Transdermal delivery device, methods of using and making the same

Abstract

Provided is a device for transdermal delivery of drugs. The device includes a separable substrate and is loaded with dual drugs based on an interpenetrating polymer network hydrogel. Also provided are methods of making and using the transdermal delivery device.

Description

[Technical Field]

[0001] This disclosure relates to a transdermal delivery device having a base and a plurality of protrusions made of interpenetrating polymer mesh hydrogel for loading and delivering drugs. [Previous Technology]

[0002] The field of drug delivery technology focuses on delivering drugs to targets in different organs, with the aim of improving drug efficacy and minimizing side effects by controlling drug release.

[0003] To date, there are many methods for delivering drugs into the human body to treat cancer, which are generally divided into oral, parenteral route injection (intravenous, intramuscular or subcutaneous) and transdermal drug delivery.

[0004] However, both oral and parenteral administration routes have several limitations that affect their effectiveness; for example, oral administration leads to enzymatic digestion and may cause liver side effects because it passes through the liver before entering the bloodstream to exert its therapeutic effect. Furthermore, off-target delivery is associated with adverse side effects, and injection can be harmful and painful for patients, especially increasing the risk of infection. Therefore, there is an urgent need for an effective cancer treatment delivery route.

[0005] To overcome the limitations of oral or parenteral administration, microneedles (MNs) that establish micropathways on the skin have been developed, allowing molecular drugs to enter and reach their target sites via percutaneous delivery devices. Percutaneous administration, by delivering drug formulations intended to diffuse into the systemic circulation, to the skin surface, is a patient-friendly and safe method of drug delivery. For example, compared to oral or parenteral administration, percutaneous administration has many advantages, including efficient delivery and the potential for subsequent therapeutic outcomes such as local immune activation and cancer therapy. Microneedle-mediated percutaneous administration is also superior to traditional intravenous, intramuscular, or subcutaneous drug delivery, minimizing the risk of drug metabolism before its action and reducing the frequency of administration.

[0006] However, unlike microneedles derived from silicon or metals, biodegradable polymeric microneedles manufactured using these conventional methods suffer from problems such as bending and deformation during skin application due to a lack of sufficient mechanical strength. In this regard, the polymerization method and monomer type are issues that need to be considered in adjusting the mechanical strength and degradation rate of the microneedles to penetrate the skin and enhance drug release.

[0007] Therefore, there is still a need for improved microneedle devices that have sufficient mechanical properties and are easy to use on the skin surface. [Summary of the Invention]

[0008] This disclosure provides a solution to many of the shortcomings of early microneedles by providing a transdermal delivery device comprising a plurality of projections, each projection comprising a first polymer formed of a first monomer having a first bond and a second polymer formed of a second monomer having a second bond; a substrate comprising a third polymer formed of the first monomer having a third bond; and a bioactive agent contained in one of the projections, wherein the projections are coupled to the substrate and configured to be at least partially insertable into the skin of an individual in need, and after the transdermal delivery device has been applied to the skin for a predetermined time, a compound used to target and destroy the third bond of the third polymer in the substrate is removed from the substrate.

[0009] Due to improved mechanical and chemical properties, the plurality of protrusions of the microneedles (MNs) disclosed herein can be used in applications where existing microneedles have failed. The microneedles disclosed herein improve performance in existing applications. For example, the improved microneedles disclosed herein can be used in applications where existing microneedles lack sufficient mechanical strength, such as penetration into the stratum corneum and other biological tissues such as blood vessels, heart valves, muscles, and skin, piercing the outer layer of skin, and permeability of compounds into the bloodstream. The improved microneedles disclosed herein can also be used for percutaneous drug delivery and tumor growth inhibition. This disclosure also provides methods for manufacturing microneedles and treating cancer by reducing tumor mass and increasing immune activation around the tumor site.

[0010] This disclosure provides an apparatus comprising an interpenetrating polymer network (IPN) hydrogel. In at least one embodiment, the hydrogel comprises sodium alginate and sulfobetaine methyl acrylate (SBMA), which are sequentially photocrosslinked with N,N'-methylenebisacrylamide (MBAAm) and then ionically crosslinked with calcium ions. The apparatus further comprises a separable substrate comprising an IPN hydrogel crosslinked with disulfide bonds and N,N-bisacryloylcystine (BISS), wherein the disulfide bonds are cleaved by dithiothreitol (DTT) and / or ethylenediaminetetraacetic acid (EDTA) and are separable from the microneedle array.

[0011] This disclosure provides a drug delivery device loaded with a dual drug, such as lipopolysaccharide (LPS) and doxorubicin (DOX), for synergistic immunochemotherapy outcomes. LPS is a well-studied immunostimulatory macromolecule composed of lipids and polysaccharides, capable of reprogramming tumor-associated macrophages (TAMs) into an M1-like phenotype with anti-tumor function by secreting potent antitumor pro-inflammatory cytokines such as tumor necrosis factor-α (TNF-α). Pro-inflammatory cytokines have been reported to have synergistic therapeutic effects with various anticancer drugs such as doxorubicin (DOX), dactinomycin, and etoposide. LPS further induces the activation of antigen-presenting cells and T cells, including CD4+, CD8+, and CD25+ cells, as well as other cancer immune surveillance cells. A sufficient quantity of drug is released from the microneedles in the initial 24 hours, followed by slow and continuous release over the next 7 days. For example, 66.1±7.4% and 59.4±5.5% of DOX and LPS, respectively, were released in the first 24 hours. In at least one embodiment, in vivo studies demonstrated that the presence of LPS enhanced the performance and activation of immunomodulators and synergistically enhanced the value of immunochemotherapy. Thus, compared with individual drugs, microneedles loaded with dual drugs, such as microneedles loaded with LPS and DOX, induced significant tumor suppression in C57BL / 6 mice carrying gliomas (p<0.05). Therefore, combined drug delivery via microneedle-mediated percutaneous administration to subcutaneous tumors is an effective drug administration method that can improve therapeutic efficacy with negligible systemic adverse reactions.

[0012] In at least one embodiment, this disclosure provides a drug delivery device comprising a separable and mechanically strong microneedle patch capable of withstanding up to 0.64 N / needle for efficient transdermal drug delivery. In at least one embodiment, the microneedles are further comprising a disulfide crosslinking agent to form a separable substrate, which can be decomposed and removed by a reducing agent or a metal chelating agent, such as DTT and / or EDTA solution, while retaining the microneedle array within the skin for continuous transdermal drug release.

[0013] In at least one embodiment, this disclosure provides a drug delivery device in the form of a microneedle array loaded with, for example, two drugs for transdermal administration, followed by removal of a separable substrate of the microneedles with a solution of dithiothreitol (DTT) and / or ethylenediaminetetraacetic acid (EDTA). In at least one embodiment, the microneedle array comprises an interpenetrating polymeric network hydrogel having a separable substrate, wherein the interpenetrating polymeric network hydrogel comprises sodium alginate and SBMA monomers, wherein the sodium alginate and SBMA are photocrosslinked with MBAAm and ionically crosslinked with calcium ions. In some embodiments, the separable substrate comprises disulfide links, for example, N,N-bisacrylcysteine.

[0014] This disclosure provides a transdermal delivery device comprising a plurality of protrusions, each of the plurality of protrusions comprising a first polymer formed of a first monomer having a first bond and a second polymer formed of a second monomer having a second bond; a substrate comprising a third polymer formed of the first monomer having a third bond; and a bioactive agent contained in one of the plurality of protrusions, wherein the plurality of protrusions are coupled to the substrate and configured to be at least partially insertable into the skin of an individual in need, and after the transdermal delivery device has been applied to the skin for a predetermined time, to target and destroy the compound forming the third bond of the third polymer in the substrate to remove the substrate.

[0015] In at least one embodiment, the first monomer is an amphoteric ion. In some embodiments, the amphoteric ion is phosphorylcholine, sulfobetaine, alkylsulfonate pyridinium salt, carboxybetaine, phosphobetaine, phosphonobetaine, phosphinobetaine, ammoniosulfate, ammoniosulfonamide, pyridiniocarboxylate, or sulfoniocarboxylate. In at least one embodiment, the phosphorylcholine is phosphorylcholine acrylate, phosphorylcholine acrylamide, phosphorylcholine methyl acrylate, alkoxydicyanoethnolate, or 2-methylacryloyloxyethylphosphocholine. In at least one embodiment, the sulfobetaine is sulfobetaine acrylate, sulfobetaine acrylamide, sulfobetaine methyl acrylate, sulfobetaine vinylimidazolium, or sulfobetaine vinylpyridine. In at least one embodiment, the carboxybetaine is carboxybetaine acrylate, carboxybetaine methyl acrylate, carboxybetaine acrylamide, carboxybetaine vinylimidazolium, carboxybetaine methacrylamide, carboxybetaine isobutylene, or carboxybetaine diallylamine. In at least one embodiment, the alkyl sulfonate pyridine salt is 3-(2-vinylpyridinium-1-yl)propane-1-sulfonate, N-(2-methacryloxy)ethyl-N,N-dimethylammonium propanesulfonate, or N-(3-methacrylylimino)propyl-N,N-dimethylaminopropanesulfonate.

