiments, the first polymer composition may be formed from two monomers, wherein the two monomers may include lactic acid and glycolic acid, wherein the two monomers are present in a molar ratio of 1:100 to 100:1.

[00155] In various embodiments, there may be more than one shell. In other words, the therapeutic delivery device may have multiple layers of shell coating the core. Each of the shell may include the second polymer composition. The second polymer composition can contain one or more polymers. The second polymer composition may include, without being limited to, hyaluronic acid and its derivatives, methacrylate hyaluronic acid, sodium alginate, collagen and its derivatives, polyurethane, poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly(lactic acid-co-glycolic acid) (PLGA), polyanhydride, polyorthoester, polyetherester, polycaprolactone (PCL), polyesteramide, poly(butyric acid), poly(valeric acid), polyvinylpyrrolidone (PVP), polyvinyl alcohol (PVA), polyethylene glycol (PEG), a block copolymer of PEG-PLA, PEG-PLA-PEG, PLA-PEG-PLA, PEG-PLGA, PEG-PLGA-PEG, PLGA-PEG-PLGA, PEG-PCL, PEG-PCL-PEG, PCL-PEG-PCL, a copolymer of ethylene glycol-propylene glycol-ethylene glycol (PEG-PPG-PEG), dextran, hetastarch, tetrastarch, pentastarch, hydroxyethyl starch, cellulose, hydroxypropyl cellulose (HPC), sodium carboxymethyl cellulose (Na CMC), thermosensitive HPMC (hydroxypropyl methyl cellulose), polyphosphazene hydroxyethyl cellulose (HEC), other polysaccharide, polyalcohol, gelatin, alginate, chitosan, sucrose, or a mixture thereof. The polymers for the second polymer composition may have a variety of molecular weights. The poly(lactic acid- co-glycolic acid) may include poly(D-lactic-co-glycolic acid), poly(L-lactic-co- glycolic acid), and/or poly(D,L-lactic-co-glycolic acid).

[00156] In various embodiments, the core and/or shell may further include a porogen. The choices of porogen may include, without being limited to, sodium chloride (NaCl), polyethylene glycol, an ionic liquid, or a mixture thereof. The ionic liquid may include a pyridinium salt or an imidazolium salt. Non-limiting examples of an imidazolium salt may include or may be 1 -butyl-3 -methyl-pyridinium dicyanamide (Bmp-dca), 1-ethyl- 3-methyl-imidazolium dicyanamide (Emim-dca), 1 -ethyl-3 -methyl-imidazolium tetracyanoborate (Emim-tcb), 1-heptyl-pyridinium tetrafluoroborate, etc.

[00157] The use of porogens (e.g. salt microparticles) can be added into the therapeutics delivery device (in core, shell, or both) to accelerate degradation thereof. The core and/or shell may be made from one or more polymers. For example, both the core and the shell may be made by a mixture of different polymers tailored specially therefor. This is easily achieved by the method of producing the therapeutic delivery device described herein, such as simply using a polymer solution having a mixture of polymers to make the core and/or shell.

[00158] In various embodiments, the porogen may be present in amount of less than 75 wt% of (i) the first polymer composition forming the core and/or (ii) the second polymer composition forming the shell. Advantageously, this renders the core sufficiently flexible for penetration into the biological and not too brittle such that the core breaks when penetrating the biological tissue. [00159] In various embodiments, the core may have a cross-sectional diameter ranging from 200 mhi to 600 mhi, 300 mhi to 600 mhi, 400 mhi to 600 mhi, 500 mhi to 600 mhi, etc.

[00160] In various embodiments, the core may further include a reservoir which houses one or more therapeutics, or the first polymer composition may house one or more therapeutics. In various embodiments, the shell may include one or more therapeutics. The one or more therapeutics housed in the core and the shell may be same or different. In various embodiments, the core may be a solid core or a hollow core. The solid core may house one or more therapeutics therein. For example, the first polymer composition used to form the solid core may house one or more therapeutics. For the hollow core configuration, the core may have a hollow reservoir defined by a polymer matrix. In certain instances, the first polymer composition used to form a polymer matrix defining the hollow core may house one or more therapeutics, or both the polymer matrix and the hollow core may house one or more therapeutics. In other words, the core may further have a polymer matrix defining a hollow reservoir, wherein either or both the polymer matrix and hollow reservoir may each house one or more therapeutics. For the hollow core configuration, the one or more therapeutics housed in the polymer matrix and the hollow core, and even the shell, may be same or different. The one or more therapeutics may include or may be a drug, a cell or a bioactive agent. Non-exclusive examples of the therapeutics may include an adipose browning agent, a lipolytic agent, an anti-obesity agent, a blood sugar lowing agent, an insulin sensitizing agent, an anti cancer agent, an anti-psychotic agent, an anti-fungal agent, an anti-microbial agent, an anti-psychotic agent, an anesthetic agent, a pro-angiogenic agent, an anti- angiogenic agent, a hair-growth promoting agent, a hormonal agent, a protein or peptide, an amino acid, an antibody or antibody fragment, an oligonucleotide, a cell or cell fragment, an exosome or microvesicle, or a combination thereof. The one or more therapeutics may be in the form of nanoparticles, nanospheres, nanocapsules, nanodots, nanorods, or a combination thereof. The one or more therapeutics may be organic or inorganic. Non exclusive examples of organic nanoparticles may includes indocyanine green, heptamethine cyanine, cryptocyanine, phthalocyanine, perylene diimide, porphyrin and its derivatives, polyaniline, poly(BIBDF-BT), polypyrrole and its derivatives, diketopyrrolopyrrole, dopamine, melanin, croconaine, squaraine, benzobisthiadiazole, or a combination thereof. Non-exclusive examples of inorganic nanoparticles may contain gold, tungsten, copper, silicon, copper sulfide, molybdenum, iron oxide, manganese dioxide, silicon dioxide, carbon-based materials, or a combination thereof. [00161] The present dislosure also provides for an applicator operable to implant the therapeutics delivery device as described herein in various embodiments of the first aspect. The applicator may include a lancing module and one or more holder tube, wherein each of the one or more holder tubes is configured for loading the therapeutics device therein, and wherein the lancing module is operable to drive the therapeutics delivery device out of the one or more holder tubes to implant the therapeutics delivery device described according to various embodiments of the first aspect into the biological tissue. Where the therapeutics delivery device is configured or operated to be in the form of multiple micro-lances, the applicator may be used to implant the multiple micro-lances simultaneously.

[00162] The lancing module herein refers to any mechanism that can lance the therapeutics delivery device loaded in the applicator into a biological tissue.

[00163] In certain embodiments, the lancing module may include a spring coupled to one or more rods and the one or more holder tubes. In other words, the applicator may include a housing that includes a spring coupled to one or more rods and one or more holder tubes, wherein each of the one or more holder tubes may be configured for loading the therapeutics delivery device therein, and wherein the spring is operable to have the one or more rods drive the one or more holder tubes to implant the therapeutics delivery device into the biological tissue. In certain embodiments, each of the one or more rods may be a hard object of any shape which is operable to strike and propel the therapeutics delivery device out of the applicator and into the biological tissue. The one or more rods may differ in shapes.

[00164] In certain embodiments, the lancing module may contain an air jet which lances or ejects the therapeutics delivery device into the biological tissue. Any other suitable mechanisms capable of lancing the therapeutics delivery device into the biological tissue may be adopted in the applicator.

[00165] Embodiments and advantages described for the present therapeutics delivery device of the first aspect can be analogously valid for the present applicator subsequently described herein, and vice versa. As the various embodiments and advantages have already been described above and examples demonstrated herein, they shall not be iterated for brevity.

[00166] In various embodiments, the applicator can be designed to launch an array of the therapeutics delivery devices. By doing this, more therapeutics can be delivered and a larger administration area can be covered. To achieve this, an example is to have the housing of the applicator configured to have multiple rod-tube combinations in order to load multiple therapeutics delivery devices. The housing can alternatively have multiple holder tubes, wherein each of them is loaded with one therapeutics delivery device and all the therapeutics delivery device can be lanced or driven by one rod or a hard object with a different shape.

[00167] In various embodiments, the housing, the lancing module, the rod, and/or the one or more holder tubes may be formed of stainless steel. The stainless steel may be medical or food grade.

[00168] The present disclosure also provides for a method of producing the therapeutics delivery device described in various embodiments of the first aspect. Embodiments and advantages described for the present therapeutics delivery device of the first aspect and the applicator can be analogously valid for the present method subsequently described herein, and vice versa. As the various embodiments and advantages have already been described above and examples demonstrated herein, they shall not be iterated for brevity.

[00169] The method may include providing a mold to form the core, wherein the mold includes a cavity having a depth of at least 1 mm and a closed end comprising a single bevel-angled shape, a double bevel-angled shape, a chiseled edge, a bullet-shape, or a frustum-shape, depositing a first polymer solution into the cavity, wherein the first polymer solution includes the first polymer composition, and drying the first polymer solution to form the core.

[00170] In various embodiments, the method may further include coating a second polymer solution on the core to form the shell, wherein the second polymer solution includes the second polymer composition. That is to say, the present method may include providing a mold to form the core, wherein the mold includes a cavity having a depth of at least 1 mm (e.g. 1 - 30 mm, 1 - 10 mm) and a closed end including a single bevel-angled shape, a double bevel-angled shape, a chiseled edge, a bullet-shape, or a frustum-shape, depositing a first polymer solution into the cavity, wherein the first polymer solution includes the first polymer composition, drying the first polymer solution to form the core, and coating a second polymer solution on the core to form the shell, wherein the second polymer solution includes the second polymer composition.

[00171] In various embodiments, the mold may be formed of a polymer material (e.g. polydimethylsiloxane, polyurethane), a ceramic material, or a metallic material (e.g. stainless steel, nickel, copper or gold). Such a mold offers advantage of a strong mechanical stability to resist deformation during a molding process. For faster micro lance drying process. The mold may be porous or/and contains porogens (e.g. polyethylene glycol) or superabsorbent polymers (e.g. a poly-acrylic acid sodium salt). The mold, or portion of it, may be subjected to surface treatments to make it more hydrophilic or hydrophobic, which make it easier for filling polymeric solution into the mold, for example, by using radiofrequency or plasma treatment, or by coating with a surfactant such as polysorbate, docusate, sodium salt, benzethonium chloride, alkyltrimethylammonium bromide or hexadecyl trimethyl ammonium bromide (CTAB).

