Microneedle electrode for continuously and sensitively monitoring physiological information as well as preparation method and application of microneedle electrode

Abstract

The invention discloses a microneedle electrode for continuously and sensitively monitoring physiological information as well as a preparation method and application of the microneedle electrode, and belongs to the technical field of electrochemistry. The microneedle electrode provided by the invention comprises a substrate and a liquid-like film layer covalently coupled to the surface of the substrate, the preparation raw materials of the liquid-like film layer comprise a low-surface-energy polymer; the surface energy of the low-surface-energy polymer is less than or equal to 78 mJ / m < 2 >. The liquid-like film layer with the low-surface-energy polymer is constructed on the surface of the substrate, and due to the flexibility, smoothness and fluidity of the low-surface-energy polymer, slippage of various liquids is facilitated, so that the interaction between the microneedle electrode and various biological substances can be reduced, adhesion of protein, bacteria, cells and the like is remarkably inhibited, and the microneedle electrode has a good application prospect. The microneedle electrode with good anti-adhesion performance has a good application prospect in blood glucose detection.

Description

Technical Field

[0001] The present invention belongs to the technical field of electrochemistry, and particularly relates to a microneedle electrode for continuously and sensitively monitoring physiological information, a preparation method thereof, and an application thereof. Background Art

[0002] Diabetes, as a major clinical chronic disease, the current routine clinical blood glucose monitoring methods mainly detect by collecting blood from fingertips or veins and then placing the blood on a blood glucose meter or test strips. This requires frequent operations 2 - 6 times a day. Frequent, invasive, and painful blood collection for testing not only brings inconvenience to the lives of patients, but also increases the risk of wounds and reduces the compliance of testing. In addition, for critically ill diabetic patients, it is also necessary to observe the dynamic changes in blood glucose, which requires continuous and real-time monitoring of blood glucose. Therefore, the full-chain development of continuous and sensitive monitoring sensing devices and integrated device systems, as well as the research to break through the bottleneck of effective monitoring of chronic diseases such as diabetes, is of great significance.

[0003] In recent years, microneedle diagnosis and treatment technology (MN), as a cutting-edge painless transdermal technology, has been booming. The length of microneedles is usually 500 - 800 microns, just piercing through the stratum corneum of the skin to contact the dermis layer, so that it will not touch the blood vessels and nerves deep in the dermis layer and effectively avoid pain. Moreover, when microneedles are integrated with sensors, they can effectively detect subcutaneous (interstitial fluid ISF, mainly present in the dermis layer) biological information (the ISF glucose level is highly correlated with blood in the stable stage). However, there is still a major key bottleneck in current microneedle-based electrochemical sensing. In complex biological fluids (such as whole plasma or blood, etc.), especially in the human body environment, the surface of the sensing electrode usually causes non-specific bioadhesion due to a large number of coexisting background species (such as proteins, cells, etc.), thus hindering the effective coupling of the target biomarker with the electrode surface, leading to sensing failure and unable to achieve continuous and sensitive monitoring. This has become a major challenge for the rapid development and application of biosensors.

[0004] Research shows that the interaction between biological substances and the substrate is an important reason for their adhesion. For example, for protein adhesion, it mainly includes the following four types of interactions: 1) ionic or electrostatic interactions, 2) hydrogen bonds, 3) hydrophobic interactions (mainly entropy-driven), 4) charge transfer or particle electron donor / acceptor type interactions. And the adhesion of biological substances such as cells and bacteria is mainly mediated by protein ligands integrin or adhesin (composed of protein and polysaccharide molecules) embedded in the membrane surface. Therefore, reducing the interaction between the surface and biological substances such as proteins, and inhibiting non-specific adhesion and contamination, have become important strategies for real-time and stable monitoring of electrode sensors.

[0005] Current biosensing devices usually suffer from non-specific adhesion of biomaterials such as protein biomolecules, cells, and bacteria in the body, which hinders the transfer of electrons to the underlying electrode, leading to sensor failure and even tissue fibrosis. Technologies for the sensor to achieve continuous and stable monitoring include physical porous or membrane filtration methods, bioengineering methods, chemical coating modification, etc. Compared with physical and bioengineering technologies, based on chemical modification methods, only by grafting or coating a layer of polymer to construct an anti-adhesion surface, it has the advantages of high universality, convenient operation, simplicity and high efficiency, and has become one of the most common strategies for designing sensors for continuous and stable monitoring. However, there are the following problems in the current development of chemical modification to inhibit anti-adhesion and achieve continuous and sensitive monitoring of sensors: First, most current sensors are mainly limited to modification based on polyethylene glycol (PEG) or zwitterions and their derivatives, and their anti-adhesion performance usually depends on the extrusion - steric effect, that is, when biomaterials such as proteins approach the modified surface, the flexible molecular chains are severely extruded to generate a large steric effect, thus forming physical and energy barriers to adhesion. This property requires a long chain and a high surface grafting / coating density to form a large steric effect when biomaterials such as proteins adhere to the surface to effectively inhibit adhesion, but this also causes serious hindrance to the transfer of the measured signal to the underlying electrode, thus severely affecting the sensitivity. In addition, PEG, etc. may suffer oxidative damage in the presence of oxygen and transition metals and may cause adverse immune reactions in the body, thus affecting their continuous action in high-level blood and serum. And zwitterionic polymers not only require complex synthesis and purification of high-charge molecules, but also need to consider the influence of external pH and electric fields on the interfacial charge state, and polypeptides or natural zwitterionic polymers may be prone to hydrolysis (degradation). Second, for transient dissolution coatings, special environments are often required, and the universality is severely insufficient. Third, for lubricant perfusion surfaces, in a complex fluid environment, the lubricant is likely to be lost, resulting in the failure of the anti-pollution function. And when using lubricant to coat the electrode, its insulating oil layer will prevent the target biomarker from coupling with the electrode surface, resulting in the loss of the sensing function. Summary of the Invention

[0006] In order to overcome at least one of the above-mentioned existing technical problems, one of the purposes of the present invention is to provide a microneedle electrode.