[0016] In at least one embodiment, the second polymer is one or more selected from the group consisting of biocompatible synthetic polymers, semi-synthetic polymers, and natural polymers. In at least one embodiment, the second polymer is selected from gums, polysaccharides, polysaccharide derivatives, alginates containing sodium alginate or calcium alginate, chitosan, chitosan derivatives, collagen, gelatin, polydextrose, poly(vinylpyrrolidone), hydroxyethyl starch, polyethylene glycol, functionalized polydextrose, trehalose-containing sugar polymers, hyaluronic acid, methacrylated hyaluronic acid, poly(methyl vinyl ether), poly(methyl vinyl ether-alt-maleic anhydride), poly(lactic acid), polyglycolic acid, poly(lactic acid-glycolic acid copolymer), polycarbonate, poly(vinyl alcohol), poly(hydroxyethyl methacrylate), poly(vinylpyrrolidone). (e.g., ketones), (2-carboxymethyl)-3-propenylaminopropyl dimethylammonium bromide, (2-carboxymethyl)-3-propenylaminopropyl dimethylammonium bromide-co-hydroxyethyl methacrylate, (2-carboxymethyl)-3-propenylaminopropyl dimethylammonium bromide-co-propenylamine, methacrylated (2-carboxymethyl)-3-propenylaminopropyl dimethylammonium bromide-co-propenylamine, poly(ε-caprolactone)poly(ε-caprolactone-co-glycolic acid), poly(2-methacryloyloxyethyl phosphocholine), poly(carboxybetaine)vinylimidazolium, poly(sulfobetaine)vinylimidazolium, and poly(sulfobetaine)vinylpyridine.

[0017] In at least one embodiment, the first monomer is cross-linked by chemical bonds, and the second monomer is cross-linked by physical bonds. In at least one embodiment, the chemical bonds and physical bonds form an interpenetrating polymer network.In at least one embodiment, the chemical bond is formed by means of N,N'-methylenebisacrylamide (MBA), bisacrylamide derivative of cystine (BISS), dimethylsubermidate, glutaraldehyde, N,N-ethylidene-bis(iodoacetamide), ethylene glycol dimethacrylate (EGDM), poly(ε-caprolactone) diacrylate, polylactic acid diacrylate, polylactic acid dimethacrylate, poly(lactic-co-glycolic acid) diacrylate, poly(lactic-co-glycolic acid) dimethacrylate, poly(ε-caprolactone-b-ethylene glycol-b-ε-caprolactone) diacrylate, and ethylene glycol-b-(lactic-co-glycolic acid) dimethylpropionate. Acrylates, polymerizable compounds containing disulfide bonds, peptide bonds, or ester bonds, poly(ε-caprolactone) dimethacrylate (MAC-PCL-MAC), poly(ε-caprolactone-b-ethylene glycol-b-ε-caprolactone) dimethacrylate (MAC-PCL-PEG-PCL-MAC), poly(lactic acid-b-ethylene glycol-b-lactic acid) diacrylate (AC-PLA-PEG-PLA-AC), poly(lactic acid-b-ethylene glycol-b-lactic acid) dimethacrylate (MAC-PLA-PEG-PLA-MAC), poly[(lactic acid-co-glycolic acid)-b-ethylene glycol-b-(lactic acid-co-glycolic acid)] diacrylate (AC-PLGA-PEG-P) Poly(lactic acid-co-glycolic acid)-b-ethylene glycol-b-(lactic acid-co-glycolic acid)] dimethacrylate (MAC-PLGA-PEG-PLGA-MAC), poly(ε-caprolactone-co-lactic acid)-diacrylate (AC-PCLA-AC), poly(ε-caprolactone-co-lactic acid) dimethacrylate (MAC-PCLA-MAC), poly(ε-caprolactone-co-glycolic acid) diacrylate (AC-PCGA-AC), poly(ε-caprolactone-co-glycolic acid) dimethacrylate (MAC-PCGA-MAC), poly(ε-caprolactone-co-glycolic acid)-b-ethylene glycol-b-(ε-caprolactone-co-lactic acid) It is formed by at least one crosslinking agent from the group consisting of diacrylate (AC-PCLA-PEG-PCLA-AC), poly(ε-caprolactone-co-lactic acid)-b-ethylene glycol-b-(ε-caprolactone-co-lactic acid) dimethacrylate (MAC-PCLA-PEG-PCLA-MAC), poly(ε-caprolactone-co-glycolic acid)-b-ethylene glycol-b-(ε-caprolactone-co-glycolic acid) diacrylate (AC-PCGA-PEG-PCGA-AC) and poly(ε-caprolactone-co-glycolic acid)-b-ethylene glycol-b-(ε-caprolactone-co-glycolic acid) dimethacrylate (MAC-PCGA-PEG-PCGA-MAC).

[0018] In at least one embodiment, the ratio of the first monomer to the second monomer is about 1:1, about 1:2, about 1:3, about 1:4, about 1:5, about 2:1, about 3:1, about 4:1 or about 5:1.

[0019] In at least one embodiment, the third bond is a disulfide bond. In at least one embodiment, the compound used to target and disrupt the third bond forming the third polymer in the substrate is dithiothreitol (DTT), ethylenediaminetetraacetic acid (EDTA), glutathione (GSH), β-mercaptoethanol, or L-cysteine.

[0020] In at least one embodiment, each of the plurality of protrusions has a gradually tapering shape. In at least one embodiment, each of the plurality of protrusions has a conical or cone-shaped form. In at least one embodiment, each of the plurality of protrusions has a height between about 25 μm and about 2,500 μm, a width between about 50 μm and about 250 μm, and a tip diameter between about 1 μm and about 25 μm.

[0021] This disclosure further provides a method for manufacturing a transdermal delivery device, comprising: preparing a first solution comprising a first monomer, a second monomer, a first crosslinking agent and at least one bioactive agent; preparing a second solution comprising the first monomer and the second crosslinking agent; applying the first solution comprising the at least one bioactive agent to a mold and centrifuging the first solution; after centrifugation, removing the upper layer of the first solution and applying the second solution to the top of the mold; covering the mold with a cover mold and centrifuging; applying a first condition suitable for causing the first solution to solidify to form a plurality of protrusions; applying a second condition suitable for causing the second solution to solidify to form a substrate; and demolding the plurality of protrusions and the substrate from the mold to obtain the transdermal delivery device.

[0022] This disclosure further provides a method for inducing bioactivity in an individual in need, comprising: providing the transdermal delivery device as described above; applying the transdermal delivery device to the skin of the individual such that the plurality of protrusions pierce the skin of the individual; and removing the substrate of the transdermal delivery device from the individual by targeting and destroying the third bond forming the third polymer in the substrate, wherein the plurality of protrusions containing the bioactive agent remain in the skin. [Simplified Explanation of the Diagram]

[0023] Figure 1A is a schematic illustration of the IPN hydrogel, which shows that the interpenetrating polymer network (IPN) hydrogel uses sequential chemical cross-linking of SBMA network, followed by ionic cross-linking of alginate network and calcium ions (Ca2+) to create a relatively tough microneedle array.

[0024] Figure 1B is a schematic diagram showing the fabrication of microneedles.

[0025] Figure 2A shows the NMR and Raman spectra of the disulfide crosslinking agent, BISS and L-cystine monomer, which have the characteristic peaks of the monomer, δ=2.89 to 3.01ppm (a,a') and δ=3.48ppm (b).

[0026] Figure 2B shows the NMR and Raman spectra of the disulfide crosslinking agent, BISS and acrylamide chloride monomer, which have monomer characteristic peaks, δ=5.62ppm(b), δ=5.80ppm(c) and δ=6.14ppm(a), used to synthesize BISS.

[0027] Figure 2C shows the NMR and Raman spectra of the disulfide crosslinking agent BISS, which have corresponding proton spectra at δ=3.05ppm(a) and 3.4ppm(a'), δ=4.8ppm(b), δ=5.85ppm(c), 6.3ppm(c') and δ=6.45ppm(d), which confirms the successful synthesis of the disulfide crosslinking agent (BISS).

[0028] Figure 2D shows the Raman spectrum of the covalent bonds formed in the disulfide crosslinking agent (BISS), and the characteristic Raman signals at 500 cm-1, 670 cm-1 and 1247 cm-1 in BISS represent SS, CS and CC bonds, respectively, indicating that the crosslinking agent was successfully synthesized.

[0029] Figure 3A shows the Raman spectrum of the IPN hydrogel, where the Raman signal appears at approximately 1133 cm⁻¹, representing the asymmetric vibration of SO₃⁻ derived from SBMA. Characteristic peaks at 807 cm⁻¹, 888 cm⁻¹, and 954 cm⁻¹ represent the stretching vibrations of CC, CO, and CCO from sodium alginate, respectively. Raman signals at 1240 cm⁻¹ and 1413 cm⁻¹ represent the stretching vibrations of COO⁻ from alginate, confirming the formation of the IPN hydrogel. The characteristic Raman spectra of the MBAAm crosslinking agent and the photoinitiator (α-ketoglutarate) are not visible in the IPN hydrogel, suggesting that the toxic monomers have completely reacted to form the IPN hydrogel, and that the potential harmful effects of the hydrogel are negligible.

[0030] Figure 3B is the FT-IR spectrum of the chemical shift of the COO- signal from 1595 cm-1 in sodium alginate to 1642 cm-1 in IPN hydrogel, indicating the formation of ionic crosslinks (-COOCa) with calcium ions.

[0031] Figures 4A to 4C are a series of FESEM images of the internal structure of IPN hydrogels with SBMA:alginate monomer ratios of 1:1, 2:1 and 3:1, respectively.

[0032] Figure 5A shows a microscopic observation of the microneedle morphology.