[00172] In various embodiments, depositing the first polymer solution may include dissolving the first polymer composition in an organic solvent to form the first polymer solution, then introducing the first polymer solution into the mold, or may include heating a first polymer composite to a high temperature (e.g. 50 - 200°C), and applying a pressure on the first polymer solution or the first polymer composite to introduce it into the mold, wherein the pressure may be a hydraulic pressure ranging from 10 to 500 pounds per square inch (psi). In this instance, the polymer composite refers to a mixture of polymers that are not dissolved in an organic solvent. Heating the first polymer composite may include melting the first polymer composite. Heating the polymers in the polymer composite above their glass transition temperature (Tg; e.g. 50 to 200 °C) renders the polymers and the polymer composite to be in a mobile rubbery state and thus moldable into the mold, without the need for a solvent and avoids damaging the therapeutics and/or the mold. Applying a pressure (e.g. 10 - 500 psi) pushes the polymer into the mold, without the need for a solvent and not adversely affecting the shape and structural integrity of the mold. [00173] In various embodiments, depositing the first polymer solution may include mixing the porogen and/or the one or more therapeutics in the first polymer solution. [00174] The present method may further include heating the mold after depositing the first polymer solution and prior to drying the first polymer solution, wherein heating the mold may include heating the mold to a temperature ranging from 50 to 200°C without damaging the therapeutics and/or the mold.

[00175] In various embodiments, the method may further include mixing the one or more therapeutics in the second polymer solution, removing the core from the mold prior to coating the second polymer solution, and immersing the core into the second polymer solution to form the shell.

[00176] The present disclosure may further provide for a method of treating a medical condition or delivering a therapeutics. Embodiments and advantages described for the present therapeutics delivery device of the first aspect, the applicator and the method of producing the therapeutics delivery device can be analogously valid for the present method of of treating a medical condition or delivering a therapeutics subsequently described herein, and vice versa. As the various embodiments and advantages have already been described above and examples demonstrated herein, they shall not be iterated for brevity.

[00177] The present method may include operating the applicator to implant the therapeutics delivery device in the biological tissue. The medical condition may include obesity, a metabolic disease, a cancer, an eye disease, a skin disease, alopecia, or schizophrenia.

[00178] In summary, herein provided is a multi-modal controlled drug (and cell) delivery system which can be easily and comfortably self-administered at home, to directly implant into the deep subcutaneous fat or tissue, and thus patient friendly and well- suited for long-term home-based management of many diseases.

[00179] The present therapeutics delivery device is operable as a self-administrable and implantable drug delivery device for controlled and targeted delivery. The therapeutics delivery device is operable with an applicator which allows for the therapeutics delivery device (i.e. micro-lance) to be released at high speed. The applicator, as described above, can include a spring and a shaft connected with a rod fitted into a housing, and a dissolvable lanced-shaped needle (ML) (solid, hollow or gutter-like) in the housing that has a porous core, a sharp tip, a shell coating layer that may be made of caboxymethylcellulose sodium (CMC) on the porous core and sharp tip, and therapeutics loaded in both the core and coating. In certain embodiments, the fabricated MLs (solid, hollow or gutter-like) are long cylindrical in shape with the diameter of -250 pm and length of -4 mm. The MLs has a sharp-pointed tip at one end, with a diameter of -30 pm and a single bevel angle of -75°, and a flat-blunted one at the other end. The central hollow channel has the diameter of -100 pm.

[00180] Advantageously, the housing of the applicator helps resist buckling during skin insertion of the needle by providing a rigid column for supporting the needle. The needle is released at a high speed to reduce pain.

[00181] Advantageously, the therapeutics delivery device allows the solid, hollow or gutter-like MLs to penetrate beyond the skin epidermis and dermis, and embedded deep into the underlying subcutaneous fat, with the tilted position, -2-6 mm from the skin’s stratum comeum.

[00182] 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 present disclosure. [00183] 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.

[00184] In the context of various embodiments, the punctuation

the term “about” or “approximately” as applied to a numeric value encompasses the exact value and a reasonable variance.

[00185] As used herein, the term “and/or” includes any and all combinations of one or more of the associated listed items.

[00186] 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. Examples

[00187] The present disclosure relates to a self-administrable and implantable therapeutics delivery device into the biological tissues in a safe and minimally-invasive manner for controlled release of the therapeutics, and a method for preparing the same. Some features of the device and method are discussed below. The therapeutics may include or may be a drug, a cell, or a bioactive agent.

[00188] The therapeutics delivery device is suitable for controlled release of therapeutics and may comprise a lance-shaped micro-implant (hereinafter, generically referred to as micro-lance) and operable with an applicator, wherein the micro-lance is made of a biocompatible and biodegradable polymer material, has a lance-shaped structure with a core and a sharp tip, is loaded with one or more biological active agent, is self-administrable and implantable into the biological tissues in a minimally-invasive manner, and the applicator is the spring-based applicator for lancing the micro-lance to pierce into the biological tissues. A plurality of micro-lances (i.e. therapeutics delivery device) can be configured as an array for implantation into a tissue.

[00189] The biocompatible and biodegradable polymer may for example be poly(D- lactic-co-glycolic acid), poly(L-lactic-co-glycolic acid), poly(D,L-lactic-co-glycolic acid) and their derivatives (hereinafter, generically referred to as poly(lactic-co-glycolic acid or PLGA), with a molecular weight from 1,000 Da to 100,000 Da, and a molar ratio of lactic acid and glycolic acid from 1:100 to 100:1. Other suitable and demonstrated polymers have been described in the detailed description section. The mixtures or composites of same (e.g. PLGA) or different polymers (e.g. polyethylene glyco) (PEG), polyvinyl alcohol (PVA), polyvinylpyrrolidone (PVP), hyaluronic acid (HA) are also usable in order to control the mechanical strength as well as the dissolution rate and release kinetic. The micro-lances can contain a different polymer, ceramic, drug, additive, or any combination or mixture thereof.

[00190] The micro-lance has a core, that can be either solid, hollow or gutter-like, and configured to have any cross-sectional shape (e.g. rounded, hexagonal, rectangular, etc.), with a diameter of 200-600 pm and the a height of 1 - 30 mm (e.g. 1 - 10 mm). The central hollow cavity or channel, with a diameter of 50-400 pm and a height of 1 - 30 mm (e.g. 1 - 10 mm), may contain different polymer, hydrogel, cryogel, nanoparticles, aqueous solution or any combination or mixture thereof, loading with different bioactive agents, proteins, peptides, oligonucleotides, cells or any combination or mixture thereof.

[00191] The polymeric micro-lance may include one or more porogen in the core for making porous micro-lance in the biological tissues, wherein the porogen, with a percentage of 1% up to 90 wt%, is any of the biocompatible fast-dissolving pore forming material or particles, preferably the inorganic compounds such as sodium chloride (NaCl), etc., polymeric material such as polyethylene glycol (PEG), and ionic liquids such as pyridinium or imidazolium salts.

[00192] The polymeric micro-lance may include one or more coating layer (outer shell), wherein the coating layer is made of any of the biocompatible fast-dissolving viscosity-enhancing coating solution, preferably carboxymethylcellulose sodium (CMC), sucrose, HA, methacrylated HA (MeHA), sodium alginate, poly-vinyl- pyrrolidone (PVP), and one or more bioactive agent is loaded into the coating layers. Other suitable and demonstrated materials for the shell have been described above in the detailed description section. The coating layers can contain a different polymer, ceramic, drug, additive, or any combination or mixture thereof.

[00193] The micro-lance may have a sharp-pointed tip, preferably a single or double bevel-angled shape, or pyramidal or conical in shape, with a diameter of 5-50 pm. [00194] The applicator may include one or more arrays of blunted or rounded hollow microchannel or tubes for loading and laterally- supporting the micro-lances, and the spring connected with one or arrays of rods perfectly fitted inside the microchannel or tubes for pushing, driving and directing the micro-lances toward the targeted tissues. [00195] The implantation site is any of the biological tissues, preferably the skin, subcutaneous tissue, subcutaneous fat, tumor, sclera, vitreous cavity, joint cavity, etc. [00196] The bioactive agents could be any drugs or compounds that can be used in obesity, metabolic diseases, tumor mass (malignant tumor such as melanoma, as well as benign tumor such as lipoma) and other diseases (e.g. skin diseases, alopecia). Non limiting examples of such bio-active agents may include adipose browning agents (e.g. b3 -adrenergic receptor agonist), lipolytic agents (e.g. deoxycholic acid), blood sugar lowing agents (e.g., PPAR agonist), anti-cancer agents (e.g. tamoxifen), anti-psychotic agents (e.g. risperidone), anti-fungal agents (e.g. amphotericin), anti- angiogenic agents (e.g. anti-VEGF agents), hair-growth promoting agents (e.g. minoxidil) and synthetic hormonal agents (e.g. levonorgestrel). The percentage of bioactive agent loaded, embedded or encapsulated in each micro-lance is from 1% up to 50% w/w.

[00197] The present therapeutics delivery device is operable for treating, preventing, ameliorating, reducing or delaying the onset of obesity, metabolic diseases, cancers, eye diseases, and other diseases. Non-limiting examples of such diseases may include obesity, metabolic diseases (e.g. type 2 diabetics, hypertriglyceridemia, and hypercholesterolemia), malignant tumor (e.g. breast cancer, prostate cancer and melanoma), benign tumor (e.g. lipoma), eye diseases (e.g. diabetes retinopathy, age- related macular degeneration), skin diseases (alopecia and deep skin infection, etc) and other diseases (e.g. schizophrenia). The device is also suitable for use in other conditions such as hormone replacement therapy, contraception and family planning, therapeutic cell delivery, immunization, etc.

[00198] The micro-lance with hollow cavity or channel filling with hydrogel or cryogel is operable for cell delivery or replacement therapy (e.g. islet cells, stem cells, neurons, melanocytes), cell-based vaccination and immunotherapy (e.g. dendritic cells), etc., in order to treat, prevent, ameliorate, reduce or delay the onset of diabetes, metabolic diseases, cancers and other diseases, or to promote skin re -pigmentation, hair- follicle growth, nerve regeneration, re- vascularization, etc.

[00199] The preparation method of polymeric micro-lance may be characterized as follows.