[0007] Another purpose of the present invention is to provide a preparation method of the above-mentioned microneedle electrode.

[0008] Another purpose of the present invention is to provide a microneedle sensor.

[0009] Another purpose of the present invention is to provide a blood glucose monitor.

[0010] In order to achieve the above purposes, the technical solutions adopted by the present invention are as follows:

[0011] The first aspect of the present invention provides a microneedle electrode, comprising a substrate and a liquid-like film layer covalently coupled to the surface of the substrate; the raw materials for preparing the liquid-like film layer include a low surface energy polymer; the surface energy of the low surface energy polymer is ≤78 mJ / m 2 .

[0012] Based on the lubrication strategy surface modification, the present invention constructs a fully hydrophobic and liquid-flowing lubricating surface. Specifically, a layer of low surface energy polymer is covalently coupled to the surface of the substrate, thereby realizing the "liquid-like" molecular modified lubricating surface, so that the surface topography of the substrate only changes at the nanoscale. And due to the flexibility, smoothness and fluidity of the low surface energy polymer, the modified surface rarely undergoes ion / electrostatic, hydrogen bond and donor / acceptor interactions with proteins; and, due to the ability to repel most oil-like and hydrophobic solvents and molecules, the hydrophobic interaction between the modified surface and protein molecules is also very small. Therefore, benefiting from the reduction of various interaction forces, biomaterials such as proteins are easily sheared off by the dynamic flow of the liquid or diffused into the medium in both dynamic and static environments, which is beneficial to the slippage of the microneedle electrode in various liquids, thus significantly inhibiting the adhesion of biomaterials such as proteins, bacteria and cells.

[0013] In some embodiments of the present invention, the surface energy of the low surface energy polymer is 1-78 mJ / m 2 ; for example, it is any value among 10, 20, 30, 40, 50, 60, 70, 78 mJ / m 2 or the range value between any two of them, such as 10-40 mJ / m 2 .

[0014] Preferably, the low surface energy polymer comprises at least one of polydimethylsiloxane (PDMS), polytetrafluoroethylene (PTFE), polyvinylidene fluoride (PVDF), polystyrene (PS) or polytrifluoroethylene (PF3E); more preferably, the low surface energy polymer comprises polydimethylsiloxane (PDMS), polytetrafluoroethylene (PTFE) or a combination thereof; even more preferably, the low surface energy polymer is selected from polydimethylsiloxane (PDMS).

[0015] The surface energy of polydimethylsiloxane (PDMS), polytetrafluoroethylene (PTFE), polyvinylidene fluoride (PVDF), polystyrene (PS), polytrifluoroethylene (PF3E) is between 18-35 mJ / m 2 , and can achieve a good anti-adhesion effect when used in the present invention.

[0016] Preferably, the thickness of the liquid-like film layer is 1-10 nm; more preferably 1.5-5 nm; even more preferably 1.8-2.5 nm.

[0017] Preferably, the substrate has a planar microneedle array; the planar microneedle array has 3 to 7 microneedles; the length of each microneedle is 600 to 1000 μm; more preferably, the length of each microneedle is 700 to 900 μm; even more preferably, the length of each microneedle is 750 to 850 μm.

[0018] By using a specific microneedle length, the microneedle electrode can just pierce through the stratum corneum of the skin to contact the dermis layer without touching the blood vessels and nerves deep in the dermis layer, thus effectively avoiding pain.

[0019] Preferably, a catalytic substance is further provided between the substrate and the liquid-like film layer; the catalytic substance includes glucose oxidase and platinum nanoparticles.

[0020] Glucose oxidase can catalyze the oxidation reaction of glucose to generate an electrical signal, thereby enabling the detection of blood glucose levels; platinum nanoparticles can improve the catalytic efficiency of glucose oxidase.

[0021] The second aspect of the present invention provides a method for preparing the microneedle electrode according to the first aspect of the present invention, comprising the following steps: preparing a liquid-like film layer on the surface of the substrate to obtain the microneedle electrode;

[0022] Alternatively, a catalytic substance is provided on the surface of the substrate, and then a liquid-like film layer is prepared to obtain the microneedle electrode.

[0023] Preferably, before preparing the liquid-like film layer, it further includes the step of hydroxyl modification of the surface of the substrate.

[0024] Hydroxyl modification is performed on the surface of the substrate, so that the modified substrate can be coupled with the low surface energy polymer, thereby preparing a liquid-like film layer.

[0025] In some specific embodiments of the present invention, the surface of the substrate is modified with a mercaptoethanol solution.

[0026] Preferably, the method for preparing the liquid-like film layer is selected from atomic layer deposition (ALD).

[0027] Preferably, the deposition time of the atomic layer deposition method is 0.5 min to 5 h.

[0028] Preferably, the preparation of the liquid-like film layer is carried out in an atmosphere of an acidic gas and a protective gas; more preferably, the acidic gas includes hydrochloric acid gas; the protective gas includes nitrogen gas.