[0033] Figures 5B and 5C are FESEM images of microneedle morphology.

[0034] Figures 6A to 6C are line graphs showing the mechanical properties of IPN hydrogels. Figure 6A is a line graph of tensile test of IPN hydrogels under different crosslinking conditions; Figure 6B is a line graph of tensile test of IPN hydrogels with different monomer ratios; and Figure 6C is a line graph of compressive stress of microneedle patches made of 1:1-UV60 and Ca2+30.

[0035] Figure 7 is a series of microscopic images showing the morphological changes of the microneedles over time, which were immersed in PBS solution (pH 7.4) at 37°C.

[0036] Figure 8A is a schematic diagram of a separable IPN hydrogel having a crosslinking agent immersed in a mixture of DTT and EDTA solutions (DTT 20mM, EDTA 100mM).

[0037] Figure 8B is a line graph showing the effect of crosslinking agent concentrations at ratios of 1:10 and 1:100 on the degradation rate of separable IPN hydrogels.

[0038] Figure 8C is a series of photographs showing hydrogels with different concentrations of crosslinking agents.

[0039] Figure 8D is a line graph showing the disintegration of separable IPN hydrogels in different disintegration solutions (n=3).

[0040] Figure 9A is a photograph of mouse skin before the microneedles were inserted.

[0041] Figure 9B is an optical image of the circular hole left on the skin after the microneedle penetrates.

[0042] Figure 9C is an H&E staining image of skin tissue after microneedle penetration.

[0043] Figure 9D is a schematic diagram of the force-holding device during the mouse skin penetration test.

[0044] Figure 9E shows the morphology of the microneedle after it penetrates the skin.

[0045] Figure 10A is a zigzag graph showing the in vitro drug release of IPN hydrogel microneedles (n=3), which represents the cumulative release of DOX.

[0046] Figure 10B is a line graph showing the in vitro drug release of IPN hydrogel microneedles (n=3), which is the cumulative release of LPS.

[0047] Figure 11A is a bar graph showing the chemical toxicity of DOX released from microneedles on CT-2A-Luc cells and NIH-3T3 cells.

[0048] Figure 11B is a line graph showing the IC50 trend of DOX released from microneedles in CT-2A-Luc cells and NIH-3T3 cells (n=8).

[0049] Figure 12A is a line graph showing the tumor inhibition, particularly tumor volume, of drug-loaded microneedles in vivo during treatment.

[0050] Figure 12B is a photograph showing the parallel trend, where the tumor size treated with microneedles loaded with LPS / DOX is significantly smaller than that of other formulations.

[0051] Figure 12C is a bar graph showing tumor weight after 18 days of treatment with dual-drug therapy (microneedles loaded with LPS / DOX) or single-drug therapy (microneedles loaded with LPS or microneedles loaded with DOX).

[0052] Figure 12D is a line graph showing the weight of mice treated with microneedles versus the weight of the control group.

[0053] Figures 13A and 13B show immunohistochemical (IHC) staining of spleen and tumor tissue.

[0054] Figures 14A and 14B show the ultraviolet-visible (UV-vis) absorption spectra and standard calibration curves of DOX and LPS at different series concentrations, respectively.

[0055] Figure 15 shows the FTIR spectra of the disulfide crosslinking agent (BISS) and its monomer.

[0056] Figures 16A and 16B show the results of the biocompatibility test (MTT assay) of IPN hydrogel on mouse glioma cells (CT-2A-Luc cells) and mouse embryonic fibroblasts (NIH-3T3 cells), respectively. (n=8)

Implementation Method

[0057] Examples of the present disclosure are illustrated with reference to a portion of the detailed description of the constituent examples. It should be understood that other examples may be used and modifications may be made without departing from the scope of the present disclosure.

[0058] Unless otherwise stated, the singular forms “a”, “a” and “the” used in this specification and the appended claims shall be deemed to include both the singular and plural forms, unless otherwise stated or clearly contradicted by the context.

[0059] The term “approximately” as used herein means approximately or close to a value in the context of the value. In one instance, the term “approximately” may include conventional rounding based on the significant number of the value. Furthermore, the phrase “approximately x to y” includes “approximately x to approximately y”.

[0060] Unless otherwise stated herein, the term "or" as used in this specification and the appended claims generally includes the meaning of "and / or". As used herein and unless otherwise stated, the conjunction "and / and / as well as" is intended to be inclusive, and the conjunction "or" is not intended to be exclusive. For example, the phrase "or alternatively" is intended to be exclusive.

[0061] Unless otherwise stated herein, the terms “comprising,” “having,” “including,” and “containing” shall be regarded as open-ended terms (i.e., meaning “including but not limited to”).

[0062] Unless otherwise stated herein, the description of a numerical range is merely a shorthand for all individual values ​​falling within that range, each of which is included in the specification as if described separately herein.

[0063] Unless otherwise stated herein or contradicted by the context, all methods described herein may be performed in any suitable order. Unless otherwise required, the use of any and all instance or illustrative terms (such as "for example" and "for instance") is for the purpose of setting forth this disclosure and not to limit its scope.

[0064] As used herein, the term "prevention" or "avoidance" is defined as the possibility of eliminating or reducing the occurrence of one or more symptoms of cancer or tumor. For example, the compositions described herein may be used to treat tumors or reduce tumor cells or to treat cancer or reduce cancer cells.

[0065] As used herein, the terms "treatment" or "therapy" refer to the administration of an effective dose of an anticancer drug to an individual in need to cure, alleviate, treat, improve, or prevent cancer, its symptoms, or the risk of developing cancer. The individual can be identified by a healthcare professional based on results from any appropriate diagnostic method.

[0066] As used herein, the term "sufficient dose of medicine" means sufficient to prevent the development, recurrence or onset of cancer and one or more of its symptoms, enhance or improve the preventive effect of another therapy, reduce the severity and stage of cancer, improve one or more of the symptoms of cancer, prevent the progression of cancer and / or enhance or improve the therapeutic effect of another therapy.

[0067] As used herein, the term "individual" is any organism that requires treatment and / or prevention of cancer. In at least one instance, the individual is a mammal, including but not limited to humans and domesticated animals (e.g., rats and mice).

[0068] The application device based on the interpenetrating polymer network hydrogel has a separable substrate for the transdermal delivery device described herein, overcoming the limitations of earlier network hydrogels. The improved interpenetrating polymer network hydrogel described herein differs from previous hydrogels in at least the following aspects: (1) mechanical strength, for example, the robust microneedle patch provided herein can withstand a compressive stress value of 76.8 N (0.64 N / needle), which confirms that the microneedles can effectively penetrate the skin (Fig. 6C); (2) drug loading and release or timely release of drugs, for example, rapid release of doxorubicin (DOX) and lipopolysaccharide (LPS) was observed in the first 4 hours, followed by stable release through the degradation of the microneedles. The degradation of microneedles based on interpenetrating polymer network (IPN) hydrogels in phosphate-buffered saline (PBS) media is attributed to sodium-calcium exchange in the hydrogel, which enhances the swelling and disintegration of the hydrogel to facilitate drug diffusion (Fig. 10A and Fig. 10B); (3) stable transdermal drug release, e.g., water absorption and swelling of the microneedles, identified over time by gradual structural loosening (Fig. 7); (4) the separability of the disulfide crosslinked substrate of the microneedles, e.g., by adjusting the disintegrant concentration to improve the disintegration rate of the disulfide-linked IPN hydrogel and the separability of the microneedles (Fig. 8C and Fig. 8D); (5) the biocompatibility of the IPN hydrogels, e.g., even at high concentrations (0.1 g / mL), cell survival rates treated with N,N-bisacrylamide (BISS) or N,N'-methylenebisacrylamide (MBAAm) crosslinked hydrogels in CT-2A-Lu (5) >85% on c cells and >92.6% on NIH-3T3 cells, indicating that the hydrogel is biocompatible and does not affect cells during transdermal application; (6) Microneedles loaded with doxorubicin treat cancer cells with higher IC50 values ​​or show better cell survival at the same concentration of DOX (Fig. 11A and Fig. 11B); (7) Synergistic effect on tumor growth inhibition, for example, tumor growth in mice treated with dual drugs (microneedles loaded with LPS and DOX) was more inhibited than in mice treated with microneedles loaded with DOX only (Fig. 12A to Fig. 12D); (8) Transdermal delivery of LPS to the epidermis rich in immune cells and triggering a wide range of immune responses such as apoptosis, for example, cancer cells undergo apoptosis in response to the drug (microneedles loaded with LPS and DOX) and inhibit tumor progression (Fig. 13A and Fig. 13B).

[0069] The effectiveness of this disclosure will be further illustrated by the following examples, which are not intended to limit the scope of this disclosure.

[0069] Example

[0070] Example 1: Fabrication of IPN hydrogel

[0071] Materials

[0072] Sodium alginate, sodium hydroxide (NaOH), sulfobetaine methyl acrylate (SBMA, 95%), α-ketoglutarate, acrylonitrile chloride (97%), calcium formate (Ca(HCOO)2), L-cysteine ​​(99.7%), ethylenediaminetetraacetic acid (EDTA, 98.5%), D,L-dithiothreitol (DTT, 99%), methylthiazolyl diphenyl-tetrazole bromide (MTT, 97.5%), doxorubicin (DOX), and lipopolysaccharide (LPS) were purchased from Sigma Aldrich, and N,N'-methylenebisacrylamide (MBAAm, 99.5%) was purchased from JTBaker. Phosphate-buffered saline (PBS), Duchenne modified Eagle medium (DMEM), and their supplements were obtained from Hyclone.