[00200] PLGA (20-50% w/v) (or other polymer) and bioactive agents are dissolved in a solvent (e.g., dimethylformamide, acetone, etc.) to obtain the polymer-drug solution. After complete evaporating the solvent, the drug-embedded polymeric matrix or rubber is filled into the micro-lance shaped mold cavities made of polydimethylsiloxane (PDMS). The mold cavities can be made of PDMS alone or PDMS mixed with sodium polyacrylate (20 wt.%). The molding process can be done by directly pressing the matrix at 50- 200 °C, using a hydraulic pressure (10 - 500 pounds per square inch). In some cases, porogens (e.g. salt microparticles) are thoroughly mixed with polymeric matrix before loading into the mold cavities, to make porogen-mixed polymeric micro lances. In some cases, PDMS -mold cavity may contain central filament, threadlike object or fiber, and thus, the fabricated micro-lance could have hollow cavity or channel. A filament, threadlike object or fiber can also be inserted into the polymeric matrix or rubber before the molding process, to make the hollow micro-lances. After air-drying at room temperature (e.g. 20 - 28 °C) or other temperature (e.g. 40 °C), micro-lances are peeled off from the mold cavities. In some cases, micro-lances can be coated with the coating solution which may also contain bioactive agents. In some cases, the hydrogel or cryogel containing cells or other bioactive agents can be loaded into the hollow cavity of micro-lance by using a syringe or centrifugation. The micro lances are then stored in the air-tight containers at the room temperature or 4 °C fridge. [00201] The present therapeutics delivery device, the applicator, method of producing the therapeutics delivery device, and their applications/uses thereof, are described in further details, by way of non-limiting examples, as set forth below.

[00202] Example 1A: Materials Used

[00203] Poly (D, L-lactide-co-glycolide) (PLGA, 50:50 ratio, Mw 54,000-69,000), carboxymethylcellulose (CMC), poloxamer 188 (Pluronic F68), sodium chloride (NaCl), polydimethylsiloxane (PDMS, SYLGARD 184) and dimethylsulfoxide (DMSO) were purchased from Sigma- Aldrich. CL316,243 (a selective p3-adrenergic receptor agonist), rosiglitazone (a selective PPARy agonist) and insulin (human recombinant) were obtained from Tocris Bioscience. Cyanine5 (Cy5) and fluorescein were acquired from Lumiprobe. DMEM/F12, fetal bovine serum (FBS), bovine serum albumin (BSA), phosphate buffered saline (PBS), penicillin- streptomycin, trypsin- EDTA and AlamarBlue cell viability reagent were obtained from ThermoFisher Scientific. All other reagents and solvents were also purchased from Sigma- Aldrich and used without further purification.

[00204] Example IB: Fabrication of Polymeric Micro-Lances [00205] In general, microlances (MLs) were prepared via a simple hot-embossing micromolding method. Briefly, PLGA and drug or fluorescence molecules were dissolved in a dimethylformamide solution to obtain the polymer-drug solution. After complete evaporating the solvent, the drug-embedded PLGA matrix or rubber was thoroughly mixed with the salt microparticles (~30 pm). The salt- mixed PLGA matrix was then casted into the PDMS micromolds, the reverse replica of the stainless- steel master- molds (33G Lancet, with the diameter of 0.25 mm and the height of 4 mm), by applying a ~50-pound per square inch load at 80°C for 30 min. After air-drying at room temperature (RT) (e.g. 20 - 28 °C) in a fume hood (~ 12 hr), polymeric MLs were peeled off from the micromolds. The micro-lances were coated with the CMC-based coating solution, with or without containing the drug or fluorescent molecules. The MLs were then stored in the air-tight containers at the RT. The salt- mixed PLGA-based MLs (PLGA/NaCl ratio of 50:50) were used in all the experiments unless otherwise indicated.

[00206] In on example, MLs were prepared via a simple thermal pressing method. Firstly, NaCl micro-particles were obtained by grinding the powder and filtering through a fine mesh with a pore size of ~50 pm. The mixture of PLGA and salt micro particles was then prepared using the solvent casting method. Briefly, PLGA was dissolved in acetone before salt micro-particles were added. After thorough mixing and complete evaporation of the solvent at room temperature (e.g. 20 - 28 °C), a homogeneous solid mixture was obtained. Rosiglitazone or fluorescent dyes were added directly into the salt-polymer solution. For hydrophilic molecules like CL316,243 and mirabegron, the deionized (DI) water solution of the drug was thoroughly mixed with acetone solution of PLGA. Then salt micro-particles were added, following by complete evaporation of solvent and water. The solid mixture was then placed on top of the PDMS mold, which is the reverse replica of a home-made stainless-steel master-mold having a vertical array of 32G ultrafine needle (diameter of 0.23 mm and height of 4 mm), before heating at ~90 °C for 10 mins and pressing with a compression force of ~20 kg for 2 hrs. PDMS mold, filled with dmg-salt-PLGA mixture, was then heated for another 15 mins without pressing to allow the PDMS mold to restore to its original shape. After cooling at room temperature (e.g. 20 - 28 °C) for 30 mins, polymeric MLs were individually taken out using forceps. Subsequently, MLs were gradually dipped into a 20 pi DI water solution containing 0.6 mg CMC and 0.2 mg Pluronic F68 and dried in air for 15 mins to form CMC-PLGA core-shelled MLs. The coating process was repeated for 3 times.

[00207] Example 1C: Characterization

[00208] The morphology of MLs was examined using a field-emission scanning electron microscope (FESEM; JSM-6700, JEOL) and a digital microscope (Leica DVM6). MLs loaded with different fluorescence dyes were visualized with a confocal laser scanning microscope (LSM800, Carl Zeiss). Drug loading amount in ML was determined using a UV-Vis spectroscopy (Shimadzu UV-1800). The mechanical property of MLs was tested using an Instron 5543 Tensile Tester. Briefly, a vertical force was applied onto a ML using a flat-headed stainless- steel cylindrical probe at a constant speed of 0.5 mm/min and the force exerted on the MN was continuously recorded. ML insertion test was performed on the isolated porcine skin as previously described. Briefly, MLs were mounted onto the cylindrical probe, and pressed perpendicular to the isolated porcine skin at a rate of 5 mm/min until a pre-set maximum load of 4 N was reached. Force exerted on the skin by the ML as a function of its displacement into skin was recorded. The insertion force was estimated when the force against the skin showed discontinuity followed by a steep slope.

[00209] To examine in vitro biocompatibility of MLs, white adipocytes differentiated from the primary human pre-adipose cells (Zen-Bio Inc.) were exposed to MLs for 2 weeks, before analyzing cell morphology using an inverted microscope (1X71, Olympus, equipped with a digital camera OlympusE330) and cytotoxicity using an almarBlue cell viability assay. To evaluate in vitro bioactivity of drug molecules loaded in MLs, these adipocytes were treated with free CL316,243 or CL316,243 released from MLs (1 pg/ml) for 7 days. Cells were then lysed with M-PER Mammalian Protein Extraction Reagent (containing Halt protease and phosphatase inhibitor cocktail, ThermoFisher Scientific) for further immunoblot experiments to test protein expression levels of UCP1, which reflects the adipose browning activity of C1316,2543.

[00210] To study the in vitro insertion capability, MLs were administered vertically into the porcine skin and its underlying subcutaneous fat tissue using a home-made lancing applicator. The applicator contains a spring and a shaft connected with a stainless-steel rod which is well-fitted into a stainless- steel holder tube. The spring is made of a -0.8 mm stainless-steel wire, and has an outer-diameter of 9 mm, a length of 40 mm, an effective coil-number of 10, a k value of -0.09 kgf/mm, and a loading force of -1.4 kgf. The porcine tissue was then excised and photographed to analyze the embedded MLs inside the tissue (insertion depth and trajectory). For histological analyses, the tissues were fixed with 4% paraformaldehyde solution for 24 hours, and cryoprotected with 30% sucrose solution for 1-2 days, before embedding in FSC22 Frozen Section Media (Leica Microsystem). The tissue sections (10 pm) were sliced using a cryostat (CM 1950 cryostat, Leica Microsystems) before staining with the hematoxylin and eosin solutions (Sigma- Aldrich), and images were taken using a digital microscope (Leica DVM6). To further evaluate the in vitro insertion trajectory, MLs were inserted vertically into the agarose hydrogel (4% w/v, in DI water, pH 7.4) using the applicator. The bright-field images of the agarose hydrogel with the embedded MLs were then analyzed using the ImageJ software (NIH.gov).

[00211] To evaluate the release profiles, MLs differently loaded with CL316,243 or Cy5 molecules were immersed in the phosphate buffer solution (PBS, pH 7.4), and placed in an incubator shaker (100 rpm, 37 °C). Fluorescent molecules released from MLs were determined using a fluorescence spectrometer (SpectraMax M5, Molecular Devices). CL316,243 molecules released from MLs were measured by a gradient reverse-phase high performance liquid chromatography (Agilent 1100 HPLC-DAD system) using an Agilent Poroshell 120 EC-C18 column (with a mobile phase of water and methanol, a flow rate of 0.5 mL/min and UV detection at 285 nm) and quantified based on the linear calibration curve created with different known concentrations. The real-time visualization of fluorescent molecules releases from MLs in agarose hydrogel (4% w/v, in DI water, pH 7.4) were also analyzed using a confocal microscopy. [00212] The in vivo fluorescence imaging of fluorescent molecules release from MLs was conducted on mice (C57BL/6J, 7 - 8-week-old male). Briefly, after shaving the mouse hairs around the inguinal region (lower left or right quadrant of dorsolateral area, adjacent to the hind limbs), MLs differently loaded with Cy5 molecules (1 pg in the PLGA-core or CMC- shell) were applied obliquely (~45 ⁰ angle from the skin surface) on the skin around the inguinal region using the applicator. The mice were then imaged immediately (dayO) or at day 1 or day7 using an in vivo imaging system (IVIS Spectrum, Perkin Elmer). To evaluate the in vivo bio-distribution of molecules released from MLs, mice treated with either ML insertion (1 pg Cy5 in the PLGA-core or CMC-shell) or intraperitoneal (IP) injection (1 pg in 50 pi PBS) were euthanized at 2-hour post treatment or at day7, and the white adipose tissue (WATs) and other major organs (liver, heart, kidney and lung) were dissected and visualized by IVIS imaging system.

