Method for manufacturing biosensing microneedle, and biosensing microneedle manufactured using same
The method of forming a porous electrode layer and enzyme layer on microneedles addresses manufacturing inefficiencies and inaccuracies in conventional glucose monitoring devices, achieving stable and cost-effective glucose measurement.
Patent Information
- Authority / Receiving Office
- WO · WO
- Patent Type
- Applications
- Current Assignee / Owner
- Filing Date
- 2025-08-28
- Publication Date
- 2026-04-02
AI Technical Summary
Conventional microneedles for glucose monitoring face challenges such as increased manufacturing costs due to prolonged drying times for glucose oxidase coating, inaccurate glucose measurement due to sweat interaction, limited enzyme lifespan, and the need for temperature-controlled storage, which complicates the manufacturing process.
A method involving the formation of a porous electrode layer by depositing a conductive material, followed by electrodepositing an enzyme-immobilizing polymer to create a stable enzyme layer, and optionally forming a platinum reaction layer using a potentiostat to enhance measurement accuracy and reduce manufacturing time.
This approach enables long-term, accurate glucose measurement with minimized manufacturing costs by ensuring stable enzyme reaction and efficient signal reception, while reducing the need for separate oxidizing enzymes on the microneedle surface.
Smart Images
Figure KR2025013186_02042026_PF_FP_ABST
Abstract
Description
Method for manufacturing a biosensing microneedle and a biosensing microneedle manufactured thereby
[0001] The present invention relates to a method for manufacturing a biosensing microneedle that is inserted into the skin to enable monitoring of health conditions, such as glucose detection, and to a biosensing microneedle manufactured thereby.
[0002] Generally, when human blood glucose levels exceed the normal range, it is classified as hyperglycemia, and when they fall below the normal range, it is classified as hypoglycemia. Hyperglycemia is associated with diabetes and requires continuous management.
[0003] To diagnose diabetes and manage it to prevent it from progressing to complications, systematic blood glucose monitoring and treatment must be carried out in parallel.
[0004] To this end, microneedle continuous glucose monitors are used in recent years to measure blood glucose levels in real time.
[0005] These microneedle continuous glucose monitors measure blood glucose levels by inserting needles measuring hundreds of micrometers into the dermis layer to receive electrical signals without damaging blood vessels or nerves. Therefore, the surface of the microneedles is coated with glucose oxidase to enable the reception of blood glucose levels as electrical signals.
[0006] However, in the case of conventional microneedles, when glucose oxidase is coated on the surface of the microneedles, the glucose oxidase is dropped onto the microneedles and then dried, which leads to an increase in manufacturing / processing costs due to the increased drying time and the resulting increase in the overall manufacturing process time.
[0007] In addition, there is a problem in that it is difficult to accurately measure the glucose concentration in the body as glucose oxidase present on the surface of the microneedle substrate that comes into contact with the skin comes into contact with sweat, converting the glucose present in the sweat into hydrogen peroxide.
[0008] Therefore, conventionally, in order to accurately measure the concentration of glucose present in the body, only the microneedle inserted into the body comes into contact with the interstitial fluid, and glucose is converted into hydrogen peroxide by glucose oxidase immobilized on the surface of the microneedle, and then the current generated as the hydrogen peroxide is oxidized in platinum must be measured.
[0009] In addition, conventional microneedles coated with oxidases have the problem that the storage period is limited due to the limited lifespan of the enzyme itself, and that separate temperature correction is required because the enzyme's activity is sensitive to temperature.
[0010] In addition, since the storage temperature of the enzyme is critical, temperature must be considered during the manufacturing process; however, there is a limitation in that temperature maintenance facilities are required to store the enzyme, leading to an increase in manufacturing process costs.
[0011] Related technologies regarding such conventional microneedles and methods for manufacturing the same are disclosed in Korean Patent Publication No. 10-2020-0113998 (October 7, 2020) and Korean Patent Publication No. 10-2023-0043276 (March 31, 2023).
[0012] The purpose of the present invention is to provide a method for manufacturing a biosensing microneedle capable of long-term use and improved measurement accuracy through a stable reaction with glucose without the need for a separate oxidizing enzyme on the surface of the microneedle, and to provide a biosensing microneedle manufactured thereby.
[0013] In addition, the present invention aims to provide a method for manufacturing biosensing microneedles and a biosensing microneedle manufactured thereby, which can minimize manufacturing costs while increasing measurement accuracy through accurate electrical signal reception by coating the surface of the microneedle with an oxidizing enzyme in a short period of time.
[0014] The present invention provides a method for manufacturing a microneedle for biosensing, comprising: a needle portion forming step of forming a needle portion on a base portion; an electrode layer forming step of forming a porous electrode layer by depositing a conductive material on the surface of the needle portion; and an enzyme layer forming step of forming an enzyme layer by electrodepositing an enzyme mixed with an enzyme-immobilizing polymer on the surface of the electrode layer.
[0015] According to another aspect, the present invention provides a method for manufacturing a microneedle for biosensing, comprising a needle portion forming step of forming a needle portion on a base portion, an electrode layer forming step of forming an electrode layer of a conductive material on the surface of the base portion and the needle portion, and a reaction layer forming step of forming a porous reaction layer by depositing platinum on the surface of the electrode layer, wherein in the reaction layer forming step, the deposition of platinum is performed by a potentiostat, and the charge density generated in the potentiostat during the deposition of platinum is set based on the electrochemical surface area (ECSA) of the electrode layer.