[0073] Synthesis of disulfide bond crosslinking agents

[0074] N,N-bisacrylcysteine ​​(BISS), a disulfide crosslinking agent, was used as a separable substrate for microneedles in the preparation of IPN hydrogels. 2.7 g (11.2 mM) L-cysteine ​​and 2 g (50 mM) sodium hydroxide (NaOH) were dissolved in 70 mL of methanol in an ice bath at 0 °C. When a clear, colorless solution was formed, 2.2 mL (27.2 mM) acrylamide chloride was added dropwise, and the reaction mixture was stirred at room temperature for 4 hours. Finally, the reaction mixture was purified by adding it dropwise to cold, vigorously stirred diethyl ether, and the precipitate was separated by centrifugation. Residual diethyl ether was removed by drying in a vacuum oven for 12 hours. The successful synthesis of BISS was verified by proton nuclear magnetic resonance spectroscopy (¹H NMR, Bruker AVANCE 600 MHz), Raman spectroscopy (JASCO NRS-5100 Laser), and Fourier transform infrared spectroscopy (FT-IR, Thermo Nicolet 6700 system).

[0075] Preparation of MBAAm cross-linked pregel hydrogel solution

[0076] Before forming a sequentially interpenetrating polymer network (IPN) hydrogel through chemical and ionic interactions, a pregel solution crosslinked with N,N'-methylenebisacrylamide (MBAAm) was prepared as follows. First, 0.6984 g (500 mM) of sulfobetaine methyl acrylate (SBMA) monomer was dissolved in 5 mL of deionized (DI) water, followed by the addition of 3.8 mg (5 mM) of crosslinking agent MBAAm. After complete dissolution using a magnetic stirrer, 3.6 mg (5 mM) of photoinitiator α-ketoglutaric acid was added, and the mixture was stirred in the dark. Once the powder was completely dissolved, 0.5 g of sodium alginate was added, and the mixture was stirred overnight to obtain a uniformly dissolved, slightly yellow, viscous solution (MBAAm pregel solution). Similarly, a BISS crosslinked pregel solution was prepared for the separable substrate of the microneedles. Here, 50 mM disulfide crosslinking agent (BISS) was used instead of 5 mM MBAAm crosslinking agent.

[0077] Fabrication of Separable Microneedles Based on IPN Hydrogel

[0078] The IPN hydrogel employs sequential chemical crosslinking of an SBMA network, followed by alginate network and calcium ion crosslinking, to fabricate a relatively robust microneedle array, as shown in Figure 1A. The microneedles are cast in a commercially available microneedle mold (Blueacre Technology, a 600 μm high PDMS mold, 11×11 array) and centrifuged using a microplate centrifuge to maintain the size, shape, and arrangement of the microneedles.

[0079] In short, 0.5 mL of MBAAm pregel hydrogel solution was added to a microneedle mold and tightly covered with a 3D-printed mold (thermoplastic polyurethane (TPU), 15 mm long, 15 mm wide, and 3 mm high) to prevent the solution from leaking out of the microneedle mold. The mold was then centrifuged at 1,500 rpm for 25 minutes, followed by centrifugation at 3,000 rpm for 10 minutes to ensure the solution completely filled the microneedle mold and was tightly stacked on top. After centrifugation, the upper 3D-printed mold was removed, and the upper solution (bottom of the microneedles) was scraped off. 0.5 mL of BISS pregel hydrogel solution was added to the top of the microneedle mold (bottom of the microneedles), covered with the 3D-printed mold, and centrifuged at 1,500 rpm for 30 minutes. Next, the pregel solution in the microneedle mold was photocrosslinked by irradiation with a full-wavelength fiber optic light source for 60 minutes, followed by ionic crosslinking by immersion in a 1M calcium ion aqueous solution (calcium formate) for 30 minutes. Then, it is placed in the air for 24 hours to evaporate the moisture, and the fully dried and separable microneedles with the desired shape are peeled off from the mold. The process is illustrated in the schematic diagram shown in Figure 1B.

[0080] Example 2: Characterization and analysis of IPN hydrogel

[0081] The formation of the IPN hydrogel was characterized using Raman spectroscopy and FT-IR. For Raman spectroscopy analysis, a small amount of hydrogel sample was placed on a silicon wafer, and representative functional groups in the polymer were detected by scanning Raman spectra between 300 and 3,000 cm⁻¹. A dried hydrogel sample was examined by FT-IR, and the formation of the IPN hydrogel was verified by comparing the characteristic signal shifts of the monomers after polymerization.

[0082] In addition, tensile and compression tests were performed using a universal testing machine (UTM) to investigate the mechanical properties of the hydrogel under different crosslinking types, various monomer molar ratios, and different crosslinking times. The internal structure and surface morphology of the IPN hydrogel and microneedles were observed under an optical microscope and a field emission scanning electron microscope (FESEM, JSM 6500F, JEOL). The physical and chemical properties were evaluated as follows.

[0083] Swelling and Degradation of IPN Hydrogels

[0084] The prepared microneedles were immersed in PBS solution (pH 7.4) at 37°C to evaluate the swelling rate of the hydrogel. The surface morphology changes of the microneedles were observed using digital electron microscopy after 30 minutes, 1 hour, 2 hours, 4 hours, 8 hours, 24 hours, and 48 hours. Similarly, the separability of the disulfide crosslinked substrate of the microneedles was evaluated by examining their degradation rate at different concentrations of DTT and / or EDTA.

[0085] Skin penetration test

[0086] The skin penetration capability of the microneedle array was tested using the skin of C57BL / 6 mice according to the standards of the International Association for the Care and Use of Laboratory Animals (IACUC-16-168). Mice were anesthetized by intramuscular injection of 0.15 mL of Zolctil 50 anesthetic at a concentration of 10 mg / mL. Hair was then carefully removed from the back of the mice at the microneedle application sites using depilatory cream. Next, as shown in Figure 9D, the microneedles were fixed to the top of the mouse skin, and a 10 N vertical downward force was applied using a self-made pusher to insert the needles into the skin. Afterward, the device and microneedles were removed, the skin was stained with trypan blue, and the puncture was observed under a digital electron microscope. Furthermore, the skin was stained with H&E to examine the skin punctured by the microneedles in detail.

[0087] Loading drugs into microneedles

[0088] In the preparation process, doxorubicin (5 mg) and / or LPS (2 mg) were loaded into the gel by mixing with 5 mL of MBAAm pregel hydrogel solution. The drug was dissolved in 5 mL of deionized water and sequentially mixed with SBMA monomer (0.6984 g, 500 mM), MBAAm crosslinking agent (3.8 mg, 5 mM), α-ketoglutarate (3.6 mg, 5 mM) and sodium alginate (0.5 g), and stirred overnight in the dark. Then, using the same steps as for manufacturing microneedles, drug-loaded microneedles (microneedles loaded with DOX, microneedles loaded with LPS, and microneedles loaded with LPS / DOX) were manufactured using a microneedle mold.

[0089] In vitro drug release

[0090] Drug-loaded microneedles were immersed in 10 mL of PBS solution (pH 7.4) and placed in a shaking incubator at 37°C and 100 rpm to simulate a real drug release environment. During the first 4 hours, approximately 2.5 mL of release medium was collected every half hour and an equal volume of PBS was added. Similarly, the released medium was collected at 5, 6, 8, 12, and 24 hours, and then every 24 hours for 7 days. The cumulative drug release was then determined by measuring the UV-Vis absorbance of DOX and LPS at 485 nm and 256 nm, respectively, and calculated using a standard calibration curve for the free drug, as shown in Figures 14A and 14B.

[0091] Biocompatibility

[0092] Mouse glioma cancer cells (CT-2A-Luc cells) and mouse embryonic fibroblasts (NIH-3T3 cells) were used as model cell lines to detect the biocompatibility of the hydrogels using the MTT assay. First, cells were cultured in T-75 culture flasks containing complete culture medium, which included DMEM (90%), fetal bovine serum (FBS) (10%), and antibiotics (1%) in a 37°C, 5% CO2 incubator. When confluence reached 80% to 90%, cells were subcultured into 96-well plates at a density of 1 × 10⁴ cells / well for 24 hours. Simultaneously, hydrogel samples were immersed in fresh culture medium in centrifuge tubes and incubated at 37°C for 24 hours. The concentration was set to 0.1 g / mL according to ISO 10993-12. Next, the old culture medium in the 96-well plates was removed, and the cells were washed with PBS. The culture medium from the hydrogel samples was then added to the cells, followed by incubation for 24 hours. Next, the old medium was replenished with 100 μL of fresh medium containing 3-(4,5-dimethylthiazolyl-2-yl)-2,5-diphenyltetrazolium bromide (MTT) dye (1 mg / mL) and incubated for 4 hours. Then, the MTT-containing medium was removed, and 100 μL of DMSO was added to dissolve the crystals, followed by incubation for 30 minutes. The absorbance of the samples at 570 nm (n=8) was then read using an ELISA microdisc analyzer (Thermo Fisher Scientific, Waltham, USA), and the percentage of cell viability was assessed using the equation shown below.