[00213] Example ID: Animal Experiments (In Vivo Studies of Polymeric Microlances)

[00214] All animal experiments were approved by Institutional Animal Care and Use Committee of Nanyang Technological University under protocol ARF-A18029. The mice (C57BL/6J) were housed in light and temperature-controlled facility (12-hr light/12-hr dark cycle, 22 °C), and allowed free access of water and standard or high- fat diet. After shaving the hairs around the inguinal region, a ML was applied using the applicator. Low-dose inhaled isoflurane was used to constrain the mice during ML application. Mice were then returned to their cages and imaged immediately to access the penetration sites. After ~5 mins, mice were recovered from anaesthesia and their behaviours were monitored to access their possible pain using Grimace scoring based on orbital tightening, nose bulge, check bulge, and ear position. The body weight and food intake were also recorded (day 1, 4 and 7). In some tested mice groups, mice were euthanized immediately after ML insertion or at dayl or day30, and skin tissue and IgWATs were collected to examine the histological changes and biocompatibility of MLs. The in vivo fluorescence from Cy5 molecules released from MLs was imaged using an in vivo imaging system (IVIS Spectrum, Perkin Elmer).

[00215] Diet-induced obese mouse model for localized drug delivery of micro-lances, e.g. mice (6-7 weeks old male) fed on high-fat diet (60% kcal from fat, Testdiet) for 3 weeks to induce obesity, were randomly divided into 5 groups for different treatments. Each ML contains 10 pg CL316,243 in the core and 5 pg CL316,243 in the shell. One ML was applied at each left and right inguinal region once every 1 week, or 2 MLs at each region once every 2 weeks (both equivalent to 1 mg/kg/week).

[00216] Lour to five mice were used in each treatment or control group, and mice were fed with high-fat diet throughout the experiments. After 5 weeks treatment, the glucose tolerance test was performed as described previously. Briefly, the overnight-fasted mice (16 hrs) were injected with glucose solution (2 g per kg in PBS) via IP route, and blood glucose was monitored over time. After 6 weeks treatment, body surface temperature was monitored on the shaved skin under fully awake condition, using infrared thermal imaging camera (LLIR T420). The mice were then euthanized by a lethal dose of carbon dioxide. After measuring the body weights, fat tissues (inguinal WAT, epididymal WAT and interscapular BAT) were excised, weighted and collected for histological analyses. Some portions of IgWATs were homogenized in radioimmunoprecipitation assay buffer (containing protease inhibitor cocktail, Roche Applied Science) for immunoblot analyses. Briefly, tissue samples with equal amount of proteins (as the loading control) were separated on 12% SDS-PAGE before being transferred onto a nitrocellulose membrane. The membrane was then blocked with Superblock blocking buffer (ThermoFisher Scientific) (2 hr at RT) and incubated with specific primary antibody (1: 200-400 dilutions, 12 hr at RT), before washing (Tris-buffered saline- Tween solution- TBST, 3 x 15 min each) and incubation with horseradish peroxidase- conjugated secondary antibody (Sigma- Aldrich) (1: 2000 - 4000; 6 hr at RT). After washing again with TBST (3 x 15 min), the protein bands were detected in a G: BOX Chemi XT4 imaging system (Syngene) using SuperSignal WestPico Chemiluminescent Substrate (ThermoFisher Scientific). Antibodies against uncoupling protein- 1 (UCP1) (PA1-24894) and PRDM16 (PA5-20872) were also obtained from ThermoFisher Scientific. Antibodies against adipocyte protein-2 (aP2, sc- 18661) (fatty acid binding protein 4), peroxisome proliferator-activated receptor gamma coactivator 1-a (PGC-la, sc-13067), and actin (sc-1616) were purchased from Santa Cruz Biotechnology. [00217] Example IE: Histological Analysis

[00218] The excised tissues were immediately fixed with 4% paraformaldehyde solution for 24 hrs, washed with PBS, and immersed in 30% sucrose solution for 2 days to cryoprotect the tissues. After embedding in FSC22 Frozen Section Media (Leica Microsystem), tissues were frozen and sliced (10-20 pm thick) using a cryostat (CM1950 cryostat, Leica Microsystems). The tissue sections were finally stained with hematoxylin and eosin solutions (Sigma- Aldrich), and images were taken using a digital microscope (Leica DVM6).

[00219] Example IF : Immunoblot Analysis

[00220] Cell or tissue samples with equal amount of proteins (as the loading control) were separated on 12% SDS-PAGE before being transferred onto a nitrocellulose membrane. The membrane was then blocked with Superblock blocking buffer (ThermoFisher Scientific) (2 hrs, at room temperature) and incubated with specific primary antibody (1:200-400 dilutions, 12 h at room temperature), before washing with Tris-buffered saline-Tween solution (TBST, 3 x 15 mins) and incubation with horseradish peroxidase-conjugated secondary antibody (Sigma- Aldrich) (1:2000- 4000; 6 hrs, at room temperature). After washing again with TBST (3 x 15 mins), the protein bands were detected in a G: BOX Chemi XT4 imaging system (Syngene) using SuperSignal WestPico Chemiluminescent Substrate (ThermoFisher Scientific). Antibodies against UCP1 (sc-6529), peroxisome proliferator-activated receptor gamma coactivator 1-a (PGC-la, sc-13067), adipocyte protein-2 (aP2, sc- 18661) (fatty acid binding protein 4), and actin (sc-1616) were purchased from Santa Cruz Biotechnology. Antibodies against UCP1 (PA1-24894) and PRDM16 (PA5-20872) were also obtained from ThermoFisher Scientific.

[00221] Example 1G: Assessment on serological parameters [00222] Serum levels of cholesterol, triglycerides, free fatty acids, and glucose were measured by the standard assay kits from Sigma-Aldrich. Serum level of insulin was determined by a mouse insulin ELISA kit from ThermoFisher Scientific.

[00223] Example 1H: Statistical Analysis

[00224] The data points were presented as mean ± standard deviation (SD). Statistical analyses were performed using one-way analysis of variance (ANOVA) followed by Turkey’s post-hoc test. Unless otherwise stated, all experiments were performed using at least four different samples per group, and the data presented are representative of two to three independent experiments. A p value of < 0.05 was considered statistically significant.

[00225] Example 2A: Discussion on Fabrication of Core-Shelled Micro-Lances [00226] In general, lance-shaped polymeric micro-implants (termed herein micro lances, MLs) were fabricated using a simple hot-embossing micro-molding method (FIG. 1H). The biodegradable and biocompatible polymer poly (lactic-co-glycolic acid) (PLGA) was used as the polymeric matrix of MLs (core), because of its strong mechanical strength and typical long-term drug releasing capability through the slow degradation process. Another biodegradable and biocompatible polymer, carboxymethyl cellulose (CMC) was used for the coating layer of MLs (shell), as the water-soluble fast-dissolving nature of CMC offers the fast release of its cargo. PLGA and CMC are also widely used in many US-FDA approved medical and food products, respectively. Combining the merits of slow-releasing PLGA and fast-releasing CMC, herein developed is the core- shell MLs for localized and controlled drug delivery into subcutaneous tissue. To speed up the release kinetics of PLGA (from several months to several weeks), salt particles as porogens were mixed in PLGA (50:50) to realize the porous structure of MLs in the tissues, as the leaching of salt particles by tissue interstitial fluid could create pores, and hence accelerate the degradation rate and release kinetic of PLGA. Briefly, salt-mixed PLGA, with or without therapeutic compounds, were casted into the mold by applying a load of ~50 pounds per square inch at ~80 °C for 1 hour. It was known that glass transition temperature (Tg) of PLGA is ~55 °C. However, because of the addition of salt which subsequently increases the Tg of polymer-salt complex, higher molding temperature (~80 °C) was used to soften and easily pack the polymer-salt complex into the mold. As the Tg of salt-mixed PLGA is above room temperature (RT), the fabricated ML behaves as a hard and rigid structure at RT. MLs were then coated with the CMC solution, with or without therapeutics, by simply using a dip-coating method (FIG. 1H).

[00227] As revealed by the scanning electron microscopy (SEM) and optical microscopy, the fabricated MLs are long cylindrical in shape with the diameter of -250 pm and length of -4 mm (FIG. 2A and FIG. IB). The ML has a sharp-pointed tip at one end, with the diameter of ~10 pm and a single bevel angle of -75°, and a flat-blunted one at the other end (FIG. IB). The ML design is based on the previous findings that sharp-tapered and thinner needles cause less pain, skin penetration force and trauma comparing to large and obtuse needles. The average pore diameter of the porous PLGA- based MLs after leaching of the salt particles is ~4 pm, with the maximum pore size of ~10 pm (FIG. ID to 1G). The confocal fluorescence imaging confirms that different fluorescence molecules, as the model therapeutic compounds, can be differently loaded in the inner PLGA core and outer coating layer (FIG. 2A). Consistent with the well- known biocompatibilities of PLGA and CMC polymers, the fabricated MLs are highly biocompatible to human cells, as evidenced by the preserved morphology and viability of human white adipocytes (FIG. 5H and 51). The bioactivity of drug molecules released from MLs was also tested by their adipose browning activity on human white adipocytes (FIG. 5J and 5K) and found that drug loading in MLs did not affect their biological activity. With the both fast and sustained releasing ability, as well as biocompatibility, these dissolving MLs implantation into the tissue are suitable for localized and controlled drug delivery. The outer-shell of CMC being exposed to tissue fluid can quickly discharge its cargo, whereas the inner-core of salt-mixed PLGA becomes porous micro-drug-reservoir which slowly degrades and releases its cargo over time.

[00228] To demonstrate on the above, one example is discussed in more detail. In this example, polymeric micro-lances (MLs) were fabricated by a simple thermal pressing method (exchangeably herein termed “hot-embossing”) using poly (lactic-co-glycolic acid) (PLGA), because the thermoplastic property of PLGA makes it moldable at a mobile rubbery state above the glass transition temperature (Tg = 55 °C) and then becomes strong and rigid below Tg. PLGA, which has good biocompatibility and high mechanical strength, is widely used in US-FDA approved drug products and implants. The biodegradation rate of PLGA can be tailored by varying the monomer ratio (PLA/PGA). PLGA with 50:50 PLA/PGA monomer ratio exhibits the fastest hydrolytic degradation at the physiological condition. To further speed up the degradation, salt micro-particles (10 - 50 pm, FIG. 1A) as porogens are added to PLGA polymer matrix. As the melting temperature of most small drug molecules is much higher than Tg of PLGA, they can be incorporated into MLs by simply mixing with PLGA molecules. [00229] As illustrated in FIG. 1H, a dry homogeneous mixture of PLGA molecules, drug compounds, and NaCl micro-particles was first placed on top of a negative mold made of polydimethylsiloxane (PDMS), and subsequently heated at 90 °C for 10 mins to make it rubbery and pressed into the mold with a compression force of ~20 kg for 2 hrs. After cooling at room temperature (e.g. 20 - 28 °C) for 30 mins, PLGA-MLs were demolded. Such thermal pressing method is simple, fast and inexpensive, and amenable for automated massive production. Moreover, to realize biphasic release of one kind of drug molecules or release two different kinds of drug molecules at distinct rates, a fast dissolving and biocompatible layer of carboxymethyl cellulose (CMC) was coated onto a PLGA-ML simply by dip-coating to produce a core-shelled ML. In fact, CMC is a commonly used additive in foods and beverages. As revealed by optical microscopy, the fabricated MLs are cylindrical with a diameter of -0.23 mm, length of -4 mm (FIG. 2A and 2B) and a sharp-pointed tip -10 pm (FIG. IB). Such dimensions are optimized based on the balanced consideration of drug loading capacity, penetration ability, and invasiveness. These MLs have similar dimensions as stainless-steel needles or lancets (32 gauge) which are widely used at home for insulin injection or blood sampling and considered virtually painless.