[0016] The method for manufacturing a biosensing microneedle according to the present invention and the biosensing microneedle manufactured thereby form a porous electrode layer by depositing a conductive aqueous material on the surface of the needle portion, and then forming an enzyme layer by electrodepositing an enzyme mixed with an enzyme-immobilizing polymer on the surface of the electrode layer. By uniformly and quickly forming the enzyme layer on the surface of the needle portion, measurement accuracy can be increased through stable electrical signal reception via the enzyme, and manufacturing costs can be minimized.
[0017] In addition, the method for manufacturing a bio-sensing microneedle according to the present invention and the bio-sensing microneedle manufactured thereby form an electrode layer by depositing a conductive material made of gold on the surface of the needle portion and then forming a porous reaction layer by controlling the overvoltage during the process of depositing platinum on the surface of the electrode layer using a constant voltage, thereby enabling a stable reaction based on the reaction sensitivity to glucose.
[0018] FIG. 1 is a flowchart illustrating a method for manufacturing a microneedle for biosensing according to one embodiment of the present invention.
[0019] Figure 2 is a state diagram showing the S100 step shown in Figure 1.
[0020] Figure 3 is an electron microscope magnified image of the needle section where the S200 step shown in Figure 1 is performed.
[0021] Figure 4 is an electron microscope magnified image of the needle section where the S300 step shown in Figure 1 is performed.
[0022] Figure 5 is a graph showing the reaction rate with a substance in the body according to the deposition time of the enzyme in the S300 step indicated in Figure 1.
[0023] FIG. 6 is a schematic cross-sectional view of a microneedle for biosensing according to one embodiment of the present invention.
[0024] FIG. 7 is an electron microscope magnified image of a microneedle for biosensing in one embodiment of the present invention.
[0025] FIG. 8 is an example of the application state of a microneedle for biosensing according to one embodiment of the present invention.
[0026] FIG. 9 is a flowchart illustrating a method for manufacturing microneedles for biosensing according to another embodiment of the present invention.
[0027] Figure 10 is a state diagram showing the S1000 step shown in Figure 9.
[0028] FIG. 11 is a schematic diagram of the manufacturing steps of a microneedle for biosensing according to another embodiment of the present invention.
[0029] Figure 12 is a graph showing the charge density state during deposition by a constant potential in the S3000 step described in Figure 9.
[0030] FIG. 13 is a schematic diagram showing the shape of the reaction layer deposited according to the overpotential mode of the electrode layer in step S3000 described in FIG. 9.
[0031] Figure 14 is an electron microscope magnified image of the shape of the reaction layer deposited according to the overpotential mode of the electrode layer in step S3000 described in Figure 9.
[0032] Figure 15 is a graph showing the reaction order of glucose in the reaction layer formed for each mode in step S3000 described in Figure 9.
[0033] Figure 16 is a graph showing the reaction sensitivity to glucose of the reaction layer formed for each mode in step S3000 described in Figure 9.
[0034] FIG. 17 is a schematic diagram showing a microneedle for biosensing according to another embodiment of the present invention.
[0035] Hereinafter, preferred embodiments according to the present invention will be described in detail with reference to the attached drawings. Prior to this, terms and words used in this specification and claims should not be interpreted as being limited to their ordinary or dictionary meanings, and should be interpreted in a meaning and concept consistent with the technical spirit of the present invention, based on the principle that the inventor can appropriately define the concept of the terms to best describe his invention.
[0036] Referring to FIG. 1, a method for manufacturing a microneedle for biosensing according to one embodiment of the present invention may include a needle portion forming step (S100), an electrode layer forming step (S200), and an enzyme layer forming step (S300).
[0037] Referring to FIG. 2, the needle portion forming step (S100) is a step of forming a needle portion (200) having a pointed end so as to be inserted into the epidermal layer of the skin on the base portion (100).
[0038] At this time, the base portion (100) is formed in the shape of a plate, and at least one needle portion (200) may be disposed on one surface of the base portion (100).
[0039] And, the needle part (200) can be formed on one side of the base part (100) through 3D printing.
[0040] In this way, the needle portion (200) is formed through 3D printing, so it can be easily and quickly formed on the base portion (100) to have a uniform size, shape, and pattern.
[0041] Here, the base portion (100) and the needle portion (200) may be made of a synthetic resin material. For example, the base portion (100) and the needle portion (200) may be made of polymethyl methacrylate (PMMA) or resin, but are not limited thereto. Also, the base portion (100) and the needle portion (2000) may be formed of the same material or different materials.
[0042] After the above needle portion formation step (S100) is completed, an electrode layer formation step (S200) is performed to form an electrode layer (300) on the surface of the needle portion (200).
[0043] The electrode layer formation step (S200) may include a first deposition step (S210) and a second deposition step (S220).
[0044] As shown in FIG. 3(a), the first deposition step (S210) forms a first conductive material layer (310) by depositing a first conductive material on the surface of the needle portion (200). Here, the first conductive material may be gold (Au).