[0093]

[0094] Animal Experiments

[0095] Seven-week-old female C57BL / 6 mice were purchased from BioLASCO (Taiwan). All animal care and experimental procedures were performed according to the standards of the Laboratory Animal Care and Use Committee of the National Defense Medical Center (IACUC-16-168). Mice were housed in a pathogen-free environment at 25±2℃ and 55±5% humidity with a 12-hour light-dark cycle, and were allowed free access to food and water. Subsequently, after one week of environmental acclimatization, CT-2A-Luc cells (1.5 × 10⁶ cells in 0.1 mL of culture medium) were subcutaneously injected into the right abdomen of each mouse as a developing tumor model. Mice were randomly divided into 5 groups (n=6 per group), and tumor size was measured every 2 days using digital calipers. When the subcutaneous tumor volume reached approximately 30 to 50 mm³ (approximately 10 days after injection), each group of mice was treated with drug-loaded microneedles of different formulations at predetermined time intervals. Group 1 (control group): no treatment; Group 2: treatment with blank microneedles; Group 3: treatment with microneedles loaded with LPS; Group 4: treatment with microneedles loaded with DOX; and Group 5: treatment with microneedles loaded with LPS / DOX at equivalent concentrations of 5 mg / kg LPS and 10 mg / kg DOX. The drug-loaded microneedles were adhered to subcutaneous tumors, and the separable substrate of the microneedles was removed within 24 hours using DTT (60 mM) and EDTA (300 mM) solutions. Throughout the experiment (18 days), tumor size and body weight were measured and recorded every 2 days to assess the antitumor effect or tumor growth inhibition rate in different treatment groups. Tumor volume (V) and tumor growth inhibition rate were calculated using equations (1) and (2) below, respectively.

[0096]

[0097]

[0098] In the formula, W and L represent the shortest and longest tumor diameters, respectively, while Vc and Vt represent the average tumor volumes of the control group and the treatment group at the end of treatment, respectively.

[0099] Finally, mice were sacrificed using cervical dislocation, and spleens and tumor tissues from each group were collected and fixed in 4% (w / v) PBS-buffered paraformaldehyde according to the Darke 2021 protocol. Then, experts from Taiwan TopSheng Technology Co., Ltd. performed tissue staining and immunohistocompatibility analysis.

[0100] Statistical Analysis

[0101] Each measurement should be repeated at least 3 times, and the values ​​should be reported as mean ± standard deviation. Statistical analysis of the variance between groups should be performed using a two-test Student's t-test, with P < 0.05 considered statistically significant. An asterisk indicates statistical significance (*P < 0.05, **P < 0.01, and ***P < 0.001).

[0102] Synthesis and Characterization of Disulfide Crosslinking Agents

[0103] An N,N-bisacrylcysteine ​​(BISS) disulfide crosslinking agent was synthesized to prepare a separable microneedle substrate that can be easily removed from the microneedle array and minimizes discomfort for the microneedle user. After the microneedles are extruded into the skin, the disulfide bonds of the crosslinked polymer forming the microneedle substrate are dissolved by a reducing agent that can break the disulfide bonds (e.g., DTT) or by a chelating agent that chelates Ca2+ ions in the hydrogel alginate network (e.g., EDTA). Therefore, the separable substrate of the microneedles is easily decomposed and removed by the reducing agent or metal chelating agent, while the microneedle array remains in the skin to achieve sustained transdermal drug release.

[0104] The successful preparation of the disulfide crosslinking agent (BISS) was first confirmed by 1H NMR. In Figure 2C, the corresponding proton spectra appear at δ=3.05ppm(a) and 3.4ppm(a'), δ=4.8ppm(b), δ=5.85ppm(c) and 6.3ppm(c'), and δ=6.45ppm(d), verifying the successful synthesis of the disulfide crosslinking agent (BISS). Figures 2A and 2B also show the characteristic peaks of the monomer. As shown in Figure 2A, the signals at δ=2.89 and 3.01ppm(a,a') and δ=3.48ppm(b) are the methylene and methine protons of L-cystine, respectively. Similarly, as shown in Figure 2B, the spectra at δ=5.62ppm(b), δ=5.80ppm(c), and δ=6.14ppm(a) are representative peaks of the acrylonitrile chloride monomer used to synthesize BISS.

[0105] Raman spectroscopy is also used to identify the covalent bonds formed in the disulfide crosslinking agent (BISS). Obviously, as shown in Figure 2D, the characteristic Raman signals at 500 cm⁻¹, 670 cm⁻¹, and 1247 cm⁻¹ in BISS represent SS, CS, and CC bonds, respectively, indicating that the crosslinking agent was successfully synthesized.

[0106] In addition, as shown in Figure 15, the FT-IR results further confirm the successful synthesis of BISS. The characteristic shift of the C=O bond vibration signal from 1810 cm⁻¹ in acrylamide chloride and 1680 cm⁻¹ in L-cystine to 1703 cm⁻¹ in BISS is an indicator of the formation of the crosslinking agent.

[0107] Example 3: Characterization of IPN hydrogel

[0108] Raman spectroscopy and FT-IR confirmed the successful preparation of IPN via sequential free radical polymerization (chemical crosslinking) of SBMA followed by ionic crosslinking of sodium alginate and calcium ions. As shown in Figure 3A, the Raman signal of the IPN hydrogel appears at approximately 1133 cm⁻¹, indicating the asymmetric vibration of SO₃⁻ originating from SBMA. Similarly, characteristic peaks at 807 cm⁻¹, 888 cm⁻¹, and 954 cm⁻¹, respectively, indicate the stretching vibrations of CC, CO, and CCO in sodium alginate. Furthermore, Raman signals at 1240 cm⁻¹ and 1413 cm⁻¹ represent the stretching vibrations of COO⁻ in alginate, confirming the formation of the IPN hydrogel. Clearly, the characteristic Raman spectra of the MBAAm crosslinking agent and the photoinitiator (α-ketoglutarate) are not visible in the IPN hydrogel, indicating that the toxic monomers have completely reacted to form the IPN hydrogel, and that any potential harmful effects of the hydrogel are negligible.

[0109] The IPN hydrogel was further characterized using FT-IR spectroscopy, which can be used to detect the functional groups of the polymer before and after polymerization. As shown in Figure 3B, the chemical shift of the COO signal from 1595 cm⁻¹ in sodium alginate to 1642 cm⁻¹ in the IPN hydrogel indicates the formation of crosslinks with calcium ions (-COOCa). Furthermore, the sharp signal at 1720 cm⁻¹ in the IPN hydrogel is due to the vibration of C=O in the SBMA monomer, clearly demonstrating the chemical crosslinking of the hydrogel. The disappearance of the representative FT-IR signals of the photoinitiator and MBAAm crosslinking agent after polymerization further confirms the formation of the IPN hydrogel.

[0110] Morphology of IPN hydrogels

[0111] The internal structure of IPN hydrogels formed by crosslinking SBMA monomers and sodium alginate with different molar ratios was studied by field emission scanning electron microscopy (FESEM). Keeping other parameters constant (60 min of photocrosslinking with MBAAm and 30 min of crosslinking with 1M Ca2+ ions), the molar ratio of SBMA:alginate was increased from 1:1 to 2:1 and 3:1, respectively, by reacting 0.5M, 0.67M, and 0.75M SBMA with 0.5M, 0.33M, and 0.25M sodium alginate.

[0112] FESEM images of the IPN hydrogel show a porous structure, primarily due to the polymerization of SBMA. As shown in Figures 4A to 4C, the size of the hydrogel pores increases with increasing SBMA:alginate molar ratio. One possible reason is that the amount of SBMA monomers polymerized is limited within a fixed time interval (60 minutes) of photopolymerization, thus reducing the crosslinking density of the hydrogel with increasing SBMA proportion. Since the pore size of the IPN hydrogel is inversely proportional to the crosslinking density of the copolymer, a relatively large pore size ratio of 3:1 was observed (Figure 4C). On the other hand, the ionic crosslinking density of the IPN hydrogel decreases with decreasing sodium alginate proportion, resulting in increased internal porosity. However, hydrogels with better mechanical properties are needed to fabricate strong microneedles that penetrate the skin, and further experiments were conducted using hydrogels with a higher crosslinking density (1:1 ratio).

[0113] Therefore, SBMA and alginate monomers in a 1:1 molar ratio were used to fabricate microneedles using a microneedle mold, and the shape of the microneedles was confirmed using optical microscopy and FESEM imaging. As shown in Figures 5A to 5C, optical and FESEM images of the dried microneedles confirmed the formation of a uniform size and shape of a microneedle array with a specified arrangement. Energy dispersive spectroscopy (EDS) analysis of the FESEM images also confirmed the uniform distribution of the polymer in the microneedle array. The tips of the microneedles are quite long; for example, the needle height is at least 600 μm, the needle base is at least 300 μm, the needle tip spacing is at least 600 μm, and the array number is at least 11 × 11 = 121. The tips of the microneedles are sharp enough to penetrate the epidermis of the skin.

[0114] Table 1: Specifications of microneedle arrays prepared from IPN hydrogels

[0115] Mechanical strength of IPN hydrogel

[0116] The mechanical properties of IPN hydrogels under different conditions were studied using a universal testing machine (UTM), such as the type of crosslinking (ionic crosslinking, photocrosslinking, or a combination of both) and the molar ratio of the monomers used for polymerization.