[0029] Preparing the liquid-like film layer in an acidic gas atmosphere is also beneficial for the hydrolysis-condensation reaction between the low surface polymer and the hydroxyl groups on the surface of the substrate; in some embodiments of the present invention, the acidic gas is also beneficial for ensuring the activity of glucose oxidase. The protective gas helps to stabilize and clean the chamber.

[0030] By controlling the deposition time, liquid film-like layers with different thicknesses can be obtained, thereby achieving the modification of the liquid film-like layer with precise thickness.

[0031] The third aspect of the present invention provides a microneedle sensor, including the microneedle electrode described in the first aspect of the present invention.

[0032] The fourth aspect of the present invention provides a blood glucose monitor, including the microneedle electrode described in the first aspect of the present invention, or the microneedle sensor described in the third aspect of the present invention.

[0033] Preferably, the blood glucose monitor further includes at least one of a wireless data transmission module, an artificial intelligence analysis algorithm module, or a user interaction module.

[0034] By adding a wireless data transmission module, the sensor can wirelessly send monitoring data to nearby receiving devices, such as smartphones or medical monitoring stations, thereby achieving remote monitoring; by adding an artificial intelligence analysis algorithm module, using artificial intelligence analysis algorithms to perform deep learning and pattern recognition on the collected physiological data, more accurate analysis results and health suggestions can be provided; by adding a user interaction module, users can easily view monitoring data, health reports, and alerts.

[0035] The beneficial effects of the present invention are: The present invention constructs a liquid film-like layer with a low surface energy polymer on the substrate surface. Due to the flexibility, smoothness, and fluidity of the low surface energy polymer, it is beneficial for the slippage of various liquids, thereby being able to reduce the interaction between the microneedle electrode and various biological substances, significantly inhibiting the adhesion of proteins, bacteria, cells, etc., and obtaining a microneedle electrode with good anti-adhesion performance, which has good application prospects in blood glucose detection. Description of the Drawings

[0036] Figure 1 It is a physical picture of the microneedle electrode in Example 1 and its SEM pictures during the modification process.

[0037] Figure 2 It is a graph of the film layer thickness change obtained by using different hydrolysis-condensation reaction times in Examples 1 to 8.

[0038] Figure 3 It is an optical microscope image of the dynamic slippage behavior of water on the electrode surface of Example 4 and the unmodified electrode surface.

[0039] Figure 4 It is the sliding angle of various liquids on the surface of the modified electrode with different film layer thicknesses.

[0040] Figure 5 It is an electrochemical performance graph of the microneedle electrode.

[0041] Figure 6 To show the fluorescence image of the adhesion of fibrinogen on the surface of the modified electrode with different thicknesses of quasi-liquid film layers.

[0042] Figure 7 To show the statistical analysis chart of the adhesion of fibrinogen on the surface of the modified electrode with different thicknesses of quasi-liquid film layers.

[0043] Figure 8 To show the statistical analysis chart of the adhesion of NIH-3T3 cells on the surface of the modified electrode with different thicknesses of quasi-liquid film layers.

[0044] Figure 9 To show the statistical analysis chart of the effect of the modified electrode on cell viability.

[0045] Figure 10 To show the fluorescence image of the adhesion of NIH-3T3 cells on the surface of the modified electrode with different thicknesses of quasi-liquid film layers and the fluorescence image of the effect of the modified electrode on cell viability.

[0046] Figure 11 The schematic diagram of the present invention for detecting blood glucose. Detailed implementation manners

[0047] The content of the present invention will be further described in detail through specific embodiments below. It should also be understood that the following embodiments are only used to further illustrate the present invention and cannot be construed as limiting the protection scope of the present invention. Some non-essential improvements and adjustments made by those skilled in the art based on the principles described in the present invention all fall within the protection scope of the present invention. The specific process parameters and the like in the following examples are also only an example within a suitable range, that is, those skilled in the art can make selections within a suitable range through the description in this article, rather than being limited to the specific data in the following examples. The raw materials, reagents or devices used in the following examples and comparative examples can be obtained from conventional commercial channels or can be obtained by existing known methods without special instructions.

[0048] In some embodiments of the present invention, the substrate has a planar microneedle array; each array has 5 microneedles, and the dimensions of the microneedles are: length 800 μm, width 200 μm, thickness 200 μm, and spacing 800 μm.

[0049] In some embodiments of the present invention, the tip of each microneedle is thinned by laser to form a sharp cone, and the sharp cone is designed into a micro-arrow structure. The tip of each micro-arrow is processed into a pinch angle of about 30°. There are micro-arrow structures on both sides of the tip. Each barbed angle has a sharp arrow, and the exposed edge is about 20 μm. This design can mechanically interlock with the tissue by the arrow tip, so that the microneedle array can better adhere to the skin surface after penetrating the skin, improving biocompatibility.

[0050] In some embodiments of the present invention, a wireless data transmission module can be added to enable the sensor to wirelessly send the monitored data to nearby receiving devices, such as smartphones or medical monitoring stations, thus realizing remote monitoring. Develop an artificial intelligence analysis algorithm to perform deep learning and pattern recognition on the collected physiological data to provide more accurate analysis results and health advice. Design a user-friendly interface and interaction system that allows users to easily view the monitored data, health reports, and alerts.