[0045] As shown in FIG. 3(b), the second deposition step (S220) forms a second conductive material layer (320) by depositing a second conductive material on the surface of the first conductive material layer (310). Here, the second conductive material may be platinum (Pt).
[0046] As shown in FIG. 4, when an electrode layer (300) is formed on the surface of the needle portion (200), the first conductive material and the second conductive material are formed on the needle portion (200) through electrodeposition, so that the electrode layer (300) is evenly coated on the surface of the needle portion (200) and can have a porous structure.
[0047] At this time, the roughness factor of the electrode layer (300) formed on the surface of the needle portion (200) may have a roughness factor that allows for stable fixation of the enzyme and stable insertion of the needle portion (200) into the epidermal layer.
[0048] For example, the roughness coefficient of the electrode layer (300) may be 13.44 to 172. Here, if the roughness coefficient of the electrode layer (300) is less than 13.44, the amount of enzyme immobilized is small due to the low specific surface area of the electrode layer (300), making it difficult to measure through the enzyme. Also, if the roughness coefficient of the electrode layer (300) exceeds 172, it may detach during the process of insertion into the epidermal layer, and the insertion performance may be degraded due to an increase in the thickness of the electrode layer (300).
[0049] After the electrode layer (300) formation step (S200) is completed, an additional step of forming a passivation layer (330) on at least a portion of the surface of the electrode layer (300) may be performed.
[0050] After the electrode layer (300) formation step (S200) is completed, an enzyme layer formation step (S300) is performed to form an enzyme layer (400) by electrodepositing an enzyme mixed with an enzyme-immobilizing polymer on the surface of the electrode layer (300).
[0051] Here, the enzyme-fixed polymer serves to fix the enzyme to the electrode layer (300), and the enzyme serves to convert a substance in the body into a substance capable of generating an electrical signal.
[0052] For example, in the enzyme layer formation step (S300), at least one of the enzyme-immobilizing polymer constituting the enzyme layer (400) may be selected from O-phenylenediamine (OPD), Polypyrrole (PPy), Polyaniline (PANI), Polythiophene (PT), Poly(m-phenylenediamine)), Poly(p-phenylenediamine)), Poly(3,4-ethylenedioxythiophene) (PEDOT), and Poly(aniline-co-anthranilic acid) may be selected.
[0053] And, the enzyme may be selected from at least one of glucose oxidase, lactate oxidase, cholesterol oxidase, uricase, amino acid oxidase, acetylcholinesterase, alcohol oxidase, creatinine diaminohydrolase, ascorbate oxidase, and dopamine oxidase.
[0054] Here, the glucose oxidase reacts with glucose to enable measurement for diabetes management.
[0055] The above lactate oxidase reacts with lactic acid to enable measurement of metabolic status during exercise and diagnosis of shock and sepsis.
[0056] The above cholesterol oxidase reacts with cholesterol to enable measurement for cardiovascular disease risk assessment.
[0057] The above urase reacts with uric acid to enable measurements for the evaluation of gout and kidney function.
[0058] The above-mentioned amino acid oxidase reacts with amino acids to enable measurement for the diagnosis of metabolic disorders.
[0059] The above acetylcholinesterase reacts with acetylcholine to enable measurements for the evaluation of nervous system function.
[0060] The above alcohol oxidase reacts with alcohol to enable the measurement of blood alcohol concentration.
[0061] The above-mentioned creatinine diaminohydrolase reacts with creatinine to enable measurement for the evaluation of kidney function.
[0062] The above-mentioned ascorbate oxidase reacts with ascorbic acid to enable measurements for the evaluation of vitamin C deficiency and oxidative stress status.
[0063] The above-mentioned dopamine oxidase reacts with dopamine to enable measurements for the evaluation of neurological disorders and mental health.
[0064] In this enzyme layer formation step (S300), when the enzyme layer (400) is formed on the surface of the electrode layer (300) by electrodeposition, the degree to which the enzyme is immobilized through the enzyme-immobilizing polymer on the surface of the electrode layer (300) can be determined according to the electrodeposition time.
[0065] For example, in the enzyme layer formation step (S300) above, if the enzyme is an enzyme that measures blood sugar, namely glucose oxidase, then as shown in Fig. 5, when the enzyme is electrodeposited on the surface of the electrode layer (300) through an enzyme-immobilizing polymer for 5 to 20 minutes, it can be seen that the reaction with glucose is stably carried out and the measurement accuracy is increased.
[0066] In this way, in the enzyme layer formation step (S300), the enzyme mixed with the enzyme-immobilizing polymer is formed on the electrode layer (300) through electrodeposition, so that the enzyme layer (400) can be formed quickly and stably at an accurate location on the electrode layer (300).
[0067] After the enzyme layer formation step (S300) is completed, a protective layer coating step (S400) may be performed to coat a protective layer (not shown) on the surface of the enzyme layer (400).
[0068] The above protective layer is coated on the surface of the enzyme layer (400) to prevent the enzyme layer (400) from falling off the electrode layer (300).
[0069] For example, the above protective layer may be made of polyurethane, but is not limited thereto.