[0117] As shown in Figure 6A and Table 2 below, the mechanical properties of IPN hydrogels with different crosslinking methods were evaluated. Compared with hydrogels formed by ionic crosslinking (1M Ca2+ for 30 minutes) (2.302 MPa) or sequential photoionic crosslinking (2.71 MPa), the hydrogel formed by photocrosslinking (UV for 60 minutes) exhibited the highest maximum tensile strength of 3.449 MPa. However, its strain was lower (20%) compared with hydrogels formed by ionic crosslinking (113%) or sequential photoionic crosslinking (80%). In this paper, the hydrogel formed solely by chemical crosslinking networks exhibited extremely high mechanical structure, high tensile strength but brittleness. Conversely, the hydrogel formed solely by ionic crosslinking had the highest elongation (strain = 113%) but the lowest tensile strength (stress = 2.301 MPa), indicating that physically crosslinked hydrogels have low mechanical strength, soft structure, and outstanding toughness. In essence, the hydrogels formed by sequential photopolymerization (chemical crosslinking) and ionic interaction (physical crosslinking) disclosed herein exhibit the best tensile strength (2.71 MPa) and strain (81%) and the highest Young's modulus (1.32 MPa), which are strong and hard enough to allow needles to penetrate the skin.

[0118] The mechanical properties of IPN hydrogels with different molar ratios of SBMA and sodium alginate monomers (1:1, 2:1, and 3:1) were further examined. The tensile test in Figure 6B depicts that the mechanical strength and toughness of the IPN hydrogel gradually decrease as the SBMA:sodium alginate ratio increases from 1:1 to 2:1 and 3:1, which is attributed to the density of the crosslinking network. As shown in Figure 4, in the FESEM results of the IPN hydrogel, the density of the crosslinking network decreases with increasing SBMA:sodium alginate ratio, followed by an increase in internal pore size. As can be seen from Figure 6B and Table 2, the 1:1 molar ratio monomers in this disclosure have a better crosslinking network density, resulting in hydrogels with relatively good mechanical strength (2.71 MPa maximum tensile strength, 81% strain, and 1.31 MPa Young's modulus). Generally speaking, sequential photocrosslinking (60 minutes) and ionic crosslinking (1M Ca2+ for 30 minutes) of SBMA and sodium alginate in a 1:1 molar ratio will produce a dense IPN hydrogel with better mechanical strength while maintaining good extensibility.

[0119] Table 2: Mechanical properties of IPN hydrogels under different crosslinking conditions and monomer ratios

[0119] 1:1-UV 60=1:1 molar ratio of SBMA and sodium alginate photopolymerization for 60 minutes.

[0119] 1:1 Ca2+ 30 = 1:1 molar ratio of SBMA and sodium alginate crosslinked with Ca2+ (1M) ions for 30 minutes

[0119] D / t = different

[0120] In addition, as shown in Figure 6C, the microneedle formulation prepared from IPN hydrogel with 1:1-Uv60, Ca2+ 30 (i.e., monomer ratio 1:1 and photocrosslinking for 60 minutes, followed by Ca2+ ion crosslinking for 30 minutes) can withstand a compressive stress of 76.8 N (0.64 N / needle), confirming that the microneedles effectively penetrate the skin. This sequential double crosslinking of the hydrogel provides a stronger microneedle array (~0.15 N / needle) than microneedles made from PLA / PLGA copolymer (Li et al., Nat Biomed Eng 3(3)(2019) 220-229) and PVA / PVP copolymer (~0.17 N / needle) (Song et al., ACS Biomater Sci Eng 6(7)(2020) 4116-4125).

[0121] Expansion and Degradation Characteristics

[0122] The swelling behavior of the microneedles was assessed by immersing them in a PBS solution (pH 7.4) at 37°C and observing morphological changes under a microscope. As shown in Figure 7, the microneedles absorb water and swell, and their structure gradually loosens over time. After 24 hours, the internal structure of the IPN hydrogel disintegrates, and most of the microneedles collapse. This swelling and subsequent collapse of the microneedle structure is attributed to the presence of sodium ions in the PBS solution, which is an ion exchange reaction with the calcium alginate hydrogel network. Calcium ions on the COO- groups of alginate in the IPN hydrogel are gradually replaced by sodium ions, increasing the electrostatic repulsion on the COO- groups. As a result, the interpenetrating network of the hydrogel gradually loosens and the hydrogel swells, followed by the collapse of the microneedle structure. This is a fundamental phenomenon for the stable transdermal drug release from microneedles.

[0123] Separability of the disulfide cross-linked substrate of microneedles

[0124] The separability of the disulfide crosslinked substrate of the microneedles depends on the decomposition rate of the disulfide crosslinked IPN hydrogel. The hydrogel was immersed in a reducing agent (DTT) and / or a metal chelating agent (EDTA) solution, and the remaining weight was measured at different time intervals of 1 hour to determine the disintegration rate of the separable disulfide crosslinked IPN hydrogel. The degradation rate was examined at different monomer:crosslinker ratios with a fixed disintegrant concentration or at different disintegrant concentrations while maintaining the monomer:crosslinker ratio.

[0125] The effect of crosslinking agent concentration on the degradation rate of separable IPN hydrogels is shown in Figures 8A and 8B. Hydrogels with a crosslinking agent concentration of 1% were immersed in a mixture of DTT and EDTA (DTT 20 mM, EDTA 100 mM) and completely disintegrated after 1 hour. However, when the crosslinking agent concentration was one-tenth of the monomer concentration, approximately 41% of the hydrogel remained within the same time interval (1 hour). Hydrogels with high crosslinking agent concentrations have a relatively dense internal structure and are less prone to disintegration in a short time. However, at low crosslinking agent concentrations, the increased porosity and surface area within the internal structure lead to a higher disintegration rate of the IPN hydrogel.

[0126] The degradation rate of separable hydrogels prepared with a crosslinking agent to monomer ratio of 1:100 was also studied in different aqueous environments, including PBS, DTT 20mM, DTT 40mM, DTT 60mM, EDTA 100mM, EDTA 200mM, EDTA 300mM, and EDTA 300mM + DTT 60mM aqueous solutions. As shown in Figures 8C and 8D, the hydrogel decomposed least in the PBS solution, and during a 30-minute incubation period, less than 10% of the hydrogel remained in the mixed solution of EDTA 300mM and DTT 60mM. Therefore, the disintegration rate of disulfide-linked IPN hydrogels and the separability of microneedles can be improved by adjusting the concentration of the disintegrant.

[0127] Skin Penetration Test

[0128] Microneedle penetration was tested using skin from C57BL / 6 mice. Microneedles were inserted into the skin using a fixed force application device and stained with trypan blue. The circular pores on the skin were observed under an optical microscope (Fig. 9B). H&E staining was used to further examine the histology of the skin tissue after microneedle penetration to investigate the depth of penetration. As shown in Fig. 9C, the outer layer of skin, possibly the stratum corneum, was completely penetrated by the microneedles (red arrow). This result indicates that the IPN hydrogel-based microneedle array has sufficient mechanical strength to penetrate the outer layer of mouse skin for effective transdermal application. As shown in Fig. 9E, the shape of the microneedle array did not show significant deformation after penetration, further demonstrating the high mechanical strength of the microneedles.

[0129] In vitro drug delivery and release

[0130] The drug release behavior of IPN hydrogels was investigated. First, DOX and / or LPS were loaded into the hydrogel by simple mixing with an IPN hydrogel pregel solution (0.5 mg DOX / microneedles and 0.2 mg LPS / microneedles), followed by casting microneedles and sequential photoionization crosslinking, to obtain drug-loaded solid microneedle arrays (DOX-loaded microneedles, LPS-loaded microneedles, and LPS / DOX-loaded microneedles). Next, the drug-loaded microneedles were immersed in 10 mL of PBS solution and kept in a rotary shaker incubator at 37°C and 100 rpm. As shown in Figures 10A and 10B, rapid release of DOX and LPS was observed in the first 4 hours due to the high diffusion rate of the drug on the surface of the microneedle array, followed by stable release due to microneedle degradation. The degradation of the IPN hydrogel-based microneedles in the PBS medium was mainly attributed to sodium-calcium exchange in the hydrogel, which enhanced the swelling and disintegration of the hydrogel to promote drug diffusion. Sodium ions in PBS slowly exchange with calcium ions in the alginate fragments of the IPN hydrogel, causing the cross-linking agent to disintegrate and the hydrogel structure to eventually collapse completely, as shown in Figure 7, releasing residual drugs, DOX, LPS, and other contents from the microneedles. However, as shown in Figure 7, the rapid degradation of microneedles in PBS is accompanied by the release of approximately 66.1±7.4% and 59.4±5.5% of DOX and LPS, respectively, within 24 hours, with the remaining drugs released over approximately 7 days. This timely drug release is one of the advantages of delivering sufficient doses of drug to the target site for effective treatment. Compared to DOX, LPS release is relatively slower, partly due to its larger molecular weight, and the chemical structure of repeating carboxyl and phosphate groups in LPS may form physical cross-links in the presence of calcium ions, leading to slower release.

[0131] Biocompatibility of IPN hydrogels

[0132] The biocompatibility of hydrogels is an essential concern for in vivo applications. Therefore, the biocompatibility of hydrogels was evaluated in vitro using MTT assays targeting CT-2A-Luc and NIH-3T3 cells. Cells were treated with the hydrogel after co-culturing with the hydrogel extract, followed by the addition of MTT dye, and the biocompatibility was determined based on the absorbance intensity of purple crystals at 570 nm in the mitochondria of living cells. Therefore, even at high concentrations (0.1 g / mL) of hydrogel samples, the survival rate of cells treated with BISS or MBAAm cross-linked hydrogels was greater than 85% in CT-2A-Luc cells and greater than 92.6% in NIH-3T3 cells, indicating that the hydrogels have significant biocompatibility and do not affect cells after transdermal drug delivery, as shown in Figure 16.