[0051] Example 1

[0052] A microneedle electrode is prepared as follows:

[0053] 1) Using a SUS304 stainless steel sheet as the substrate, the stainless steel substrate is etched by laser micro-etching technology to obtain a planar microneedle array. It can be seen that there are 5 microneedles in each array. The dimensions of the microneedles are: length 800 μm, width 200 μm, thickness 200 μm, spacing 800 μm, and the tip part is thinned by laser to form a sharp cone. The sharp cone is designed into a micro-arrow structure, and the tip of each micro-arrow is processed into a pinch angle of about 30°. There are micro-arrow structures on both sides of the tip, and each barb angle has a sharp arrow, and the exposed edge is about 20 μm. Then, all surfaces of the microneedle array are cleaned by plasma, and then a gold layer is electroplated using electrochemical deposition technology. Subsequently, the microneedle electrode is placed in a mercaptoethanol solution to introduce hydroxyl groups on the electrode surface, so that the electrode can be coupled with liquid-like molecules (PDMS) for functionalization. In a platinum sulfite solution, platinum nanoparticles are electroplated using electrochemical deposition technology to improve the sensing sensitivity, and then washed with deionized water and dried. Subsequently, BSA (80 mg / mL) and glutaraldehyde (2.5%) are used as the viscous solution and mixed with glucose oxidase (GOx, 50 mg / mL) in a volume ratio of 5:2:1, and mixed well on a shaker. Then, the mixture is uniformly coated on the electrode surface by spraying technology and dried overnight. Ag is magnetron sputtered on the electrode surface using a mask, and then the microneedle is placed in a 1 M concentration of KCl / HCl buffer solution, and a current of 1 μA is applied for 1 min to prepare an Ag / AgCl reference electrode. The above hydroxylated gold electrode is directly used as the counter electrode.

[0054] 2) Place the above microneedle array electrode in an ALD reactor, evacuate the reactor to a vacuum state, and use an inert gas nitrogen (N2) to stabilize and clean the chamber. Then, in a chamber temperature environment of 37 °C, weakly acidic water vapor (H2O / HCl), N2, and PDMS (Mw are 550, 2000 - 6000, 350000 in sequence; introduced into the chamber at a rate of 50 ± 10 Pa dose) are sequentially introduced. A hydrolysis-condensation reaction occurs on the electrode surface, and the reaction time is controlled to be 1 min, so as to realize the modification of a liquid-like film layer with an accurate thickness on the surface of the microneedle electrode.

[0055] Example 2

[0056] A microneedle electrode, the difference in the preparation steps from Example 1 lies in that the reaction time of the hydrolysis-condensation reaction on the electrode surface is controlled to be 2 min. Other raw materials and steps are the same as those in Example 1.

[0057] Example 3

[0058] A microneedle electrode, the difference in the preparation steps from Example 1 lies in that the reaction time of the hydrolysis-condensation reaction on the electrode surface is controlled to be 5 min. Other raw materials and steps are the same as those in Example 1.

[0059] Example 4

[0060] A microneedle electrode, the difference in the preparation steps from Example 1 lies in that the reaction time of the hydrolysis-condensation reaction on the electrode surface is controlled to be 10 min. Other raw materials and steps are the same as those in Example 1.

[0061] Example 5

[0062] A microneedle electrode, the difference in the preparation steps from Example 1 lies in that the reaction time of the hydrolysis-condensation reaction on the electrode surface is controlled to be 20 min. Other raw materials and steps are the same as those in Example 1.

[0063] Example 6

[0064] A microneedle electrode, the difference in the preparation steps from Example 1 lies in that the reaction time of the hydrolysis-condensation reaction on the electrode surface is controlled to be 1 h. Other raw materials and steps are the same as those in Example 1.

[0065] Example 7

[0066] A microneedle electrode, the difference in the preparation steps from Example 1 lies in that the reaction time of the hydrolysis-condensation reaction on the electrode surface is controlled to be 2 h. Other raw materials and steps are the same as those in Example 1.

[0067] Example 8

[0068] A microneedle electrode, the difference in the preparation steps from Example 1 lies in that the reaction time of the hydrolysis-condensation reaction on the electrode surface is controlled to be 4 h. Other raw materials and steps are the same as those in Example 1.

[0069] Comparative Example 1

[0070] A microneedle electrode, the difference in the preparation steps from Example 1 is that in step 2), the microneedle array electrode is placed in an ALD reactor, and the reactor is evacuated to a vacuum state to remove impurities that may affect the deposition process; without using H2O / HCl and N2, PDMS is directly introduced into the chamber by evaporation at a rate of 50±10 Pa dose, and the reaction time (1 min, 2 min, 5 min, 10 min, 20 min, 1 h, 2 h, 4 h) is controlled. Step 1) is the same as in Example 1. The microneedle electrode in this example is obtained.

[0071] Performance test

[0072] (1) XPS, ellipsometer, SEM, AFM, etc. are respectively used to characterize the film thickness, coverage rate, etc. of the electrode surface.

[0073] Figure 1 Fig. is the physical diagram of the microneedle electrode in Example 1 and its SEM diagrams during the modification process. Among them, (a) is the physical diagram of the microneedle electrode, (b) is the SEM diagram of the microneedle electrode, (c) is the SEM diagram of the material after electroplating a gold layer on the substrate, (d) is the SEM diagram of the material after electroplating platinum nanoparticles and spraying glucose oxidase, and (e) is the SEM diagram of the material after atomic deposition of PDMS. As can be seen from Figure 1 It can be seen that the physical diagram intuitively shows the physical morphology and structure of the microneedle electrode, including the layout and size of the microneedle array on the substrate; while the SEM diagram provides the microscopic details of the surface of the microneedle electrode, including the shape of the microneedles, surface texture, and the uniformity and coverage of the liquid-like film layer. Through the SEM images, the distribution of the liquid-like film layer on the surface of the microneedles and the changes in the surface morphology after modification can be observed, which is crucial for evaluating the quality and performance of the liquid-like film layer.

