CN210990322U - Highly sensitive wearable non-invasive glucose sensor - Google Patents

Abstract

The utility model relates to a glucose sensor, especially a wearable noninvasive glucose sensor of high sensitivity, it includes: a base layer (10); a counter electrode layer (20) and a working electrode layer (30) formed on the base layer (10); a nano noble metal particle accumulation layer (40) is formed above the working electrode layer (30); an electron mediator layer (50) is formed on the noble metal nanoparticle deposited layer (40), and a glucose oxidase layer (60) is provided on the electron mediator layer (50). The utility model discloses a design can laminate the stratum basale of skin and make glucose sensor have wearable, through set up nanometer noble metal granule accumulation layer in order to play the catalytic action between glucose oxidase layer and electrode layer to showing sensitivity and the stability that improves glucose sensor, impel the further development and the application of not having wound type glucose sensor.

Description

Highly sensitive wearable non-invasive glucose sensor

technical field

The utility model relates to a glucose sensor, in particular to a high-sensitivity wearable non-invasive glucose sensor.

Background technique

Diabetes, which prevents the body from maintaining normal blood sugar levels, is one of the common diseases that can lead to incurable death. According to the International Diabetes Federation, in 2017, the number of people with diabetes in the world has reached 424 million, and by 2045, the number of people with diabetes will reach 628 million. The number of diabetic patients in China has exceeded 114 million, ranking first in the world. Accurate detection of glucose levels in human blood is essential for the treatment and prevention of diabetes. Blood glucose monitoring methods can be divided into invasive measurement, minimally invasive measurement and non-invasive measurement. At present, blood glucose is mainly monitored by invasive measurement (such as intravenous indwelling needle monitoring and fingerstick monitoring), which not only brings pain to patients, but also easily causes wound infection. Minimally invasive measurement generally obtains glucose signals through a probe implanted in the human skin. Although it can relieve the patient's pain to a certain extent, it needs to be implanted into the skin through surgery, which is very inconvenient to use.

Most of the existing non-invasive glucose sensors detect metabolic fluids such as sweat, saliva, and tears to determine the blood sugar content of the human body. However, due to the low glucose content in body fluids, the measurement results are inaccurate and the stability is poor. However, there are fewer interfering factors in tissue fluid, which has obvious advantages. The extraction of interstitial fluid is mainly by reverse ion osmosis. Its working principle is that under the action of an external electric field, sodium ions and chloride ions undergo electromigration in the subcutaneous interstitial fluid to move to the positive and negative electrodes of the electrodes to form tiny direct current channels. Since the human skin itself has a negative charge, it is mainly the ion flow formed by the migration of sodium ions to the negative electrode under the action of electric potential. Using this ion current as a channel can carry out neutral glucose molecules in the tissue fluid at the negative electrode of the electrode, and at the same time, the hyaluronic acid solution is applied to the surface of the skin to increase the surface charge transfer rate to increase the amount of tissue fluid exported. This part of the tissue fluid can replace blood as a test fluid to detect blood sugar levels. This method delivers molecules through the skin without damaging the skin's surface or coming into contact with the blood. However, due to the small amount of extracted human tissue fluid and the low glucose content, it is urgent to improve the sensitivity and detection stability of the glucose sensor, and at the same time improve the biocompatibility of the sensor applied to human skin.

Utility model content

(1) Technical problems to be solved

In order to solve the above problems of the prior art, the present invention provides a high-sensitivity wearable non-invasive glucose sensor design scheme. The glucose sensor is wearable by designing a base layer that can be attached to the skin. A nano-precious metal particle stacking layer is arranged between the electrode layer to improve the catalytic effect, thereby effectively and greatly improving the sensitivity and stability of the glucose sensor, and promoting the development and application of the non-invasive glucose sensor.

(2) Technical solutions

In order to achieve the above-mentioned purpose, the main technical scheme adopted by the present utility model includes:

A highly sensitive wearable non-invasive glucose sensor comprising:

a base layer (10); a counter electrode layer (20) and a working electrode layer (30) formed on the base layer (10);

A nano-precious metal particle accumulation layer (40) is formed above the working electrode layer (30); an electron mediator layer (50) is formed on the nano-precious metal particle accumulation layer (40), and an electron mediator layer (50) is formed on the nano-precious metal particle accumulation layer (40). A glucose oxidase layer (60) is provided on (50).