[0070] In this embodiment, the method for manufacturing a microneedle for biosensing involves depositing a conductive aqueous material on the surface of a needle portion (200) to form a porous electrode layer (300), and then electrodepositing an enzyme mixed with an enzyme-immobilizing polymer on the surface of the electrode layer (300) to form an enzyme layer (400). This allows the enzyme layer (400) to be formed uniformly and quickly on the surface of the needle portion (200), thereby increasing measurement accuracy through stable electrical signal reception via the enzyme and minimizing manufacturing costs.
[0071] FIG. 6 is a schematic cross-sectional view of a bio-sensing microneedle according to one embodiment of the present invention, and FIG. 7 is an electron microscope magnified image of a bio-sensing microneedle according to one embodiment of the present invention.
[0072] Referring to FIGS. 6 and 7, a biosensing microneedle of one embodiment may include a base portion (100), a needle portion (200), an electrode layer (300), and an enzyme layer (400).
[0073] The base portion (100) supports the needle portion (200) and prevents the needle portion (200) from being fully inserted into the user's epidermal layer.
[0074] Here, the base portion (100) is positioned outside the epidermal layer, allowing the needle portion (200) inserted into the epidermal layer to be stably removed from the epidermal layer.
[0075] For example, the base portion (100) may be formed in the shape of a plate.
[0076] The above needle portion (200) is a part that is inserted into the epidermal layer.
[0077] At least one of the above needle portions (200) may be disposed on one side of the base portion (100).
[0078] The above needle portion (200) may have a pointed cross-sectional shape so as to be easily inserted into the epidermal layer.
[0079] For example, the needle portion (200) may have a horn-shaped cross-section, and more preferably, may have a cross-sectional shape that decreases as it moves away from one side of the base portion (100).
[0080] Here, the needle portion (200) can be formed on one side of the base portion (100) through 3D printing.
[0081] In this way, the needle portion (200) is formed on the base portion (100) through 3D printing, so it can be easily and quickly formed on the base portion (100) to have a uniform size, shape, and pattern.
[0082] Here, the base portion (100) and the needle portion (200) may be made of a synthetic resin material. For example, the base portion (100) and the needle portion (200) may be made of polymethyl methacrylate (PMMA) or resin, but are not limited thereto. Also, the base portion (100) and the needle portion (2000) may be formed of the same material or different materials.
[0083] The electrode layer (300) may be provided on the surface of the needle portion (200). Here, the electrode layer (300) may be provided on at least a portion of the entire surface of the needle portion (200).
[0084] The electrode layer (300) may have a porous structure so that the enzyme layer (400), which will be described later, can be stably fixed.
[0085] Here, the electrode layer (300) can be formed by depositing a conductive material on the surface of the needle portion (200). For example, when the electrode layer (300) is formed on the surface of the needle portion (200) by electrodeposition, the electrode layer (300) can be evenly coated on the surface of the needle portion (200) and have a porous structure.
[0086] The electrode layer (300) may include a first conductive material layer (310) and a second conductive material layer (320).
[0087] The first conductive material layer (310) and the second conductive material layer (320) may be made of different conductive materials.
[0088] For example, the first conductive material layer (310) may be made of gold (Au), and the second conductive material layer (320) may be made of platinum (Pt).
[0089] The roughness factor of the electrode layer (300) may have a roughness factor that enables stable fixation of the enzyme layer (400) provided on the surface of the electrode layer (300) and stable insertion of the needle portion (200) into the epidermal layer.
[0090] For example, the roughness coefficient of the electrode layer (300) may be 13.44 to 172. Here, if the roughness coefficient of the electrode layer (300) is less than 13.44, the amount of the enzyme layer (400) fixed is small due to the low specific surface area of the electrode layer (300), making it difficult to measure through the enzyme. Also, if the roughness coefficient of the electrode layer (300) exceeds 172, it may detach during the process of insertion into the epidermal layer, and the insertion performance may be reduced due to an increase in the thickness of the electrode layer (300).
[0091] The enzyme layer (400) may be provided on the surface of the electrode layer (300). Here, the enzyme layer (400) may be provided on at least a portion of the entire surface of the electrode layer (300).
[0092] At this time, the enzyme layer (400) can be formed on the surface of the electrode layer (300) through electrodeposition.
[0093] And, the enzyme layer (400) may be composed of a mixture of enzyme-fixed polymer and enzyme.
[0094] Here, the enzyme-immobilized polymer serves to immobilize the enzyme on the electrode layer (300), and the enzyme reacts with a substance in the body to generate an electrical signal through reduction, thereby enabling measurement.
[0095] For example, the enzyme-immobilized polymer may be selected from at least one of O-phenylenediamine (OPD), Polypyrrole (PPy), Polyaniline (PANI), Polythiophene (PT), Poly(m-phenylenediamine)), Poly(p-phenylenediamine)), Poly(3,4-ethylenedioxythiophene) (PEDOT), and Poly(aniline-co-anthranilic acid).
[0096] The above enzyme may be selected from at least one of glucose oxidase, lactate oxidase, cholesterol oxidase, uricase, amino acid oxidase, acetylcholinesterase, alcohol oxidase, creatinine diaminohydrolase, ascorbate oxidase, and dopamine oxidase.
[0097] In addition, in the case where the micro-needle for biosensing of one embodiment is formed with a plurality of needle portions (200), an enzyme layer (400) having different enzymes may be formed on the surface of any one of the plurality of needle portions (200) and the needle portion (200) adjacent to said needle portion (200).