[0074] Figure 2 Fig. is the graph of the film thickness change obtained by using different hydrolysis-condensation reaction times in Examples 1 to 8. As can be seen from the figure, the change in the hydrolysis-condensation reaction time is crucial for controlling the precise thickness of the liquid-like film layer, because the film thickness directly affects the anti-adhesion property and electrochemical performance of the microneedle electrode. Multiple data points are shown in the figure, and each point represents the film thickness measured at a specific reaction time, so that the change in the film thickness with the increase of the reaction time can be observed; the blue curve array in the figure represents the film layer, showing the change of the film thickness with the reaction time.

[0075] Σ is used to represent the number of free non-overlapping polymer chains filling a fixed area. When 1 < Σ < 5, the transition from the "mushroom" state to the "brush" state occurs, the grafted chains are moderately stacked, the substrate surface is evenly covered and is more likely to freely stretch and rotate, thus significantly increasing flexibility and fluidity. Moreover, when the grafting thickness is about 2 nm, Σ ≈ 1.8, which promotes the slippage of various liquids.

[0076] (2) Anti - adhesion performance test: Select fluorescent fibrinogen (Fg) (both at a concentration of 0.1 mg / mL) as the representative protein, and co - incubate it with the microneedle electrode sheets modified with different film layer thicknesses at 37 °C for 24 h (short - term) and 7 d (long - term). Meanwhile, the unmodified electrode serves as the control group (Control). After incubation, the experimental group and control group electrodes are washed once with PBS buffer (0.01 M, pH = 7.4), then protein adhesion images are recorded with a fluorescence microscope at the same exposure time, and the relative fluorescence intensity on the substrate surface is quantitatively analyzed by ImageJ to evaluate the anti - protein adhesion level on the surface of the electrodes modified with different thicknesses of liquid - like film layers.

[0077] RAW cells and NIH / 3T3 cells are respectively co - incubated with the microneedle electrode sheets modified with different film layer thicknesses in a 37 °C, 5% CO₂ incubator for 24 h (short - term) and 7 d (long - term). Meanwhile, the unmodified electrode serves as the control group. After incubation, the cells are stained and calibrated with 10 μg / mL Hoechst33342 for the two types of cells. Then, cell adhesion images are recorded with a fluorescence microscope, and the anti - cell adhesion level on the surface of the electrodes modified with different thicknesses of liquid - like film layers is evaluated by cell counting statistics.

[0078] Microneedle electrode sheets with different thicknesses (including working electrode, reference electrode, and counter electrode) are placed horizontally on the measurement stage. Meanwhile, the unmodified electrode serves as the control group. Using the sessile drop method, 3 μL of probe liquid (deionized water, ethanol, toluene, and Dulbecco's modified Eagle medium with fetal bovine serum (DMEM - FBS)) is dropped at three random sites on the surface, and droplet images are collected within 5 s. The contact angles of different probe liquids are calculated by contact angle analysis software combined with statistical analysis to evaluate the hydrophilic - hydrophobic characteristics of each electrode. For the sliding angle test, 10 μL of the above - mentioned probe liquid is dropped, and the angle is recorded by tilting the substrate until the droplet starts to slide along the surface, and the slip adhesion is recorded.

[0079] Figure 3 Figure shows the optical microscope images of the dynamic sliding behavior of water on the surface of the electrode in Example 4 and the unmodified electrode surface. Among them, (a1) shows the state of water on the unmodified electrode (i.e., SUS304 stainless steel sheet) before sliding, (a2) shows the state of water on the unmodified electrode after sliding, and (a3) shows the wetting area left by water on the unmodified electrode after sliding; (b1) shows the state of water on the modified electrode (i.e., the microneedle electrode obtained in Example 4) before sliding, (b2) shows the state of water on the modified electrode after sliding, and (b3) shows the wetting area left by water on the modified electrode after sliding. The black part in (a3) and (b3) represents the area contaminated on the surface after liquid sliding, that is, the wetting area. From Figure 3As can be seen, when appropriately tilted, the droplets slide off the surface of the modified electrode without residue, while the unmodified electrode is wetted. The image shows the sliding process of the water droplets when tilted. Especially on the surface of the modified electrode, the water droplets can slide off rapidly without leaving any residue, indicating that the modified surface has excellent anti-adhesion properties and superhydrophobic properties. In contrast, the surface of the unmodified electrode shows the wetting phenomenon of the water droplets, indicating that the water droplets are more likely to spread and adhere on the unmodified surface.

[0080] Figure 4 The sliding angles of the modified electrode surfaces with different film layer thicknesses for various liquids, including water, ethanol, toluene, and DMEM-FBS, are shown. The sliding angle is an important parameter to measure the anti-adhesion property of the surface, which reflects the angle required for the droplet to start sliding on the inclined surface. A smaller sliding angle indicates better anti-adhesion performance of the surface and the droplet is more likely to slide on the surface. In Examples 1-8, the film layer thicknesses of Examples 4-8 are in the range of 1-4 nm, and they all have smaller contact angles, good hydrophobicity, and good anti-slip adhesion.