According to a preferred embodiment of the present invention, the working electrode layer (30) and the counter electrode layer (20) are conductive carbon films, metallic silver films, metallic copper films or gold films.

According to a preferred embodiment of the present invention, wherein the working electrode layer (30) and the counter electrode layer (20) are deposited on the base layer (10) by an evaporation method; or, the working electrode layer (30) and the counter electrode layer (20) are formed on the base layer (10) by means of spin coating, brush coating, spray coating, screen printing or printing.

According to a preferred embodiment of the present invention, the nano-precious metal particle accumulation layer (40) is a nano-gold, nano-platinum or nano-silver particle accumulation layer.

According to a preferred embodiment of the present invention, the nano-precious metal particle accumulation layer (40) is a nano-gold particle accumulation layer.

According to a preferred embodiment of the present invention, wherein, the nano-precious metal particle accumulation layer (40) is formed on the working electrode layer (30) by a logistic method such as ion beam sputtering, magnetron sputtering or physical vapor deposition. )superior.

Ion beam sputtering has the characteristics of simple operation technology, easy and precise control, and suitable for mass production, and ion beam sputtering can form a stacking layer of nano-metal particles with smaller diameters, so that the nano-precious metal particles have good catalytic activity and larger of specific surface area, roughness and good biocompatibility.

According to a preferred embodiment of the present invention, the particle size of the nano-precious metal particles in the nano-precious metal particle accumulation layer (40) is 0.1 nm-100 nm. Preferably, the nano-precious metal particle accumulation layer (40) is a nano-gold particle accumulation layer, and the particle size of the nano-gold particles is less than 5 nm.

According to a preferred embodiment of the present invention, the thickness of the nano-precious metal particle accumulation layer (40) is 1 nm-1 mm.

According to a preferred embodiment of the present invention, wherein the electron mediator layer (50) is ferric ferrocyanide (PB), ferrocene and its derivatives, flavin adenine dinucleotide (FAD) ), benzoquinone or polytetrafluoroethylene, and organic dyes.

The electron mediator has the advantages of promoting the electron transfer process and broadening the linear range. The commonly used material for the electron mediator layer (50) is PB or FAD. The PB can be formed on the nano-precious metal particle stacking layer (40) by an electrochemical deposition method.

According to a preferred embodiment of the present invention, the glucose oxidase layer (60) is made by fixing the glucose oxidase on the electron mediator layer (50) through a porous polymer. Since simple glucose oxidase is not easy to form and immobilize, when making glucose sensors, it is usually necessary to disperse glucose oxidase in a conductive polymer (preferably a porous conductive polymer) to achieve immobilization and forming (molding into a film or layer). so the concept of the glucose oxidase layer (60) belongs to the prior art.

Preferably, the porous polymer is chitosan or other conductive polymer aerogels. Chitosan is a biocompatible polymer matrix, which can well disperse and immobilize glucose oxidase (GOD), and has good film-forming ability and high water permeability, as well as certain electrical conductivity. It is therefore the preferred material for immobilizing glucose oxidase.

According to a preferred embodiment of the present invention, the base layer (10) is a flexible base layer or a hard base layer; the flexible base layer is a flexible polymer resin material substrate, etc., and the hard base layer For rigid plastic substrate or rigid paper, etc.

According to a preferred embodiment of the present invention, the base layer (10) is a flexible base layer, which comprises a polymethyl methacrylate (PMMA) layer (11) located below and a polyimide layer located above An amine (PI) film layer (12), the counter electrode layer (20) and the working electrode layer (30) are arranged on the polyimide film layer (12).

Preferably, the Young's modulus of the flexible base layer is 1-200 MPa, preferably 4-5 MPa, which can well bend and fit the skin.

Among them, the PMMA layer has high toughness, high elongation rate and high transparency, high tenacity can provide the strength of the entire wearable non-invasive glucose sensor, and its high elongation rate can improve the comfort of patients when using it, which can be adjusted according to the user The body part to be pasted is properly bent and stretched for a good fit. One side of the PI film layer faces the user's skin side, the PI has good thermal stability and good biocompatibility, and can reduce skin discomfort (foreign body sensation) when it is closely attached to the skin. PMMA and PI together constitute a composite flexible substrate with flexibility and stretch rate, which can better fit the skin, improve the user's comfort, and improve the biocompatibility of the wearable non-invasive glucose sensor.