[0098] For example, as shown in FIG. 8, a plurality of needle sections (200) may be separated into several needle sections (200a, 200b, 200c, 200d, 200e, 200f), and then an enzyme layer (400) having different enzymes may be formed on the surface of the needle sections (200) placed in each needle section (200a, 200b, 200c, 200d, 200e, 200f). Through this, it may be possible to measure and evaluate various biomarkers by reacting with substances in the body in the enzyme layer (400) having different enzymes on the needle sections (200) placed in the needle sections (200a, 200b, 200c, 200d, 200e, 200f).
[0099] In addition, the biosensing microneedle of one embodiment may include a protective layer (not shown) provided on the surface of the enzyme layer (400).
[0100] This protective layer prevents the enzyme layer (400) from falling off the electrode layer (300).
[0101] For example, the above protective layer may be made of polyurethane, but is not limited thereto.
[0102] In this way, the microneedle for biosensing of one embodiment is provided with a porous electrode layer (300) on the surface of the needle portion (200) by depositing a conductive material, and an enzyme layer (400) on the surface of the electrode layer (300) by electrodeposition of an enzyme mixed with an enzyme-immobilizing polymer. By forming the enzyme layer (400) uniformly and quickly on the surface of the needle portion (200), measurement accuracy can be increased through stable electrical signal reception via the enzyme, and manufacturing costs can be minimized.
[0103] Referring to FIG. 9, a method for manufacturing a microneedle for biosensing according to another embodiment of the present invention may include a needle portion forming step (S1000), an electrode layer forming step (S2000), and a reaction layer forming step (S3000).
[0104] As shown in FIG. 10, the needle portion forming step (S1000) is a step of forming a needle portion (2000) having a pointed end on a base portion (1000) so as to be inserted into the epidermal layer of the skin.
[0105] At this time, the base portion (1000) is formed in the shape of a plate, and at least one needle portion (2000) may be disposed on one surface of the base portion (1000).
[0106] Referring to FIG. 11, the needle portion (2000) can be formed by the steps of forming a needle groove (2000b) on one surface of a mold plate (2000a) by laser processing, injecting a synthetic resin material such as resin into one surface of the mold plate (2000a) and curing it through UV, and removing the cured synthetic resin material from the mold plate (2000a).
[0107] In addition, in another embodiment, the needle portion (2000) may be formed on one surface of the base portion (1000) through 3D printing. At this time, the needle portion (2000) formed on the base portion (1000) through 3D printing may undergo an additional curing step.
[0108] In this way, when the needle part (2000) is formed by 3D printing in the needle part forming step (S1000), it can be easily and quickly formed on the base part (1000) to have a uniform size, shape, and pattern.
[0109] Here, when the needle part (2000) is formed by 3D printing in the needle part formation step (S1000), the base part (1000) and the needle part (2000) may be made of a synthetic resin material. For example, the base part (1000) and the needle part (2000) may be made of polymethyl methacrylate (PMMA) or resin, but are not limited thereto. Also, the base part (1000) and the needle part (2000) may be formed of the same material, but are not limited thereto and may be formed of different materials. Here, the resin may be selectively applied as either UV-curing resin or thermo-curing resin.
[0110] In addition, after the needle portion (2000) is formed on the base portion (1000) in the needle formation step (S1000), a polymer membrane solution can be spin-coated onto the surfaces of the base portion (1000) and the needle portion (2000) to form a membrane layer (2100). This membrane layer (2100) serves to ensure stable adhesion during the deposition of a conductive material to form an electrode layer (3000).
[0111] After the above needle portion formation step (S1000) is completed, an electrode layer formation step (S2000) is performed to form an electrode layer (3000) on the surface of the base portion (1000) and the needle portion (2000).
[0112] In this electrode layer formation step (S2000), a conductive material is deposited on the surface of the needle portion (2000) and the base portion (1000) to form an electrode layer (3000). Here, the conductive material may be gold (Au).
[0113] After the above electrode layer formation step (S2000) is completed, an additional step of forming a passivation layer (3200) on at least a portion of the surface of the electrode layer (3000) can be performed.
[0114] More specifically, the passivation layer (3200) may be formed on the surface of the electrode layer (3000) disposed in the needle portion (2000) and base portion (1000) regions, excluding a portion of the electrode layer (3000) formed on the upper side of the needle portion (2000).
[0115] After the electrode layer formation step (S2000) is completed, a reaction layer formation step (S3000) is performed to form a porous reaction layer (4000) by depositing platinum (Pt) on the surface of the electrode layer (3000).
[0116] Here, platinum (Pt) exhibits excellent catalytic activity toward glucose by adsorbing glucose and forming -OH groups, and thus plays a role in generating electrical signals by reacting with glucose in the body.
[0117] The above reaction layer (4000) can be formed by the deposition of platinum using a potentiostat.
[0118] At this time, when the reaction layer (4000) is formed by depositing platinum using a potentiostat, it is formed to have porosity, thereby increasing the reaction efficiency with glucose through an increase in surface area.
[0119] Here, the reaction layer (4000) can be set based on the charge density generated in the potentiostat during platinum deposition using the potentiostat, which is based on the electrochemical surface area (ECSA) of the electrode layer (3000).