[0081] (3) Electrochemical performance test: The performance of the sensor is evaluated by cyclic voltammetry. The electrode is cycled between the oxidation potential and the reduction potential of the electrochemically active [Fe(CN)6] 3- / 4- solution. The current density and the peak-to-peak potential (ΔEp) of the redox process are used to evaluate the electron transfer kinetics between the electrode surface and the solution, thereby reflecting the overall quality and state of the solid-liquid interface. That is, the standardized gold electrodes modified with different types of liquid film layer thicknesses are studied in 5 mM [Fe(CN)6] 3- / 4- (i.e., [Fe(CN)6] 3- and [Fe(CN)6] 4- ) and 0.1 mM KCl solution to study the influence of the coating thickness on the current density and the peak-to-peak potential by cyclic voltammetry (fixed scan rate of 50 mV / s, scan voltage range of -0.5 V to 0.5 V).

[0082] The electron transfer resistance characteristics at the electrode interface are analyzed by EIS. That is, in 5 mM [Fe(CN)6] 3- / 4- and 0.1 mM KCl solution, the EIS of the sensor systems modified with different types of liquid film layer thicknesses is studied in the frequency range of 0.1-106 Hz.

[0083] Electrochemical detection (amperometry) was used to analyze the effect of the thickness of the liquid-like film layer on the sensitivity of the sensor. Glucose simulated solution detection test. In glucose simulated solutions with different concentration gradients, the voltage corresponding to the characteristic peak was obtained according to the cyclic voltammetry method. Then, this voltage was used as the preset bias voltage (the potential difference between the working electrode and the reference electrode) of the amperometry method to obtain the glucose response currents at different concentrations, thereby quantitatively analyzing the sensitivity and linear range of the working electrode. When analyzing the analyte concentration by electrochemical measurement, a constant bias voltage was applied to the working electrode. After the current stabilized for 2 min, the current-time (I-t) signal was recorded, and the analyte concentration was gradually increased at intervals to establish the signal-concentration correlation. The response sensitivity of different sensor systems modified with liquid-like film layers of different types to glucose was tested, analyzed, and evaluated. At the same time, the unmodified sensor system was used as a control group.

[0084] Figure 5 It is a graph of the electrochemical performance and anti-adhesion performance of the microneedle electrode. Among them, (a1) is the CV curve of the microneedle electrode with different film layer thicknesses; (a2) is the statistical graph of the current density and peak-to-peak potential of the microneedle electrode with different film layer thicknesses; (b1) is the current-time curve of the glucose sensing performance of different sensors including the liquid-like film layer (~2 nm) before and after modification and before and after protein contamination; (b2) is the signal-concentration correlation of the glucose sensing performance of different sensors including the liquid-like film layer (~2 nm) before and after modification and before and after protein contamination; (c) is the sensitivity analysis of the glucose sensing performance of different sensors including the liquid-like film layer (~2 nm) before and after modification and before and after protein contamination; (d) is the fluorescence image of the glucose sensing performance of different sensors including the liquid-like film layer (~2 nm) before and after modification and before and after protein contamination; (e) is the statistical analysis of the glucose sensing performance of different sensors including the liquid-like film layer (~2 nm) before and after modification and before and after protein contamination. The figure shows the adhesion of protein (Fg, green fluorescence) on the surfaces of different electrodes (including the working electrode and the reference electrode) after 7 days of incubation. Figure 5 It shows that the microneedle electrode prepared in the example has good electrochemical performance and anti-adhesion performance, and good results can be obtained by modification with the liquid-like film layer.

[0085] (4) Continuous sensitive monitoring: Under in vitro conditions, evaluate the continuous sensitive monitoring performance of the modified microneedle blood glucose sensor. Specifically, it includes the following: a) Evaluation of the transdermal performance of the microneedle array: Coat the tip of the microneedle array with a fluorescent dye (such as rhodamine B), insert it into the skin tissue of an in vitro rat and then remove it, and evaluate the transdermal performance through a fluorescence microscope. b) Evaluation of the continuous sensitive monitoring performance of subcutaneous glucose in vitro: In the present invention, a 0.1 mg / mL fibrinogen solution (Fg solution, fluorescent, used for subsequent analysis of the continuous sensitive monitoring mechanism) is selected as the complex biological fluid. Place the Fg solution in a Franz diffusion device, spread fresh rat skin over the device, apply the above-mentioned optimally modified microneedle blood glucose sensor with a liquid film layer on the in vitro skin and penetrate the skin, and continuously monitor the glucose concentration in the Fg solution (for 7 days). Among them, the glucose concentration is changed by dropping glucose solution at specific times to form fluctuations, and the electrical signal corresponding to the glucose level in the complex biological fluid is measured, and readings are taken once every 5 minutes. At the same time, the unmodified microneedle blood glucose sensor is used as a control group to evaluate the stability of continuous sensitive monitoring.