[0133] Cytotoxicity of DOX-loaded microneedles

[0134] The anticancer effect of DOX-loaded microneedles on CT-2A-Luc and NIH-3T3 cells was evaluated using MTT assay. As shown in Figure 11A, with increasing DOX concentration, cell survival rate (%) significantly decreased, resulting in cell survival rates of 35.7±4.7% for CT-2A-Luc and 41.1±3.4% for NIH-3T3 cells, indicating the effectiveness of DOX-loaded microneedles in treating cancer cells at comparable concentrations. Furthermore, the half-maximal inhibitory concentration (IC50) of DOX-loaded microneedles on CT-2A-Luc cells was assessed using a Gaussian formula, which showed an IC50 of 1.088 μg / mL and 3.23 μg / mL on NIH-3T3 cells (Figure 11B). A lower IC50 value indicates a stronger inhibitory effect on cell growth. Here, unlike cancer cells (CT-2A-Luc cells), NIH-3T3 cells have a higher IC50 value for DOX, or exhibit better cell survival at the same concentration of DOX, indicating that DOX internalization is enhanced in cancer cells compared to normal cells.

[0135] In vivo antitumor efficacy of dual-drug-loaded microneedles

[0136] The efficiency of microneedle-mediated percutaneous co-delivery of LPS and DOX and their synergistic effect on tumor growth inhibition were evaluated in mice with C57BL / 6 gliomas. Clearly, direct intratumoral (IT), peritumoral (PT), or intravenous (IV) injection of tumor-inhibiting agents can lead to drug leakage into nearby tissues, thereby reducing therapeutic efficacy and safety. However, microneedle-mediated local percutaneous drug delivery can avoid these problems by improving drug diffusion into deeper tumors and leading to high tumor accumulation. Furthermore, microneedles play an important role in combination therapy, with multiple drugs loaded into microneedles and applied continuously to achieve synergistic value. Due to the accumulation of high-density immune cells including Langerhans cells (LC), dendritic cells (DC), and macrophages in skin tissue, microneedle-mediated percutaneous co-delivery of LPS and DOX will have a potent synergistic effect on immunochemotherapy in cancer treatment, as the immunostimulatory molecule LPS has a good opportunity to interact with and activate these immune cells to enhance the therapeutic effect of the chemotherapeutic drug DOX.

[0137] As described above, drug-loaded separable microneedles (microneedles loaded with LPS, microneedles loaded with DOX, and microneedles loaded with LPS / DOX) and blank separable microneedles were applied to subcutaneous tumors to deliver anticancer drugs percutaneously. After 24 hours, the separable substrate of the microneedles was removed with a small amount of EDTA (300 mM) and DTT (60 mM), below the minimum toxicity dose. Then, as shown in Figure 12B, tumor growth was monitored over 18 days by measuring external tumor volume. The drug-loaded microneedle array penetrated the stratum corneum of the skin and released the drug locally and stably to effectively inhibit the tumor. Finally, the tumor growth profile for each group is summarized in Figures 12A and 12B.

[0138] As shown in Figure 12A, mice treated with drug-loaded microneedles showed significant inhibition of tumor growth compared to the control group (untreated) or mice treated with blank microneedles that showed rapid tumor growth. At the end of the treatment period (18 days), the mean tumor volumes of mice treated with LPS-loaded microneedles, DOX-loaded microneedles, and LPS / DOX-loaded microneedles were 552.4±205.6 mm3, 345.4±220.2 mm3, and 253.0±176.9 mm3, respectively, which were significantly lower than those of mice treated with the control group (1696.8±447.6 mm3) or blank microneedles (1468.31±460.68 mm3). Notably, tumor growth in mice treated with dual-drug therapy (LPS / DOX-loaded microneedles) was even more significantly inhibited than that in mice treated with DOX-loaded microneedles. This can be attributed to the synergistic effect of chemotherapy drugs (DOX) and the significant immunostimulatory effect of LPS, accompanied by the upregulation of different anti-cancer immune cells. The tumor photographs in Figure 12B also show a parallel trend where tumors treated with LPS / DOX-loaded microneedles were significantly smaller than those with other formulations. Furthermore, the tumor weight shown in Figure 12C further validates that the dual-drug approach (LPS / DOX-loaded microneedles) has better tumor growth inhibition than either the single-drug approach (LPS-loaded microneedles or DOX-loaded microneedles). The mean body weight (21.8 ± 1.1 to 23.3 ± 0.5 g) in the different treatment groups did not change significantly during treatment, and no animals died. However, as shown in Figure 12D, unlike the control group, the body weight of mice treated with microneedles decreased significantly during the first two days of treatment, then increased in parallel with the control group. This was due to the anesthetic injection during the fixation of the separable microneedles to the tumor and the discomfort caused by the microneedles to the mice until the separable matrix was removed within 24 hours. The separable microparticles presented herein are convenient and effective for the transdermal delivery of anticancer drugs in the vicinity, and provide a method for minimizing systemic toxicity and side effects of the drugs.

[0139] Immunohistochemistry

[0140] LPS can inhibit dendritic cell (DC) apoptosis and enhance dendritic cell-mediated CD4+ T cell proliferation, and upregulate the expression of CD80, CD86, CD69, and CD25, as well as the secretion of cytokines such as TNF-α and IL-6. Therefore, to elucidate the immunostimulatory effect of LPS on in vivo anticancer activity, spleen and tumor tissues from each group were collected and stained to detect CD25+, CD4+, CD69+, CD8+, TNF-α, and activated caspase 3 levels as representative immunomodulatory biomarkers. Histopathological data were obtained by examining H&E-stained slides using an optical microscope. The percentage of organs affected by proliferative and non-proliferative changes in experimental animals was based on the International Harmonization of Nomenclature and Diagnostic Criteria (INHAND).

[0141] CD25+ is a type I transmembrane protein present on activated T cells and activated B cells, and is a reliable immunohistochemical marker for the upregulation of immune cells. Therefore, CD25+ was mainly highly expressed in mice treated with LPS containing microneedles, with positive signals mainly appearing in the splenic sinuses of the red pulp, but a small number of signals were also found in the germinal centers of the spleen. The positive signal in the microneedle treatment group loaded with LPS / DOX was mainly higher than that in other groups. As shown in Figure 13A, CD4+ and CD69+ T cells were significantly upregulated in the spleen tissue of mice treated with LPS (microneedles loaded with LPS and microneedles loaded with LPS / DOX), while they were relatively low in the control group, indicating that LPS activates and proliferates major antigen-presenting cells, B cells. Positive signals were frequently found in the periarteriolar lymphoid sheath (PALS) of the germinal centers or in the splenic sinuses of the red pulp. In LPS-treated tumor tissues, the expression of CD8+ T cells and TNF-α was significantly increased, accompanied by tumor growth inhibition, indicating that LPS can locally enhance the immune cells in cancer immunotherapy. LPS stimulates monocytes to transform into M1 macrophages, which then play an important role in eliminating cancer cells indirectly by directly attacking or recruiting other immune cells or secreting various cytokines. Positive CD8+ signals were more commonly found in necrotic tissue and the peritumoral region, while TNF-α was mainly present in tumor cells and inflammatory cells in the peritumoral tissue. Although no tumor cells were found in the microneedles loaded with LPS / DOX, significantly higher TNF-α signals were observed in some inflammatory cells in the connective tissue, indicating that the combined drugs effectively eliminated cancer cells.

[0142] As shown in Figure 13B, lysed caspase 3 or activated caspase 3 is another biomarker of apoptotic cells, capable of degrading various cellular proteins and DNA fragments during apoptosis. This was detected by the IHC test shown in Figure 13B, indicating overexpression of lysed caspase 3 in tumor tissues of groups 3, 4, and 5. This confirms that cancer cells primarily respond to the combined drug (LPS / DOX / microneedles) to undergo apoptosis, and tumor progression is significantly inhibited. The results, as demonstrated in Figure 12 above, show a consistent reduction in tumor size.

[0143] Positive signals are mostly found in the central and necrotic areas, with stronger signals in the cytoplasm and nucleus. Recently, despite the significant clinical success of chemotherapy, its effectiveness is strongly influenced by the heterogeneous tumor microenvironment. On the other hand, immunotherapy or cancer vaccination, which relies on modulating the patient's immune system to recognize and destroy malignant cells, has become a paradigm shift in cancer treatment. However, most biomolecules injected into the circulatory system as vaccines, such as peptides and nucleic acids, are more susceptible to enzymatic degradation in body fluids. Therefore, the ingenious use of microneedles as carriers for percutaneous delivery of immunostimulatory molecules or vaccines represents a major breakthrough in cancer immunotherapy, in which biomolecules are protected from enzymatic attack and delivered to the skin, where antigen-presenting cells (APCs) are densely packed.

[0144] Here, the microneedle array penetrates the stratum corneum and delivers LPS to the immune-cell-rich epidermis, triggering a broad immune response, including the maturation of dendritic cells. Subsequently, the maturation of dendritic cells can activate B cells and CD8+ T cells for humoral and cellular immunity, respectively, and provide synergistic value for chemotherapy. Therefore, the microneedle-mediated percutaneous delivery system is an attractive cancer immunogenicity and therapeutic approach.