[00230] Example 2B: Present Micro-Lances for Bi-phasic Drug Delivery and Strong Mechanical Strength

[00231] As shown in FIG. 2A, confocal fluorescence imaging confirms the ability of the MLs to pack different fluorescence molecules as the model drug molecules (red cy5 in inner PLGA core, green fluorescein in outer CMC shell). To examine in vitro release kinetics, MLs differently loaded with fluorescence dyes were inserted into agarose hydrogel as the skin mimics, and continuously monitored under confocal microscopy. Immediately after insertion, fast-dissolving outer-shell of CMC quickly released its cargo within 5 mins (FIG. 2B and 1C), while the salt-mixed PLGA inner-core (50% salt) gradually degraded, and discharged its cargo over several weeks (4 - 6 week, FIG. 2B). As revealed by scanning electron microscopy (SEM) (FIG. ID to 1G), PLGA MLs became porous after dissolution of NaCl micro-particles, and gradually degraded over time. Using high performance liquid chromatography (HPLC), in vitro drug release kinetics was further tested by monitoring the release profile of a browning agent (CL316,243 - a b3 -adrenergic receptor agonist) encapsulated in inner-core or outer- shell or both in PBS solution. As shown in FIG. 2C, release from the shell is fast while release from the inner core has the time constant of ~12 days based on exponential curve fitting. In comparison, drug release from PLGA ML without the salt porogens exhibits a much slower release time constant of ~27 days. More specifically, -15% was released from PLGA-only ML while -50% was released from PLGA ML containing 50% salt in 2 weeks. At week 6-8, salt-mixed MLs were totally disintegrated into highly-porous microspheres (FIG. 2E) which then degraded into its two monomers (lactic and glycolic acids) and metabolized into carbon dioxide and water via the tricarboxylic acid cycle. Taken together, the present core-shelled MLs provide bi- phasic release kinetics, i.e. the initial burst release followed by the sustained release over several weeks.

[00232] As ML needs to be strong enough to penetrate through skin and implant entirely into underlying sWAT, the ML’s mechanical strength was assessed by compression test. As shown in FIG. 2D, PLGA MLs with or without 50% salt micro particles can sustain a compressive force of -0.9 N, without significant deformation. It means that these MLs are strong enough for skin penetration because the force required for a sharp needle (-30 pm tip) to penetrate human skin is less 0.1 N. Addition of 50% salt micro-particles does not compromise the mechanical strength significantly. Salt- blended ML may be more brittle because it breaks beyond the critical compression load (-1.5 N) instead of elastic bending as pure PLGA MLs (FIG. 3J and 3K). Nevertheless, higher salt percentage accelerates ML degradation and drug release (FIG. 2F and 2G). FIG. 3L and 3M show the corresponding mechanical strength performance. To elaborate, for example, MLs with 75% salt may be too brittle for skin penetration and suitable for applications that require such properties. Additional CMC coating does not reinforce the mechanical strength (FIG. 3N), indicating that the mechanical property is dictated by the PLGA inner core. In addition, drug loading (10%) does not significantly compromise the mechanical properties of MLs (FIG. 30).

[00233] Example 2C: Lancing MLs into sWAT

[00234] Mechanical failure of a needle during skin penetration is often caused by lateral bending, which may be a typical mechanical failure mode for the needles, particularly the polymeric ones, during insertion into the skin is buckling, the lateral bending when the axial compressive force exceeds the critical load of the needles. However, the presence of lateral support can largely increase the critical buckling load. For example, the lateral support of mosquito’s labium (a protective sheath surrounding the fascicle) increases the critical load of fascicle (a microneedle made of chitin) by a factor of 5. Moreover, a straight fascicle can penetrate more easily. As such, ML can be laterally supported by a rigid tube to resist buckling during skin insertion. This can also alleviate any possible fracture due to compressive force during insertion. On the other hand, in order to reduce the penetration duration and hence discomfort, the speed of a piercing ML should be fast. To meet these two requirements, a high-speed lancing- device or applicator can be included, which involves a spring and a shaft connected with a stainless-steel rod. The rod is fitted into a stainless-steel tube holder, into which a ML can be snugly loaded (FIG. 41). Once the compressive force of the spring is released, the rod instantly strikes the ML, lancing ML swiftly into skin. The lancing speed is ~ 1 m/s, and duration is only ~50 ms (FIG. 4J). These values are comparable to those achieved by the lancet lancing devices used by diabetes patients at home for blood sampling. The lancing speed, thus the skin penetration depth, can be adjusted by varying the spring compression.

[00235] Using the applicator, a core- shelled ML can be easily lanced into skin- mimicking agarose gel (4%), without buckling or fracture (FIG. 4K). The distance between the penetration point on the surface and the blunted end of ML is 2-3 mm (greater than the thickness of human skin of ~2 mm at abdominal area), suggesting that MLs can be embedded deep into sWAT. Notably, the embedded ML is titled -22°, because of the resistive force from the tissue acted on the bevelled tip of the ML (FIG. 4K to 4N). This is consistent with the observation that the insertion trajectory of bevel- tipped needles is curved. Porcine skin closely resembles human skin and is the widely used skin model. Similar to human sWAT, porcine sWAT is continuously associated with dermal layer. As demonstrated in FIG. 3A, ML can be easily lanced into sWAT (tip reaches 5 - 6 mm deep from the skin surface). Subsequent histological study shows the narrow insertion path, and the implanted ML (FIG. 3B). The insertion force for ML upon its penetration into the porcine skin is ~0.1 N (FIG. 3P) as indicated by a small dip in the force-displacement curve followed by a steeper slope. Taken together, these experiments demonstrate that MLs are strong enough to pierce the skin and reach into the subcutaneous fat, where the implanted MLs can function as the micro-drug- reservoirs for the localized and controlled drug release.

[00236] Example 2D: Safe, Painless and Minimally Invasive Implantation of MLs as Embedded Drug Reservoirs

[00237] The safety and drug releases of core-shell MLs were further investigated in vivo mouse model. The MLs were inserted into the skin around the inguinal area using the applicator at the angle of -45° from the skin surface to embed the MLs within the underlying subcutaneous white adipose tissue (WAT; also called inguinal WAT). Inguinal WAT, which lies just 1 - 2 mm below the skin, is the largest subcutaneous WAT in rodents, and comparable in terms of location to the large gluteofemoral subcutaneous WAT in humans. The insertion site, body weight and food intake were monitored over 1 week.

[00238] Consistent with the well-known biocompatibility of PLGA and CMC polymers, the fabricated MLs are highly biocompatible to human cells, as evidenced by the well-preserved morphology and viability of human white adipocytes (FIG. 5H and 51). For in vivo experiments, MLs were administrated by the applicator at the inguinal area to implant MLs into the sWAT (inguinal WAT — IgWAT). IgWAT, which lies just 1 - 2 mm below the skin, is the largest sWAT in rodents. As shown in FIG. 3C, a tiny insertion spot/mark (-0.17 mm²) was observed immediately after insertion, and completely disappeared by day 2, without leaving any scar or other abnormalities (e.g. skin discoloration, swelling, bruising, hemorrhage or infection). The mark left by ML is similar to those (-0.15 mm²) created by stainless-steel needles of 32-gauge size. During 1 week of observation period, there were no significant differences in body weight and food intake between the control and tested groups (FIG. 3D and 3E). In addition, no sign of pain was observed from the mice administrated with ML, as assessed by the grimace scoring (FIG. 3F). The pain-free penetration is attributable to the sharp tip, small diameter, good mechanical strength, and fast lancing speed of MLs. No abnormal behavior (e.g. scratching due to itching) was noted in any of the tested mice. Furthermore, as shown in FIG. 3G, a ML penetrated into IgWAT and did not cause any noticeable tissue reactions (e.g. inflammatory reaction) at the implanting site. Previous studies demonstrated that even a large-sized PLGA implant (more than 1 cm) was highly inert in the body with no evidence of inflammatory infiltration. Similar to the degradation in the buffer solution (FIG. 2E), MLs were mostly degraded in vivo after 6 weeks (FIG. 6K to 6P). All these observations indicate that ML penetration and implantation is minimally-invasive and painless, without causing harmful effects on the targeted tissue and body.

[00239] In vivo distribution of released dye molecules from MLs was also examined. As shown in FIG. 3H, fluorescence signal of Cy5 encapsulated in the ML was strong at the penetration site. Fluorescence was absent in other areas and organs, e.g. liver, kidney, lung and heart. When Cy5 was loaded only in the shell, the fluorescence faded away within a day; when it was loaded in the core, the florescence remains strong even at day 7. Cy5 molecules released from either the shell or core only accumulated in the ML-embedded IgWAT, but not in the IgWAT on the other side or other organs (FIG. 31). In contrast, Cy5 molecules administrated systemically, i.e., intraperitoneal (IP) injection predominately accumulated in liver and kidney, but not in IgWAT. Hence, the micro-lance implantation allows targeted drug delivery into subcutaneous WAT, without passing through the systemic circulation, hepatic first pass metabolism, etc. At day 7, only the IgWAT embedded with a ML having a Cy5-loaded core still showed strong fluorescence. These experiments demonstrate that core-shelled ML can act as a drug reservoir inside sWAT to give highly localized biphasic drug release. Because of the localized and controlled drug releasing capabilities, micro-lanceimplantation would allow avoidance of large dose with frequent dosing-intervals required by conventional drug administration.