(3) Beneficial effects

The beneficial effects of the present invention are:

The utility model can greatly improve the detection sensitivity and detection stability of the glucose sensor by arranging a catalytic layer formed by stacking a layer of nano precious metal particles on the electronic mediator layer and the working electrode layer, so that the non-invasive glucose sensor, especially the interstitial fluid reverse ion The application of osmotic non-invasive glucose sensors has been further developed.

Experiments show that the glucose sensor with the nano-precious metal particle stacking layer of the present invention is 2-3 times higher than the glucose sensor without the nano-precious metal particle stacking layer.

In the present invention, the working electrode layer, the electronic medium layer, the glucose oxidase layer, the counter electrode layer, etc. are further arranged on the composite flexible base layer made of PMMA and PI, and the flexible base layer can be attached to the user's skin. The high biocompatibility of PI and the high toughness and tensile properties of PMMA enhance the biocompatibility of the wearable non-invasive glucose sensor, thereby improving the user's comfort.

Description of drawings

FIG. 1 is a schematic structural diagram of a preferred embodiment of the glucose sensor of the present invention.

Fig. 2 is the working principle diagram of the glucose sensor of the present invention.

FIG. 3 is a SEM image of the surface of the gold film working electrode formed by the evaporation method (ion beam sputtering for 0 seconds).

FIG. 4 is a SEM image of the surface of the gold nanoparticle accumulation layer formed by ion beam sputtering for 60 seconds.

FIG. 5 is an AFM (atomic force microscope) image of the gold film on the surface of the working electrode formed by evaporation method (ion beam sputtering for 0 seconds).

FIG. 6 is an AFM image of the surface of the gold nanoparticle accumulation layer formed by ion beam sputtering for 60 seconds.

FIG. 7 is the time-current curve of the glucose sensor without the gold nanoparticle stacking layer prepared by ion beam sputtering for 0 seconds.

FIG. 8 is the time-current curve of the glucose sensor prepared by ion beam sputtering for 60 seconds and containing the stacked layer of gold nanoparticles.

FIG. 9 is the tensile curve of the flexible substrate of the glucose sensor of the present invention.

[Description of reference numerals]

10 base layer; 20 pair electrode layer; 30 working electrode layer; 40 nanometer precious metal particle stacking layer; 50 electron mediator layer; 60 glucose oxidase layer; 11 polymethyl methacrylate layer; 12 polyimide film layer.

Detailed ways

In order to better explain the present invention and facilitate understanding, the present invention will be described in detail below with reference to the accompanying drawings and through specific embodiments.

As shown in FIG. 1, the high-sensitivity wearable non-invasive glucose sensor of the present invention includes: a base layer 10, a counter electrode layer 20, a working electrode layer 30, a nano-precious metal particle stacking layer 40, an electronic mediator layer 50, a glucose Oxidase layer 60 .

Wherein, the counter electrode layer 20 and the working electrode layer 30 are formed on the base layer 10, the nano-precious metal particle accumulation layer 40 is formed on the working electrode layer 30; the electron mediator layer 50 is formed on the nano-precious metal particle accumulation layer 40, and the electronic media A glucose oxidase layer 60 is provided on the bulk layer 50 . The characteristics and functions of the above layers are described as follows:

The working electrode layer 30 and the counter electrode layer 20 are conductive carbon films, metallic silver films, metallic copper films or gold films, which can be deposited on the base layer 10 by an evaporation method. Preferably, the working electrode layer 30 and the counter electrode layer 20 are a layer of gold film formed on the flexible substrate 10 by evaporation method, and the thickness is 50-300 nm.

The nano-precious metal particle stacking layer 40 is a stacking layer of nano-platinum, nano-silver, nano-gold and other precious metal particles, which not only has high conductivity but also has high catalytic activity. Electrocatalytic activity of hydrogen.