Claims

1. A transdermal delivery device, comprising: The plurality of protrusions each comprise a first polymer formed of a first monomer having a first bond and a second polymer formed of a second monomer having a second bond; The substrate comprises a third polymer formed from the first monomer having a third bond; and A bioactive agent, which is contained in one of the plurality of protrusions. in, The plurality of protrusions are coupled to the substrate and configured to be at least partially insertable into the skin of an individual in need, and to target and destroy the compound forming the third polymer in the substrate to remove the substrate after the percutaneous delivery device has been applied to the skin for a predetermined time.

2. The transdermal delivery device as claimed in claim 1, wherein, The first monomer is a zwitterion.

3. The transdermal delivery device as claimed in claim 2, wherein, The zwitterion is phosphatidylcholine, sulfobetaine, alkyl sulfonate pyridinium salt, carboxybetaine, phosphate betaine, phosphatidyl betaine, phosphatidyl betaine, ammonium sulfate, aminosulfonamide, pyridylcarboxylate, or sulfobetaine.

4. The transdermal delivery device as claimed in claim 3, wherein, The phoscholine is phoscholine acrylate, phoscholine acrylamide, phoscholine methyl acrylate, alkoxydicyanvinyl alcohol ester, or 2-methacryloyloxyethyl phoscholine.

5. The transdermal delivery device as claimed in claim 3, wherein, The sulfobetaine is sulfobetaine acrylate, sulfobetaine acrylamide, sulfobetaine methyl acrylate, sulfobetaine vinylimidazolium, or sulfobetaine vinylpyridine.

6. The transdermal delivery device as claimed in claim 3, wherein, The carboxybetaine is carboxybetaine acrylate, carboxybetaine methyl acrylate, carboxybetaine acrylamide, carboxybetaine vinylimidazolium, carboxybetaine methacrylamide, carboxybetaine isobutylene, or carboxybetaine diallylamine.

7. The transdermal delivery device as claimed in claim 3, wherein, The alkyl sulfonate pyridine salt is 3-(2-vinylpyridinium-1-yl)propane-1-sulfonate, N-(2-methacryloxy)ethyl-N,N-dimethylammonium propanesulfonate, or N-(3-methacryloxyimino)propyl-N,N-dimethylammonium propanesulfonate.

8. The transdermal delivery device as claimed in claim 1, wherein, The second polymer is one or more selected from the group consisting of biocompatible synthetic polymers, semi-synthetic polymers and natural polymers.

9. The transdermal delivery device as claimed in claim 1, wherein, The second polymer is selected from gums, polysaccharides, polysaccharide derivatives, alginates containing sodium alginate or calcium alginate, chitosan, chitosan derivatives, collagen, gelatin, polydextrose, poly(vinylpyrrolidone), hydroxyethyl starch, polyethylene glycol, functionalized polydextrose, trehalose-containing sugar polymers, hyaluronic acid, methacrylated hyaluronic acid, poly(methyl vinyl ether), poly(methyl vinyl ether-alt-maleic anhydride), poly(lactic acid), polyglycolic acid, poly(lactic acid-glycolic acid copolymer), polycarbonate, poly(vinyl alcohol), poly(hydroxyethyl methacrylate), poly(vinylpyrrolidone), (2- The group consisting of (2-carboxymethyl)-3-propenylaminopropyl dimethylammonium bromide, (2-carboxymethyl)-3-propenylaminopropyl dimethylammonium bromide-co-hydroxyethyl methacrylate, (2-carboxymethyl)-3-propenylaminopropyl dimethylammonium bromide-co-propenylamine, methacrylated (2-carboxymethyl)-3-propenylaminopropyl dimethylammonium bromide-co-propenylamine, poly(ε-caprolactone)poly(ε-caprolactone-co-glycolic acid), poly(2-methacryloyloxyethyl phosphocholine), poly(carboxybetaine)vinylimidazolium, poly(sulfobetaine)vinylimidazolium, and poly(sulfobetaine)vinylpyridine.

10. The transdermal delivery device as claimed in claim 1, wherein, The first monolithic system is cross-linked by chemical bonds, and the second monolithic system is cross-linked by physical bonds.

11. The transdermal delivery device as claimed in claim 10, wherein, The chemical and physical bonds form an interpenetrating polymer network.

12. The transdermal delivery device as claimed in claim 10, wherein, The chemical bond is formed by polymerizable compounds selected from N,N'-methylenebisacrylamide (MBA), bisacrylamide derivatives of cystine (BISS), dimethyl octyldiimide, glutaraldehyde, N,N-ethylene-bis(iodoacetamide), ethylene glycol dimethacrylate (EGDM), poly(ε-caprolactone) diacrylate, polylactic acid diacrylate, polylactic acid dimethacrylate, poly(lactic-co-glycolic acid) diacrylate, poly(lactic-co-glycolic acid) dimethacrylate, poly(ε-caprolactone-β-ethylene glycol-β-ε-caprolactone) diacrylate, ethylene glycol-β-(lactic-co-glycolic acid) dimethacrylate, including disulfide bonds, peptide bonds, or ester bonds. Poly(ε-caprolactone) dimethacrylate (MAC-PCL-MAC), poly(ε-caprolactone-β-ethylene glycol-β-ε-caprolactone) dimethacrylate (MAC-PCL-PEG-PCL-MAC), poly(lactic acid-β-ethylene glycol-β-lactic acid) diacrylate (AC-PLA-PEG-PLA-AC), poly(lactic acid-β-ethylene glycol-β-lactic acid) dimethacrylate (MAC-PLA-PEG-PLA-MAC), poly[(lactic acid-co-glycolic acid)-β-ethylene glycol-β-(lactic acid-co-glycolic acid)] diacrylate (AC-PLGA-PEG-PLGA-AC), poly[(lactic acid-co-glycolic acid)-β-ethylene glycol-β-(lactic acid-co-glycolic acid)] diacrylate (AC-PLGA-PEG-PLGA-AC), poly[(lactic acid-co-glycolic acid)- Poly(ε-caprolactone-co-glycolic acid)-β-ethylene glycol-β-(lactic acid-co-glycolic acid) dimethacrylate (MAC-PLGA-PEG-PLGA-MAC), poly(ε-caprolactone-co-glycolic acid)-diacrylate (AC-PCLA-AC), poly(ε-caprolactone-co-glycolic acid)-dimethacrylate (MAC-PCLA-MAC), poly(ε-caprolactone-co-glycolic acid)-diacrylate (AC-PCGA-AC), poly(ε-caprolactone-co-glycolic acid)-dimethacrylate (MAC-PCGA-MAC), poly(ε-caprolactone-co-glycolic acid)-β-ethylene glycol-β-(ε-caprolactone-co-glycolic acid)-diacrylate (... It is formed by at least one crosslinking agent from the group consisting of AC-PCLA-PEG-PCLA-AC, poly(ε-caprolactone-co-lactic acid)-β-ethylene glycol-β-(ε-caprolactone-co-lactic acid) dimethacrylate (MAC-PCLA-PEG-PCLA-MAC), poly(ε-caprolactone-co-glycolic acid)-β-ethylene glycol-β-(ε-caprolactone-co-glycolic acid) diacrylate (AC-PCGA-PEG-PCGA-AC), and poly(ε-caprolactone-co-glycolic acid)-β-ethylene glycol-β-(ε-caprolactone-co-glycolic acid) dimethacrylate (MAC-PCGA-PEG-PCGA-MAC).

13. The transdermal delivery device as claimed in claim 10, wherein, The ratio of the first monomer to the second monomer is approximately 1:1, approximately 1:2, approximately 1:3, approximately 1:4, approximately 1:5, approximately 2:1, approximately 3:1, approximately 4:1, or approximately 5:

1.

14. The transdermal delivery device as claimed in claim 1, wherein, The third bond is a disulfide bond.

15. The transdermal delivery device as claimed in claim 1, wherein, The compound used to target and disrupt the third bond forming the third polymer in the substrate is dithiothreitol (DTT), ethylenediaminetetraacetic acid (EDTA), glutathione (GSH), β-mercaptoethanol, or L-cysteine.

16. The transdermal delivery device as claimed in claim 1, wherein, Each of the multiple protrusions has a gradually tapering shape.

17. The transdermal delivery device as claimed in claim 16, wherein, Each of the multiple protrusions is conical or cone-shaped.

18. The transdermal delivery device as claimed in claim 16, wherein, Each of the plurality of protrusions has a height between about 25 μm and about 2,500 μm, a width between about 50 μm and about 250 μm, and a tip diameter between about 1 μm and about 25 μm.

19. A method of manufacturing a transdermal delivery device, comprising: A first solution comprising a first monomer, a second monomer, a first crosslinking agent, and at least one bioactive agent is prepared. Prepare a second solution comprising the first monomer and the second crosslinking agent; A first solution containing at least one bioactive agent is applied to a mold, and then the first solution is centrifuged. After centrifugation, the upper layer of the first solution is removed and the second solution is applied to the top of the mold; Cover the mold with a capping mold for centrifugation; Apply first conditions suitable for causing the first solution to solidify to form a plurality of protrusions; Apply second conditions suitable for causing the second solution to solidify to form a substrate; as well as The plurality of protrusions and the base are demolded from the mold to obtain the transdermal delivery device.

20. A method for inducing biological activity in individuals in need, comprising: Provide a transdermal delivery device as described in claim 1; The percutaneous delivery device is applied to the individual's skin so that the plurality of protrusions pierce the individual's skin; and The compound used to target and disrupt the third bond forming the third polymer in the substrate removes the substrate of the transdermal delivery device from the individual, while the plurality of protrusions containing the bioactive agent remain in the skin.

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