[00240] Example 2E: Gutter-Like MLs [00241] FIG. 13 A and 13B show, in the upper lane, representative bright field and corresponding confocal images of gutter-like ML made of poly(D, L-lactide-co- glycolide) (PLGA, 50:50, Mw 54,000-69,000), loaded with cells.

[00242] FIG. 13C and 13D show, in the lower lane, representative bright field and corresponding confocal images of cell-loaded PLGA-based gutter-like ML inserted in a 4% agarose hydrogel. Scale bar denotes 50 pm. As shown in the figures, the fabricated MLs are long cylindrical gutter-like in shape with the diameter of -250 pm and length of -4 mm. The ML has a sharp-pointed tip at one end, with the diameter of -30 pm and a single bevel angle of -75°, and a flat-blunted one at the other end. The central channel has a diameter of -150 pm, and provides a reservoir space for cell loading. The volume of the central channel was 0.05 to 0.1 mm³ (pL) (5 x 10⁶- 10 x 10⁷ pm³). Depending on the cell types (e.g. neutrophil - 300 pm³, beta cell - 1000 pm³, cardiomyocyte - 15,000 pm³), the maximal dosage of cells in each ML ranges from 6 x 10³ to 3 x 10⁵ cells. Such MLs attain the mechanical force of more than 1 N per needle (stiffness of -3.5 N/mm) without breaking or deformation. Such MLs are also strong enough to penetrate porcine skin. The porous PLGA-based lateral wall of the ML core (~ 0.5 - 5 pm pore size) allows nutrient and gas exchange to the surrounding environment. [00243] The surface modification and cell loading of PLGA-based gutter- like MLs are discussed below.

[00244] Coating with Poly-L-Lysine (PLL) - Poly-L-Lysine (PLL) is a biocompatible cationic polyamino acid that facilitates the attachment of cells and proteins to solid surfaces in biological application, hence promoting cell adhesion, proliferation and regeneration at the interface of the biomaterial. PLL demonstrates exceptional antimicrobial property, water solubility, stability and safety. PLL improves cell adhesion by enhancing the electrostatic interaction between the positive charges on the PLL surface and negative charges on the cell membrane surface. When adsorbed to the biomaterial, it increases the number of positively charged sites available for cell binding. This surface modification strategy is convenient, simple and displays great potential in biomaterial applications.

[00245] PLL coating was performed as followed. MLs (e.g. solid or hollow or gutter like MLs) were immersed in 0.1 M NaOH solution for 30 minutes at room temperature (e.g. 20 - 28 °C) to introduce -COOH groups onto the MLs’ surface. MLs were washed with distilled water five times to remove the remaining NaOH solution. The MLs were then immersed in PLL solution (P4707, Sigma Aldrich), with a concentration of 0.01% and molecular weight of 70,000 to 150,000, overnight at 4°C, before washing with distilled water to get rid of excess PLL solution.

[00246] Oxygen Plasma Treatment - Oxygen plasma treatment is another surface modification strategy to enhance cell adhesion. Gas plasma treatment is used for chemical modification of PLGA surface by creating reactive groups and increasing its hydrophilicity to improve cell adhesion. Besides hydrophilicity and surface chemistry, surface roughness may also influences cell spreading and growth as cells may orientate themselves along the grooves of a surface, a phenomenon known as contact guidance. Studies showed that oxygen plasma-treated samples had significantly higher cell attachment.

[00247] MLs (solid, hollow or gutter-like MLs) were placed in a petri dish and inserted into a plasma cleaner (Harrick Plasma) with high power setting (~45 W) and treated for different durations (e.g. 10 - 60 minutes).

[00248] Cell loading into MLs - Cells were loaded into MLs (solid, hollow or gutter like MLs) by simply mixing MLs with cell suspension solution before incubating overnight in the C02 incubator at 37 °C. In some cases, cells may be loaded with centrifugal forces (e.g. ~ 100 - 400 g) using a centrifuge.

[00249] Example 3A: Present Device and Treatment/Prevention of Diseases [00250] The present device and method provides a new transdermal delivery technique and demonstrated its use to directly deliver browning and insulin sensitizing agents into subcutaneous white adipose tissue (sWAT). Ultrathin, core-shelled, and lance-shaped polymeric drug reservoirs (micro-lances - MLs) were readily fabricated by thermal pressing molding and subsequent dip-coating. They can be rapidly lanced through the skin and totally implanted into sWAT for localized and biphasic drug release. The excellent therapeutic effectiveness to prevent development of obesity and associated metabolic diseases was demonstrated using an obese mouse model induced by high-fat diet. The present device provides a self-administrable and minimally-invasive ML approach, which can also be employed for long-term home-based treatment of other chronic diseases. [00251] For example, obesity is a serious epidemic health problem that can cause many other diseases including type 2 diabetes and cardiovascular diseases. Existing approaches to combat obesity suffer from low effectiveness and adverse side effects. The present device provides for a self-administrable and minimally-invasive transdermal drug delivery strategy for home-based long-term treatment of obesity and other diseases. Specifically, ultrathin, core-shelled, and lance-shaped polymeric drug- reservoirs (micro-lances - MLs) were readily fabricated by a thermal pressing molding method and totally implanted into subcutaneous fat by lancing through the skin. Using a diet- induced obese mouse model, development of obesity and associated metabolic disorders was shown to be effectively inhibited by applying therapeutic core- shelled MLs once every 2 weeks. The outstanding therapeutic effects can be attributable to highly localized and biphasic drug release, as well as combination therapy based on browning transformation of white fat and enhanced insulin sensitivity.

[00252] Obesity is increasingly prevalent worldwide and it may often be a root cause to many diseases such as diabetes, cardiovascular diseases, and cancers. Lifestyle intervention (diet and physical activity) alone may be insufficient to combat obesity because of problematic long-term patient compliance, body adaptation, and the fact that obesity may be a consequence of definite biological reasons. Although surgical intervention may be effective, there are concerns about risks, side-effects, high cost, and long-term efficacy. On the other hand, only a few weight management medications may be available, but they all act indirectly with limited efficacy, e.g. either by reducing fat absorption in gastrointestinal tract or suppressing appetite in brain. The adverse side- effects and high cost limit their use. Clearly, it becomes imperative to seek for new therapeutic strategy, which the present device addresses.

[00253] White adipose tissue (WAT) stores energy in the form of triglycerides. In obesity, WAT is excessively accumulated and releases various deleterious factors (e.g. free fatty acids- FFA, inflammatory cytokines, reactive oxygen species) which lead to various metabolic problems, such as insulin resistance. In contrast, brown adipose tissue (BAT) specializes in burning energy using glucose and FFA as the major fuels. Studies have demonstrated BAT as a negative regulator of adiposity and insulin resistance. Although BAT is scarce in adults and even less in obese people, brown-like (or beige) adipocytes have been identified in WAT, particularly in subcutaneous fat depot. Also, white adipocytes can transform into brown-like adipocytes (termed as browning) upon induction by some agents, (e.g. b3 -adrenergic receptor agonist). Therefore, stimulating browning of subcutaneous WAT (sWAT) may be more effective strategy to combat obesity and associated metabolic diseases. Nevertheless, the clinical use of browning or anti-obesity agents is prevented by the issues associated with the conventional drug delivery methods (e.g. oral intake, intravenous injection). Using these systemic delivery methods, the bioavailability at WAT is low due to poor GI absorption, hepatic first pass effect, renal clearance, enzymatic degradation and lack of ability to specifically target on WAT. Moreover, these methods tend to only give a burst release of drugs, leading to short therapeutic duration and hence the need of frequent administration. Consequently, the required high dosage and accumulation in non- targeted tissues unavoidably cause systemic side-effects. Not surprisingly, nanoparticle delivery of anti-obesity agents through the systemic route has significant off-target accumulation.

[00254] Microneedle skin patches for localized and controlled transdermal delivery of browning agents to sWAT have been demonstrated. Such a method offers outstanding anti-obesity efficacy with low effective dosage (thus presumably low side-effects) and minimal invasiveness (thus suitable for home-based healthcare). Despite its superiority over conventional methods, the microneedle approach still faces several limitations. Firstly, microneedle patches are typically fabricated using micro-molding method which is time-consuming and not easy to be scaled up due to involvement of centrifugation and lengthy drying process. Secondly, the length of microneedles is usually limited to be less than 1 mm in order to ensure painless penetration and possibility to demold microneedles without breaking. Such length is too short to reach sWAT. Therefore, drug molecules discharged from the microneedles embedded in dermis rely on diffusion into sWAT. Thirdly, loading capacity of microneedles is small. This problem tends to be worsened by two issues: (i) some molecules enter systemic circulation via dermal capillary networks before reaching sWAT and (ii) microneedles cannot fully insert into the skin (typically only -75% of total volume being actually utilized) due to elastic deformation of skin.

[00255] The present device addresses one or more of the limitations mentioned above. [00256] Example 3B: Effective Anti- Obesity Treatment Enabled by Present ML Approach

[00257] Studies have shown that adipose browning could potentially increase energy expenditure, reduce adiposity, and improve metabolic health. Many browning molecules have been discovered (e.g. b-adrenergic hormone, apelin, etc.), however, their low bioavailability in adipose tissue, enzymatic degradation in circulation, rapid renal clearance and off-target effects associated with systemic administration (e.g. intravenous injection) dissuade their practical usage in metabolic diseases. Here, the advantages of micro-lance (ML) for improving therapeutic efficacy of the browning agent in obesity treatment are demonstrated.

[00258] CL316,243, a p3-adrenergic receptor agonist, was used as the browning agent in our mouse experiments because of its proven browning effect on rodent sWAT. It is worthy to mention that a US-FDA approved human b3 -adrenergic receptor agonist (mirabegron) is clinically used for overactive bladder syndrome. Therefore, repurposing it for anti-obesity will largely lower the translational hurdle because its toxicity is well characterized. The bioactivity of mirabegron molecules (1 pg) released from MLs using human white adipocytes was first tested, and observed that browning effects can be induced by mirabegron-laden MLs, similar to the same amount of free mirabegron molecules added in the culture medium, as evidenced by the increased expression of UCP1 protein which is a key biomarker of brown adipocytes responsible for energy consumption (FIG. 5J and 5K). This experiment demonstrates that ML encapsulation does not affect the drug bioactivity, and mirabegron can serve as the browning agent for human.