[0120] At this time, the electrochemical surface area (ECSA) of the electrode layer (3000), more specifically the electrochemical surface area (ECSA) of a portion of the electrode layer (3000) above the needle portion (2000) where the reaction layer (4000) is deposited, can be set through the following mathematical formula 1.
[0121]
[0122] Here, is the reduction charge, is the surface coverage ratio, is 390μC / cm 2 am.
[0123] And, when forming the reaction layer (4000) by depositing platinum using the above potentiostat, the voltage applied to the potentiostat can be maintained in an overpotential state lower than the reference potential at which platinum is deposited.
[0124] With reference to FIG. 12, the current of the electrode layer is shown as a function of the applied voltage for platinum deposition by a constant potential, FIG. 13 is a schematic diagram showing the shape of the reaction layer deposited according to the overpotential mode of the electrode layer, and FIG. 14 is an electron microscope magnified image of the shape of the reaction layer deposited according to the overpotential mode of the electrode layer.
[0125] In this reaction layer formation step (S3000), when depositing platinum using a potentiostat, the deposition overpotential can be set to 'Nucleation', 'Growth', or 'Nucleation & Growth' modes.
[0126] For example, in the case of nucleation deposition, a high overpotential of -0.2V is applied, which can be driven by diffusion control kinetics at a high overpotential. At this time, the reaction layer (4000) on the surface of the electrode layer (3000) has a dendritic structure with significant roughness.
[0127] In contrast, in the growth deposition, a low overpotential of 0.15 V was applied, which, under the influence of surface control kinetics at the low overpotential, causes the reaction layer (4000) to have a structure of spherical particles with relatively minimal roughness through gradual platinum reduction.
[0128] Finally, an alternating potential sequence of nucleation and growth conditions can be applied in the nucleation and growth deposition. That is, when -0.2V for 20 seconds, 0.15V for 100 seconds, -0.2V for 100 seconds, and 0.15V for 20 seconds are applied, the shape of the resulting reaction layer (4000) includes dendritic and spherical particles. Thus, the nucleation and growth have intermediate surface roughness compared to that observed in the nucleation and growth processes.
[0129] In this reaction layer formation step (S3000), the amount of platinum deposited using a potentiostat is controlled by monitoring the charge during each deposition and controlling the deposition time.
[0130] Here, the amount of deposited platinum can be controlled by manipulating the charge density (CD), which is the total charge for platinum reduction normalized to the electrochemical surface area (ECSA) of the electrode layer (3000).
[0131] In this way, the deposition of platinum using a potentiostat allows for easy control of the shape of the reaction layer (4000) by controlling the applied overpotential, and the electrochemical surface area (ECSA) of the reaction layer (4000) can be increased by increasing the charge density during deposition.
[0132] Referring to FIG. 15, it can be seen that the reaction order of the reaction layer (4000) formed for each mode in the reaction layer formation step (S3000) for glucose was an average of 0.821 for 'nucleation', 0.854 for nucleation and growth, and gradually increased to 0.919 for growth. This indicates that the nonlinearity of the sensor response increases as the electrochemical surface area (ECSA) of the deposited platinum increases.
[0133] In addition, the average rate constants obtained during nucleation, nucleation and growth, and growth of the reaction layer (4000) are each 0.406×10 -18 , 0.187×10 -18 , 0.039×10 -18As such, it can be seen that the reaction rate increases as the electrochemical surface area (ECSA) of the deposited platinum increases.
[0134] Referring to FIG. 16, when the reaction layer (4000) is formed through nucleation, nucleation and growth, and growth, it can be seen that in the case of the reaction layer (4000) having a higher electrochemical surface area (ECSA), that is, the reaction layer (4000) deposited in the nucleation mode, the sensitivity is increased but shows a non-linear response to changes in glucose concentration. In the case of the reaction layer (4000) having a lower electrochemical surface area (ECSA), that is, the reaction layer (4000) deposited in the growth mode, the sensitivity is decreased but shows a more linear response within the concentration range.
[0135] In this way, after the reaction layer formation step (S3000) is completed, a protective layer coating step may be performed to coat a protective layer on the surface of the reaction layer (4000).
[0136] The above protective layer is coated on the surface of the reaction layer (4000) to prevent the reaction layer (4000) from falling off the electrode layer (3000).
[0137] For example, the above protective layer may be made of polyurethane, but is not limited thereto.
[0138] In this other embodiment of the method for manufacturing a microneedle for biosensing, a conductive material made of gold is deposited on the surface of the needle portion (2000) to form an electrode layer (3000), and then platinum is deposited on the surface of the electrode layer (3000) in an overpotential state by a certain potential to form a porous reaction layer (4000), thereby enabling a stable reaction rate and reaction sensitivity with glucose.
[0139] FIG. 17 is a schematic cross-sectional view of a microneedle for biosensing according to another embodiment of the present invention.
[0140] Referring to FIG. 17, a biosensing microneedle of another embodiment may include a base portion (1000), a needle portion (2000), an electrode layer (3000), and a reaction layer (4000).
[0141] The base portion (1000) supports the needle portion (2000) and prevents the needle portion (2000) from being fully inserted into the user's epidermal layer.