[0086] Evaluate the continuous sensitive monitoring performance of the microneedle blood glucose sensor using type I diabetic model rats. Select rats weighing 200 - 250 g for diabetes modeling. 12 hours before modeling, fast the rats without water deprivation treatment, and then intraperitoneally inject streptozotocin (60 mg·kg -1 ) into the rats. Feed them normally for 7 days, and test the blood glucose (BG) concentration. When it exceeds 16.7 mmol / L for two consecutive days, the modeling is successful. Place the above-mentioned optimally modified microneedle blood glucose sensor with a liquid film layer on the back skin of the diabetic rats for continuous blood glucose monitoring (for 7 days), measure the electrical signal corresponding to the interstitial fluid blood glucose in the subcutaneous tissue, take readings once every 5 minutes, and use the unmodified microneedle blood glucose sensor as a control group. During the entire recording period, tail blood is taken at specific time points, and the actual blood glucose value is measured using a commercial blood glucose meter for comparison. In addition, during the entire test process, a quantitative glucose solution is intraperitoneally injected into the rats at different time points to induce blood glucose fluctuations in the body, so as to verify the accuracy of blood glucose monitoring under different blood glucose states. Analyze the error between the blood glucose values measured by the modified sensor and the unmodified sensor and the actual values, and evaluate the sensitivity, stability, and accuracy of continuous blood glucose monitoring of the modified sensor.

[0087] The biosafety of the microneedle blood glucose sensor is evaluated by measuring the tissue inflammation, allergic reactions, tissue fibrosis, etc. that may be induced during the transdermal monitoring process. The specific experiment is divided into 5 groups. Group 1: The microneedle blood glucose sensor is removed after being implanted subcutaneously for 5 minutes. Groups 2 and 3: The microneedle blood glucose sensor is removed after being implanted subcutaneously for 1 day and 7 days respectively. Groups 4 and 5: The microneedle blood glucose sensor is removed after being implanted subcutaneously for 1 day and 7 days respectively, but a bias voltage of 0.5 - 1 V is continuously applied during this period to monitor blood glucose fluctuations (sampling once every 5 minutes). After the microneedle sensor is removed, all rats are continued to be fed for 7 days, and the skin at the sensor application site is fixed, embedded, sectioned, stained with H&E, and observed under a microscope for skin tissue inflammation, allergy, and fibrosis conditions. The normal group (skin tissue in the same area without applying the sensor) is used as the control group to evaluate the biosafety of the sensor application.

[0088] Figure 6 To show the fluorescence images of the adhesion of fluorescent fibrinogen on the surface of modified electrodes with different thicknesses of quasi-liquid film layers. Among them, the green fluorescence shows fluorescent fibrinogen Fg. (a) is the fluorescence image of Fg adhesion for 1 day, and (b) is the fluorescence image of Fg adhesion for 7 days. This figure captures the adhesion of fibrinogen on the modified electrodes with different thicknesses of quasi-liquid film layers through a fluorescence microscope. The distribution and density of the green fluorescence intensity can intuitively display the degree of protein adhesion. By comparing the images after 1 day and 7 days, the change of protein adhesion over time and the effect of different thicknesses of quasi-liquid film layers on protein adhesion can be evaluated.

[0089] Figure 7 To show the statistical analysis chart of the adhesion of fluorescent fibrinogen on the surface of modified electrodes with different thicknesses of quasi-liquid film layers. Among them, the green fluorescence shows fluorescent fibrinogen Fg. (a) is the statistical analysis chart of Fg adhesion for 1 day, and (b) is the statistical analysis chart of Fg adhesion for 7 days. This figure provides Figure 6 the quantitative data of the fluorescence images in. The fluorescence intensity is quantified through statistical analysis methods, so as to obtain the accurate numerical values of the fibrinogen adhesion amount on the surface of modified electrodes with different thicknesses of quasi-liquid film layers. These data can be used to evaluate the trend of protein adhesion over time and the effect of different thicknesses of quasi-liquid film layers on reducing protein adhesion. It can be seen from Figures 6 - 7 that compared with the unmodified electrode, the microneedle electrode prepared in the embodiment of the present invention has a good anti-protein adhesion effect.

[0090] Figure 8To show the statistical analysis chart of the adhesion of NIH-3T3 cells on the surface of modified electrodes with different thicknesses of liquid-like film layers, the adhesion of NIH-3T3 cells is marked with green fluorescence, and the adhesion time is 1 day. This chart provides the adhesion situation of the cells on the modified electrodes with different thicknesses of liquid-like film layers after 1 day through statistical analysis. The statistical data of the number and distribution of the cells marked with green fluorescence can be used to evaluate the influence of different thicknesses of liquid-like film layers on cell adhesion and the regulatory effect of the coating on cell adhesion behavior.

[0091] Figure 9 To show the statistical analysis chart of the influence of the modified electrode on cell viability. Through statistical analysis methods, the data on the influence of the modified electrode on cell viability are provided. This may include the measurement results of cell survival rate, proliferation rate or other cell viability indicators. These data are crucial for evaluating the biocompatibility and potential cytotoxicity of the modified electrode.

[0092] Figure 10 To show the fluorescence images of the adhesion of NIH-3T3 cells on the surface of modified electrodes with different thicknesses of liquid-like film layers and the fluorescence images of the influence of the modified electrode on cell viability. Among them, the green fluorescence shows NIH-3T3 cells.

[0093] (a) is the fluorescence image of the adhesion of NIH-3T3 cells on the surface of modified electrodes with different thicknesses of liquid-like film layers after 1 day, and (b) is the fluorescence image of the influence of the modified electrode on cell viability. This figure combines Figure 8 and Figure 9 information, and intuitively shows the adhesion situation of NIH-3T3 cells on the modified electrodes with different thicknesses of liquid-like film layers and the influence of the modified electrode on cell viability through fluorescence images. The cell distribution marked with green fluorescence and viability indicators (such as the ratio of live cells to dead cells) can intuitively show the biocompatibility of the modified electrode and its influence on cell behavior. It can be seen from Figures 8 - 10 that compared with the unmodified electrode, the microneedle electrode prepared in the embodiment of the present invention has good anti-cell adhesion effect, good biocompatibility and low cytotoxicity.