[00259] Anti-obesity effects of CL316,243 loaded MLs were investigated on diet- induced obese mouse model, which was induced by feeding the mice (6-week old male) with high-fat diet (60% kcal from fat). Obese mice (~30 g) were treated once every 1 or 2 weeks with CL316,243 delivered through IP injection or ML application (but with the same dosage of 1 mg/kg per week). Because of the controlled releasing capability, prolonged dosing-intervals (1-week or 2-week dosing-interval) were chosen to show the long-term effectiveness of ML in obesity treatment. Notably, mice with different treatments had similar food intake as the untreated ones (FIG. 7A and 7B). After 6 weeks, mice continuously fed with high-fat diet gained -23% of body weight (FIG. 4A to 4C) and -70% of the gain was due to the increase of white fat mass including IgWAT and epididymal WAT - EpiWAT (FIG. 4D to 4E). Consistent with the previous studies, IP injection of CL316,243 failed to significantly suppress weight gain. Specifically, -14% and 18% of weight gain were still observed in the groups received IP injection once a week or once two weeks (FIG. 4C). Fluctuation of body weight changes was observed in the latter group presumably because doubled dosage caused weight loss in the initial a few days, and weight gain then restored quickly afterwards. The weight re gaining is likely because drug molecules cannot retain in the fat tissue for a long time. [00260] In comparison, treatment by core-shelled MFs with CF316,243 loaded in both core and shell virtually halted weight gain induced by high-fat diet. Specifically, only -4% and -6 % of weight gain was observed in once-a-week and once-2-week groups, respectively (FIG. 4A to 4C). Consistently, white fat masses (IgWAT and EpiWAT) were reduced by -50% and -40% in these groups (FIG. 4D and 4E). Such outstanding efficacy is attributable to the localized and controlled drug delivery enabled by MFs, which not only provide the initial bolus dose from its fast-releasing shell to quickly reach the therapeutic level, but also maintain therapeutic effect for a long period via its slow-releasing core. The intimate link between obesity and type 2 diabetes is well recognized because excessive fat accumulation causes insulin resistance mainly due to decrease of insulin-stimulated glucose transport and metabolism in adiposes and skeletal muscles. Consistently, it was shown that glucose clearance rate in blood is much slower in obese mice compared to normal mice. As shown in FIG. 4F, insulin resistance in the two ME treated groups was alleviated. Furthermore, the mice in the ME treated groups, but not untreated or IP treated groups, showed increased body surface temperature (FIG. 4G and 4H). Both enhanced insulin sensitivity and elevated body temperature can be ascribed to the transformation of sWAT into thermogenic brown-like adipose tissue.

[00261] Supportively, histological analyses revealed abundant multilocular brown-like adipocytes in IgWATs of ME treated mice, but not in IP injection and control groups (FIG. 5A). Brown-like adipocytes contain multiple lipid droplets and abundant mitochondria, and express thermogenic proteins including UCP1. In addition, white adipocytes in IgWATs of ME treated mice became much smaller than that in untreated mice due to size reduction of the lipid droplets inside. Conceivably, this is because lipid storage in the white adipocytes was largely consumed by the nearby brown-like cells. As demonstrated in FIG. 5B and 5C, the transcription factors that direct browning (PGCla and PRDM16) as well as brown- specific UCP1 were upregulated in the IgWAT of ML treated mice, further confirming CL316,243 induced adipose browning. [00262] Comparing to normal mice, untreated obese mice expectedly had hyperglycaemia, hypercholesterolemia, hypertriglyceridemia, and hyperinsulinemia. These problems can cause various diseases including type 2 diabetes, atherosclerosis and other cardiovascular diseases. Intriguingly, ML treated mice showed a significant reduction in the serum levels of cholesterol, triglycerides, and insulin compared to the untreated group (FIG. 5D to 5G).

[00263] Taken together, ML treatment can suppress the development of obesity and associated metabolic diseases by inducing browning of sWAT through localized and controlled delivery of browning agent. In addition, it was observed that fat browning and reduction effects from biphasic delivery of CL316,243 (in both shell and core) are much better than those from fast releasing of CL316,243 (same amount but in shell only). Specifically, -14% weight gain was still observed in the group received with the latter treatment (FIG. 8A). Consistently, larger fat mass (FIG. 8B) and lesser brown like adipocytes (FIG. 8C) were observed in this group than in the group treated with biphasic delivery of CL316,243. This is because the drug molecules released from ML shell can only dwell within fat tissue for -1 day (FIG. 3H and 31). On the other hand, the same amount of CL316,243 encapsulated in the core for slow release only was also inferior to the biphasic drug administration (-13% weight gain, larger fat mass, and lesser brown-like cells) (FIG. 8A to 8C), conceivably because without the initial burst release, the therapeutic level for browning is not reached. As fat transformation is a reversible process, a sufficient therapeutic level for browning is essential to rectify the whitening effect induced by high-fat diet. Consistently, the metabolic and insulin sensitizing effects from biphasic delivery of CL316,243 are much better than those from fast only or slow only releasing of CL316,243 (same amount in shell only or core only) (FIG. 8D to 8H). These experiments testify the advantages of biphasic release.

[00264] Example 3C: Combination Therapy for Obesity and Metabolic Disorders [00265] Combination therapy that involves co-delivery of several drugs for synergistic or multiple therapeutic effects are highly desired. To simultaneously tackle obesity and the resulting insulin resistance, core and shell of MLs are differentially loaded with adipose browning agent (CL316,243) and insulin- sensitizing drug (rosiglitazone which is an agonist of the peroxisome proliferator-activated receptor-g, PPARy). Rosiglitazone not only is a US-FDA approved anti-diabetic drug but also is recognized as a browning agent by some studies. As shown in FIG. 6A, mice treated once every 2 weeks for 6 weeks, with MLs loaded with 40 pg CL316,243 in inner-core and 20 pg rosiglitazone in outer-shell (equivalent to the combined dosage of 1 mg/kg per week) significantly reduced weight gain induced by high-fat diet (only -8% weight gain vs. -24% in control group). Subcutaneous injection of both CL316,243 and rosiglitazone with the same dosage and frequency failed to exert obvious weight control effect (-20% weight gain). Interestingly, the ML treatment with CL316,243 in the shell and rosiglitazone in the core failed to control the weight gain (-19% gain, FIG. 6A). The much better treatment offered by the burst release of rosiglitazone followed by sustained release of CL316,243 than the otherwise may be explained as follows. Hydrophilic CL316,243 acts on the cell surface via binding with b3 -adrenergic receptor and it is vulnerable to fast tissue clearance. In contrast, rosiglitazone which has a polar head and a hydrophobic tail can readily enter the cell and act intracellularly via binding with the nuclear PPARy receptor. In addition, rosiglitazone promotes angiogenesis that supports the browning process.

[00266] Consistently, as compared with other treatments (subcutaneous injection of rosiglitazone and CL316,243, or MLs with rosiglitazone in the core and CL316,243 in the shell), MLs with CL316,243 in the core and rosiglitazone in the shell performed the best in reducing adipose masses (IgWAT and EpiWAT) (FIG. 6B and 6C), increasing WAT browning as evidenced by increased multilocular brown- like adipocytes in IgWAT (FIG. 6D), and enhancing energy expenditure as evidenced by increased body surface temperature (FIG. 7C), and decreasing serum levels of cholesterol and triglycerides (FIG. 6E and 6F). Although the browning effects from such combination therapy was comparable to the treatment by ML with CL316,243 in both core and shell, it enabled faster glucose clearance rate (FIG. 61 and 6J) and more reduction of serum insulin level (-4 ng/ml vs. -6 ng/ml) (FIG. 5G and 6H). Furthermore, the adipose insulin resistance index (IR = insulin x FFA) was lowest in the mice received the combination therapy (FIG. 7D and 7E). Evidently, the combination therapy not only exerts significant anti-obesity effects via inducing browning but also enhances the insulin sensitivity whereby reducing the risks to develop obesity associated metabolic diseases.

[00267] Example 4: Summary

[00268] Skin, which is the largest organ, contains a vast network of capillaries and lymphatic vessels and has various tissues lying underneath including subcutaneous fat (sWAT), muscles, tendons, ligments, and fibrous joint capsules. As the outer covering of the body, it offers an conveniently accessible interface for drug administration. Transdermal drug delivery can bypass gastrointestinal tract absorption and hepatic first pass effect, and be locally applied at the site just above the targeted or suitable tissues. Herein, a new method, that is, lancing polymeric micro-lances (MLs) through the skin for total implantation inside the underlying tissue to give localized and controlled drug release has been developed.

[00269] The present disclosure demonstrates the use of such paradiagm-shifting approach for combating obesity and associated metabolic diseases via inducing adipose browning and increasing insulin sensitivity. Using an obese mouse model, the outstanding therapeutic effects are evidenced by halted weight gain under high-fat diet, reduced fat masses, appearance of multilocular brown-like adipocytes, upregulated expression of brown- specific proteins including thermogenic UCP1, elevated body surface temperature, increased glucose clearance rate, as well as decreased serum levels of cholesterol, triglycerides and insulin. The present ML approach significantly outperforms the conventional systemic administration methods (IP or subcutaneous injection) because of high drug bioavailability owing to localized and direct delivery, as well as biphasic drug release kinetics and combination therapy enabled by core- shelled MLs. It is also superior to reported microneedle-based transdermal delivery approach in terms of much less frequent treatment (here only once every 2 weeks) and less weight gain using the same weekly dosage of CL316,243. This is attributable to the total implantation of ML inside sWAT and more desirable drug release profiles. [00270] The human clinical studies of both b3 -adrenergic receptor agonists (CL316,243 and mirabegron) using conventional systemic delivery were disappointing. Specifically, chronic and high-dose oral intake of CL316,243 (1500 mg/day, 8 weeks) only enhanced insulin action and fatty acid oxidation, but without any effects on body weight, fat mass and metabolic activity. Mirabegron is more potent than CL316,243. Although, oral intake of mirabegron (50 mg/day, 10 weeks) increased the expression of several brown- specific proteins, no changes on body weight and fat mass were observed. A study has shown that only -30% of orally intaken mirabegron reaches circulation with a plasma half-life of only -40 h. A very small percentage of the circulating mirabegron actually goes to adipose tissue while the majority is absorbed by other tissues. High-dose oral intake of mirabegron (e.g. 200 mg/day) causes adverse cardiovascular problems, such as increasing heart rate and blood pressure. The present ML approach may make mirabegron as well as other browning agents clinically feasible by reducing the effective dosage because of high drug bioavailability and optimal release kinetics. Abdominal and gluteofemoral areas are suitable sites for ML application because they are the largest sWAT depots in human and the least pain sensitive compared to other body parts.