[0142] Here, the base portion (1000) is positioned outside the epidermal layer, allowing the needle portion (2000) inserted into the epidermal layer to be stably removed from the epidermal layer.
[0143] For example, the base portion (1000) may be formed in the shape of a plate.
[0144] The above needle portion (2000) is a part that is inserted into the epidermal layer.
[0145] At least one of the above needle portions (2000) may be disposed on one side of the base portion (1000).
[0146] The above needle portion (2000) may have a pointed cross-sectional shape so as to be easily inserted into the epidermal layer.
[0147] For example, the needle portion (2000) may have a horn-shaped cross-section, and more preferably, may have a cross-sectional shape that decreases as it moves away from one side of the base portion (1000).
[0148] In this way, the needle portion (2000) may be formed on the base portion (1000) through 3D printing. When the needle portion (2000) is formed through 3D printing, it can be easily and quickly formed on the base portion (1000) to have a uniform size, shape, and pattern.
[0149] Here, the base portion (1000) and the needle portion (2000) may be made of a synthetic resin material. For example, the base portion (1000) and the needle portion (2000) may be made of polymethyl methacrylate (PMMA) or resin, but are not limited thereto. Also, the base portion (1000) and the needle portion (2000) may be formed of the same material, but are not limited thereto and may be formed of different materials. Here, the resin may be selectively applied as either UV-curing resin or thermo-curing resin.
[0150] In addition, a membrane layer (2100) may be formed on the surface of the base portion (1000) and the needle portion (2000). This membrane layer (2100) serves to ensure stable adhesion during the deposition of a conductive material to form an electrode layer (3000) to be described later.
[0151] The electrode layer (3000) can be provided on the surface of the base portion (1000) and the surface of the needle portion (2000).
[0152] The electrode layer (3000) can be formed by depositing a conductive material on the surface of the base portion (1000) and the surface of the needle portion (2000).
[0153] For example, the conductive material forming the electrode layer (3000) may be gold (Au).
[0154] The above reaction layer (4000) can be formed on the surface of the electrode layer (3000).
[0155] The reaction layer (4000) may be formed on a portion of the surface of the electrode layer (3000). For example, the reaction layer (4000) may be provided in a portion of the electrode layer (3000) formed on the upper side of the needle portion (2000).
[0156] Here, the reaction layer (4000) may be made of platinum (Pt), which has excellent reactivity with glucose.
[0157] At this time, the reaction layer (4000) can be formed on the surface of the electrode layer (3000) by deposition.
[0158] And, the reaction layer (4000) can be formed on the surface of the electrode layer (3000) by deposition through a charge density set at a constant potential based on the electrochemical surface area (ECSA) of the electrode layer (3000).
[0159] In this way, the reaction layer (4000) has porosity when formed on the surface of the electrode layer (3000) by deposition with a charge density set at a certain voltage, thereby increasing the reaction efficiency with glucose through an increase in surface area.
[0160] For example, when the reaction layer (4000) is formed by the deposition of platinum using a constant potential, it can be formed in a state where the constant potential is set to have an overpotential charge density.
[0161] In addition, for a biosensing microneedle of another embodiment, a passivation layer (320) may be formed on the surface of an electrode layer (3000) where a reaction layer (4000) is not formed.
[0162] In addition, an exchange membrane (5000) may be formed on the surface of the electrode layer (3000) and the reaction layer (4000).
[0163] Here, the exchange membrane (5000) may be made of a copolymer that enables the movement of hydrogen ions. As an example, the exchange membrane (5000) may be Nafion or a mixture of Nafion.
[0164] For example, the above protective layer may be made of polyurethane, but is not limited thereto.
[0165] In this other embodiment, the microneedle for biosensing has an electrode layer (3000) formed by depositing a conductive material made of gold on the surface of the needle portion (2000), and a porous reaction layer (4000) formed by depositing platinum in an overpotential state by a certain potential on the surface of the electrode layer (3000), thereby enabling a stable reaction based on the reaction sensitivity to glucose.
[0166] The present invention has been described with reference to embodiments illustrated in the drawings, but this is merely illustrative, and those skilled in the art will understand that various modifications and equivalent alternative embodiments are possible therefrom. Accordingly, the true technical scope of protection of the present invention should be determined by the technical spirit of the appended claims.
Claims
1. A needle portion forming step of forming a needle portion on a base portion; and An electrode layer formation step of forming a porous electrode layer by depositing a conductive material on the surface of the above needle portion; A method for manufacturing a microneedle for biosensing, comprising: an enzyme layer formation step of electrodepositing an enzyme mixed with an enzyme-immobilizing polymer onto the surface of the electrode layer to form an enzyme layer.
2. In Claim 1, The above needle portion forming step is a method for manufacturing a microneedle for biosensing in which the needle portion is formed by 3D printing.
3. In Claim 1, The above electrode layer formation step is A first deposition step of forming a first conductive material layer by depositing a first conductive material on the surface of the needle portion, and It includes a second deposition step of forming a second conductive material layer by depositing a second conductive material on the surface of the first conductive material layer, and A method for manufacturing a bio-sensing microneedle in which the first conductive material is gold (Au) and the second conductive material is platinum (Pt).