[0094] Figure 11This is the schematic diagram of the present invention for blood glucose detection. Among them, (a) is the detailed schematic diagram of blood glucose detection, (b1) is the skin structure diagram during the blood glucose detection process, (b2) is the partial enlarged view of the skin structure, and (c1) and (c2) respectively represent the blood glucose detection curve and anti-adhesion curve obtained during the detection process. It can be seen that by using the microneedle electrode of the present invention and controlling the thickness of the liquid-like film layer, slippage can be promoted, thereby achieving high anti-adhesion performance; good electrochemical performance can also be maintained, thereby achieving high sensitivity; thus, continuous and sensitive blood glucose monitoring can be realized. The microneedle sensor prepared by the microneedle electrode of the present invention has good blood glucose monitoring effect, and the specific process is as follows: 1) The microneedle blood glucose sensor is inserted into the skin, and the liquid-like film layer covers the microneedles; 2) Glucose enters the sensor through the microneedles and reacts with glucose oxidase; 3) Pt nanoparticles act as catalysts to accelerate the oxidation reaction of glucose; 4) The oxidation reaction generates an electrical signal, and the actual blood glucose value is converted through the current change, thereby testing the blood glucose concentration; 5) The anti-adhesion performance of the liquid-like film layer with a certain thickness reduces the adhesion of proteins and cells, maintaining the high sensitivity and electrochemical performance of the sensor; 6) The system can continuously monitor the blood glucose level, provide real-time data, and help with the remote management of diabetes.

[0095] As can be seen from the embodiments of the present invention, by using the microneedle electrode of the present invention, long-term inhibition of biomaterial adhesion (more than 7 days) can be achieved, and the inhibition efficiency of biomaterial adhesion reaches more than 90%, ensuring the stability of continuous monitoring. It is feasible to prepare the liquid-like film layer with a thickness controllable at the nanoscale, with a thickness of about 2 nm, which can efficiently inhibit the adhesion of biomaterials and ensure that the sensitivity is not affected, thus making continuous and sensitive monitoring theoretically feasible. The "liquid-like" molecular modified lubricating surface, that is, covalently coupling a layer of low surface energy polymer (such as PDMS), only changes the surface topography at the nanoscale, and due to the flexibility and fluidity of this polymer, it is beneficial for the slippage of various liquids, thereby being able to reduce the interaction with biomaterials and significantly inhibit the adhesion of proteins, bacteria, and cells; furthermore, it can reduce adverse reactions and tissue damage and improve biocompatibility. The device developed using the microneedle electrode in the embodiments of the present invention can achieve continuous and sensitive blood glucose monitoring in in vitro complex biological fluids and in vivo environments without causing inflammation. And through cell culture and animal models, it is proved that the level of inflammatory cytokines can be reduced.

[0096] In summary, the present invention constructs a liquid-like film layer with a low surface energy polymer on the substrate surface. Due to the flexibility, smoothness, and fluidity of the low surface energy polymer, it is beneficial for the slippage of various liquids, thereby being able to reduce the interaction between the microneedle electrode and various biomaterials, significantly inhibit the adhesion of proteins, bacteria, cells, etc., and obtain a microneedle electrode with good anti-adhesion performance, which has good application prospects in blood glucose detection.

Claims

1. A microneedle electrode, characterized in that: The invention comprises a substrate and a liquid film-like layer covalently coupled to the surface of the substrate; the raw material for preparing the liquid film-like layer comprises a low surface energy polymer; the surface energy of the low surface energy polymer is ≤78mJ / m 2 .

2. The microneedle electrode according to claim 1, characterized in that: The low surface energy polymer includes at least one of polydimethylsiloxane, polytetrafluoroethylene, polyvinylidene fluoride, polystyrene or polytrifluoroethylene; And / or, the thickness of the liquid-film-like layer is 1 to 10 nm.

3. The microneedle electrode according to claim 1, characterized in that: The substrate has a planar microneedle array; the planar microneedle array has 3 to 7 microneedles; and the length of each microneedle is 600 to 1000 μm.

4. The microneedle electrode according to claim 1, characterized in that: A catalytic substance is also arranged between the substrate and the liquid-like film layer; the catalytic substance includes glucose oxidase and platinum nanoparticles.

5. A method for preparing a microneedle electrode according to any one of claims 1 to 4, characterized in that: The following steps are involved: Prepare a liquid film-like layer on the surface of the substrate to obtain the microneedle electrode; Alternatively, a catalytic substance is disposed on the surface of the substrate, and then a liquid-like film layer is prepared to obtain the microneedle electrode.

6. The preparation method according to claim 5, characterized in that: Before preparing the liquid-like film layer, the method further includes a step of modifying the surface of the substrate with hydroxyl groups.

7. The preparation method according to claim 5, characterized in that: The method for preparing the liquid-like film layer is selected from the atomic deposition method; Preferably, the deposition time of the atomic deposition method is 0.5 min to 5 h.

8. A microneedle sensor, characterized in that: The microneedle electrode comprises the microneedle electrode according to any one of claims 1 to 4.

9. A blood glucose monitor, characterized in that: It comprises the microneedle electrode according to any one of claims 1 to 4, or the microneedle sensor according to claim 8.

10. The blood glucose monitor according to claim 9, characterized in that: It also includes at least one of a wireless data transmission module, an artificial intelligence analysis algorithm module or a user interaction module.

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