[00271] Lancing ML into these fat depots should be more patient friendly than lancing the stainless-steel lancets into the most pain- sensitive fingertips for blood sampling. For the latter, a study reported that >90% patients considered painless. In addition to frequent blood sampling, diabetes patients also need to do insulin injection using an insulin needle (e.g. 33G x 4 mm) several times a day. Owing to its sustainable releasing property, ML treatment can be much less frequent (e.g. once every 2 weeks). Using a simple applicator, ML can be easily self-administrated. ML wholly implanted in sWAT can then serve as micro-drug-reservoir. In comparison to microneedle-based transdermal drug delivery, total implantation of ML in sWAT ensures a higher loading capacity, 100% bioavailability in the subcutaneous tissues, precise drug dosage, and minimized skin reactions. In addition, implantation depth can be controlled by simply adjusting the applicator’s spring force or injection angle into the skin. Powder or liquid jet injection has been demonstrated for transdermal delivery. But for these methods, -10% of sprayed drugs are loss at the skin surface and the majority of the remaining are dispersed in epidermis and dermis, which are easily adsorbed into circulation before reaching subcutaneous tissues (e.g., sWAT). In addition, as the velocity of jet stream requires at least 100 m/s in order to penetrate the skin, such method not only needs a sophisticated high-pressure jet injector, but also often causes pain and bruising due to the high impact. Having a unique lance-shaped structure with a sharp tip -10 pm, our ML can be easily and swiftly lanced into sWAT using a simple applicator with the lancing speed of ~1 m/s, without causing skin injury and pain, and losing drug on skin surface or into circulation. Also, the applicator can be further devised to simultaneously lance an array of MLs. MLs may also be constructed using different polymers. For example, shell can be made of crosslinked hyaluronic acids (FIG. 9) which gives a prolonged release (~3 days) as compared to CMC. In addition, other drug combinations and loading of multiple drugs in shell or core or both can be used.

[00272] ML approach promises for other diseases as well, for examples, implanting anti-diabetic agents (e.g. sulfonylureas- glimepiride) for diabetes, synthetic hormonal agents (e.g. levonorgestrel) for contraception and hormone replacement therapy, anti psychotic agents (e.g. risperidone) for schizophrenia and other mental disorders, anticancer drugs to tumor tissues (e.g., tamoxifen for breast cancer). It may also be useful for intraocular drug delivery by bypasing anatomical barriers (e.g. cornea, sclera). For example, MLs can be implanted into sclera for sustained delivery of corticosteriods for uveitis or b-adrenergic receptor blockers for glaucoma. In summary, such safe, painless and convenient transdermal delivery method may provide unprecedented long-term home-based solution for obesity and many other diseases. [00273] Example 5: Commercial and Potential Applications [00274] Existing IDD devices, including drug infusion pumps, intraocular drug delivery devices, and contraceptive drug delivery devices can be replaced by the present patient-friendly self-administrable and implantable micro-lance therapeutics delivery devices, because those conventional devices may inconveniently require surgery and clinic visit.

[00275] Advantageously, the present device can be applied in home -based care of people with obesity or overweight. It has enormous potential to replace currently- available obesity treatments (liposuction, lipolysis), or FDA-approved “anti-obesity drugs”, because it is safe, cost-effective, and patient-friendly.

[00276] While the present disclosure 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 present disclosure as defined by the appended claims. The scope of the present disclosure 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 therapeutics delivery device implantable in a biological tissue, comprising: a core which is biodegradable, wherein the core comprises a first polymer composition, wherein the core has a height of at least 1 mm, and wherein the core has one end shaped as a sharp tip sufficient to penetrate the biological tissue, wherein the sharp tip comprises a single bevel-angled shape, a double bevel-angled shape, a chiseled edge, a bullet-shape, or a frustum-shape.

2. The therapeutics delivery device of claim 1, wherein the first polymer composition comprises poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly(lactic acid-co-glycolic acid) (PLGA), polyanhydride, polyorthoester, polyetherester, polycaprolactone (PCL), polyesteramide, poly(butyric acid), poly(valeric acid), polyvinylpyrrolidone (PVP), polyvinyl alcohol (PVA), polyethylene glycol (PEG), PEG-PLA, PEG-PLA-PEG, PLA-PEG-PLA, PEG-PLGA, PEG-PLGA-PEG, PLGA- PEG-PLGA,PEG-PCL, PEG-PCL-PEG, PCL-PEG-PCL, ethylene glycol-propylene glycol-ethylene glycol (PEG-PPG-PEG), hyaluronic acid, a derivative thereof, or a mixture thereof.

3. The therapeutics delivery device of claim 1, wherein the first polymer composition is formed from two monomers, wherein the two monomers comprise lactic acid and glycolic acid, wherein the two monomers are present in a molar ratio of 1 : 100 to 100:1.

4. The therapeutics delivery device of any one of claims 1 to 3, further comprising: a shell which is biodegradable, wherein the shell comprises a second polymer composition, wherein the first polymer composition and the second polymer composition have different degradation rate.

5. The therapeutics delivery device of claim 4, wherein the second polymer composition comprises hyaluronic acid and its derivatives, methacrylate hyaluronic acid, sodium alginate, collagen and its derivatives, polyurethane, poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly(lactic acid-co-glycolic acid) (PLGA), polyanhydride, polyorthoester, polyetherester, polycaprolactone (PCL), polyesteramide, poly(butyric acid), poly(valeric acid), polyvinylpyrrolidone (PVP), polyvinyl alcohol (PVA), polyethylene glycol (PEG), PEG-PLA, PEG-PLA-PEG, PLA-PEG-PLA, PEG-PLGA, PEG-PLGA-PEG, PLGA-PEG-PLGA, PEG-PCL, PEG- PCL-PEG, PCL-PEG-PCL, ethylene glycol-propylene glycol-ethylene glycol (PEG- PPG-PEG), dextran, hetastarch, tetrastarch, pentastarch, hydroxyethyl starch, cellulose, hydroxypropyl cellulose (HPC), sodium carboxymethyl cellulose (Na CMC), thermosensitive HPMC (hydroxypropyl methyl cellulose), polyphosphazene hydroxyethyl cellulose (HEC), a polysaccharide, polyalcohol, gelatin, alginate, chitosan, sucrose, or a mixture thereof.

6. The therapeutics delivery device of claim 4 or 5, wherein the core and/or shell further comprises a porogen, wherein the porogen comprises sodium chloride (NaCl), polyethylene glycol, an ionic liquid, or a mixture thereof.

7. The therapeutics delivery device of claim 6, wherein the porogen is present in amount of less than 75 wt% of (i) the first polymer composition forming the core and/or (ii) the second polymer composition forming the shell.

8. The therapeutics delivery device of any one of claims 1 to 7, wherein the core has a cross-sectional diameter ranging from 200 pm to 600 pm.

9. The therapeutics delivery device of any one of claims 1 to 8, wherein: the core is a solid core, wherein the solid core houses one or more therapeutics therein; or the core further comprises a polymer matrix defining a hollow reservoir, wherein either or both of the polymer matrix and the hollow reservoir each houses one or more therapeutics.

10. The therapeutics delivery device of any one of claims 1 to 9, wherein the shell comprises one or more therapeutics.

11. The therapeutics delivery device of claim 9 or 10, wherein the one or more therapeutics comprise an adipose browning agent, a lipolytic agent, an anti-obesity agent, a blood sugar lowing agent, an insulin sensitizing agent, an anti-cancer agent, an anti-psychotic agent, an anti-fungal agent, an anti-microbial agent, an anti-psychotic agent, an anesthetic agent, a pro-angiogenic agent, an anti-angiogenic agent, a hair- growth promoting agent, a hormonal agent, a protein or peptide, an amino acid, an antibody or antibody fragment, an oligonucleotide, a cell or cell fragment, an exosome or microvesicle, or a combination thereof.

12. An applicator operable to implant the therapeutics delivery device of any one of claims 1 to 11, comprising: a housing comprising a lancing module and one or more holder tubes, wherein each of the one or more holder tubes is configured for loading the therapeutics delivery device therein, and wherein the lancing module is operable to drive the therapeutics delivery device out of the one or more holder tubes to implant the therapeutics delivery device of any one of claims 1 to 11 into the biological tissue.

13. The applicator of claim 12, wherein the housing, the lancing module, and/or the one or more holder tubes is formed of stainless steel.

14. A method of producing the therapeutics delivery device of any one of claims 1 to 11, comprising: providing a mold to form the core, wherein the mold comprises a cavity having a depth of at least 1 mm and a closed end comprising a single bevel-angled shape, a double bevel-angled shape, a chiseled edge, a bullet-shape, or a frustum-shape; depositing a first polymer solution into the cavity, wherein the first polymer solution comprises the first polymer composition; and drying the first polymer solution to form the core.

15. The method of claim 14, further comprising coating a second polymer solution on the core to form the shell, wherein the second polymer solution comprises the second polymer composition.

16. The method of claim 14 or 15, wherein the mold is formed of a polymeric material, a ceramic material, or a metallic material.

17. The method of any one of claims 14 to 16, wherein depositing the first polymer solution comprises: dissolving the first polymer composition in an organic solvent to form the first polymer solution, and applying a pressure on the first polymer solution in the mold, wherein the pressure comprises a hydraulic pressure ranging from 10 to 500 pounds per square inch (psi).

18. The method of any one of claims 14 to 17, wherein depositing the first polymer solution comprises mixing the porogen and/or the one or more therapeutics in the first polymer solution.

19. The method of any one of claims 14 to 18, further comprising heating the mold after depositing the first polymer solution and prior to drying the first polymer solution, wherein heating the mold comprises heating the mold to a temperature ranging from 50 to 200°C.

20. The method of any one of claims 15 to 19, further comprising: mixing the one or more therapeutics in the second polymer solution; removing the core from the mold prior to coating the second polymer solution, and immersing the core into the second polymer solution to form the shell.

21. A method of treating a medical condition or delivering a therapeutics, wherein the method comprises: operating the applicator of claim 12 or 13 to implant the therapeutics delivery device of any one of claims 1 to 11 in the biological tissue.

22. The method of claim 21, wherein the medical condition comprises obesity, a metabolic disease, a cancer, an eye disease, a skin disease, alopecia, or schizophrenia.