4. In Claim 1, A method for manufacturing a microneedle for biosensing in which the roughness coefficient of the electrode layer formed in the above electrode layer formation step is 13.44 to 172.
5. In Claim 1, A method for manufacturing a microneedle for biosensing in which the enzyme-immobilized polymer comprises at least one of O-phenylenediamine (OPD), Polypyrrole (PPy), Polyaniline (PANI), Polythiophene (PT), Poly(m-phenylenediamine)), Poly(p-phenylenediamine)), Poly(3,4-ethylenedioxythiophene) (PEDOT), and Poly(aniline-co-anthranilic acid).
6. In Claim 1, A method for manufacturing a biosensing microneedle comprising at least one of the following enzymes: glucose oxidase, lactate oxidase, cholesterol oxidase, uricase, amino acid oxidase, acetylcholinesterase, alcohol oxidase, creatinine diaminohydrolase, ascorbate oxidase, and dopamine oxidase.
7. In Claim 1, A method for manufacturing a biosensing microneedle in which electrodeposition is performed for 5 to 20 minutes when the enzyme is a blood glucose measuring enzyme in the enzyme layer formation step.
8. The bass part; At least one needle portion disposed on one surface of the base portion; A porous electrode layer provided on the surface of the above needle portion; Includes an enzyme layer provided on the surface of the electrode layer; The above electrode layer is formed by the deposition of a conductive material on the surface of the needle portion, and The above enzyme layer is a biosensing microneedle formed by electrodeposition of an enzyme mixed with an enzyme-immobilizing polymer on the surface of the above electrode layer.
9. In Claim 8, The above needle portion is a bio-sensing microneedle formed by 3D printing.
10. In Claim 8, The above electrode layer comprises a first conductive material layer disposed on the surface of the needle portion, and It includes a second conductive material layer disposed on the surface of the first conductive material, A biosensing microneedle in which the first conductive material layer is made of gold (Au) and the second conductive material layer is made of platinum (Pt).
11. In Claim 8, A microneedle for biosensing in which the roughness coefficient of the electrode layer is 13.44 to 172.
12. In claim 8, The above enzyme-immobilized polymer is a biosensing microneedle composed of at least one of O-phenylenediamine (OPD), polypyrrole (PPy), polyaniline (PANI), polythiophene (PT), poly(m-phenylenediamine)), poly(p-phenylenediamine)), poly(3,4-ethylenedioxythiophene) (PEDOT), and poly(aniline-co-anthranilic acid).
13. In Claim 8, The above enzyme is a biosensing microneedle composed of at least one of glucose oxidase, lactate oxidase, cholesterol oxidase, urase, amino acid oxidase, acetylcholinesterase, alcohol oxidase, creatinine diaminohydrolase, ascorbate oxidase, and dopamine oxidase.
14. In Claim 8, The above needle portion is formed in multiple numbers, and A biosensing microneedle having a plurality of needle portions and an enzyme layer composed of different enzymes formed on the surface of a needle portion adjacent to said needle portion.
15. A needle portion forming step (S1000) for forming a needle portion on a base portion; An electrode layer forming step (S2000) of forming an electrode layer of a conductive material on the surface of the base portion and the needle portion; A reaction layer forming step (S3000) in which platinum is deposited on the surface of the electrode layer to form a porous reaction layer; is included, In the reaction layer formation step above, the deposition of platinum is carried out by a potentiostat, and A method for manufacturing microneedles for biosensing in which the charge density generated at the constant potential during the deposition of the above platinum is set based on the electrochemical surface area (ECSA) of the electrode layer.
16. In Claim 15, The above needle part forming step (S1000) is A step of forming needle grooves on one surface of a mold plate by laser processing, and A method for manufacturing a bio-sensing microneedle comprising the step of injecting a synthetic resin material into one surface of the above-mentioned mold plate and then curing it.
17. In Claim 15, The above needle part forming step (S1000) is a method for manufacturing a microneedle for biosensing by 3D printing.
18. In Claim 15, A method for manufacturing a bio-sensing microneedle in which the charge density generated at the constant potential during the reaction layer formation step (S3000) is maintained in an overpotential state.
19. In Claim 15, A method for manufacturing bio-sensing microneedles in which the conductive material is gold (Au).
20. In Claim 15, The electrochemical surface area (ECSA) of the above electrode layer is A method for manufacturing microneedles for biosensing configured through. (Here, is the reduction charge, is the surface coverage ratio, is 390μC / cm 2 It could be.) 21. The bass part; At least one needle portion disposed on one surface of the base portion; An electrode layer made of a conductive material, provided on the surface of the base portion and the needle portion; A porous reaction layer provided on a portion of the surface of the electrode layer; comprising The above reaction layer is made of platinum, and The above reaction layer is a biosensing microneedle formed on the surface of an electrode layer by deposition with a charge density set based on the electrochemical surface area (ECSA) of the electrode layer.
22. In Claim 21, The above base portion and needle portion are microneedles for biosensing formed by 3D printing.
23. In Claim 21, The above base portion and needle portion are microneedles for biosensing formed by curing of a synthetic resin material.
24. In Claim 21, The above conductive material is a microneedle for biosensing, which is gold (Au).
25. In Claim 21, The above charge density is a biosensing microneedle maintained in an overpotential state.