Preferably, the nano-precious metal particle accumulation layer 40 is a accumulation layer of nano-gold particles, and the particle size of the nano-gold particles is 0.1 nm-100 nm, more preferably ≤ 5 nm. In order to form the gold nanoparticle stacking layer with catalytic properties, the ion beam sputtering method is preferentially formed. The method of ion beam sputtering can not only form a stacked layer of gold nanoparticles, but also form a rough surface with unevenness (see Figures 4 and 6), so as to obtain good catalytic activity, a larger specific surface area, and a larger roughness and good biocompatibility. Preferably, the thickness of the nano-precious metal particle accumulation layer 40 is 1 nm-1 mm; the thickness can be precisely controlled by the application time of ion beam sputtering, and the ion beam sputtering time is 5-150 seconds.

The electron mediator layer 50 is a reducing compound, which can promote the electron transfer process and widen the linear range. The material used for the electron mediator layer 50 is ferric ferrocyanide (PB Prussian blue), ferrocene and its derivatives, flavin adenine dinucleotide (FAD), benzoquinone, polytetrafluoroethylene or organic Dyes, etc. Wherein, when PB is used, it can be deposited on the nano-precious metal particle stacking layer 40 by an electrochemical deposition method. The most commonly used material for the electronic mediator layer 50 is PB or FAD.

The glucose oxidase layer 60 is formed by immobilizing the glucose oxidase on the electron mediator layer 50 through a porous polymer. Since simple glucose oxidase is not easy to form and immobilize, when making a glucose sensor, it is usually necessary to disperse glucose oxidase in a conductive polymer (preferably a porous conductive polymer) to achieve immobilization and forming (molding into a film or layered).

In the preparation process of the glucose oxidase layer 60, the glucose oxidase can be dissolved in deionized water, then mixed with the chitosan solution and stirred until it is clear without bubbles, and the mixed glucose oxidase and chitosan mixed solution is dripped onto the On the electron mediator layer 50, the mixed solution solvent is evaporated to dryness after being placed at room temperature for 2-7 hours. Chitosan is a biocompatible polymer matrix, which can well disperse and immobilize glucose oxidase, and has good film-forming ability and high water permeability. Preferred material for immobilizing glucose oxidase.

Further, the base layer 10 is a flexible base layer with a Young's modulus of 4-5 MPa, which can well bend and fit the skin, and includes a polymethyl methacrylate (PMMA) film layer 11 located below and a layer located above. The polyimide (PI) film layer 12, the counter electrode layer 20 and the working electrode layer 30 are arranged above the PI film layer 12, and one side of the PI film layer 12 is used to stick to the surface of the human skin, and the PMMA film layer can support and The PI film layer, the counter electrode layer 20, the working electrode layer 30, the nano-precious metal particle accumulation layer 40, the electron mediator layer 50 and the glucose oxidase layer 60 on the protective inner side are tightly attached to the human skin, and by contacting the human skin Transmitted cells and tissue fluid, the reaction process shown in Figure 2 occurs. PI has good thermal stability and good biocompatibility, and can reduce skin discomfort when it is closely attached to the skin.

In addition to being a flexible base layer, the base layer 10 can also be a rigid base layer, such as a rigid plastic sheet substrate or a rigid paper substrate. However, in the present invention, the flexible base layer of the PMMA+PI composite structure is preferentially used. On the one hand, these two materials have very excellent properties, such as stretchability, transparency, high temperature resistance and biocompatibility, etc.; On the other hand, the flexible base layer of the composite structure has a very suitable Young's modulus.

As shown in Figure 2, glucose is oxidized by O ₂ and loses electrons to become gluconic acid. The electrons combine with water and O ₂ to generate H ₂ O ₂ . The electrons are transferred by the electron transfer mediator layer 50 (PB) to form a current. The nano-gold particles in the noble metal particle stacking layer 40 play a catalytic role, which can significantly enhance the sensitivity of the sensor. The detection terminal collects the current and voltage signals brought by the electron flow, and realizes the quantitative detection of the glucose concentration by utilizing the characteristic that the current or voltage signal has a linear relationship with the glucose concentration.

Example 1

The highly sensitive wearable non-invasive glucose sensor of this embodiment can be prepared according to the following method:

Step 1: Polymethyl methacrylate (PMMA, dissolved in chlorobenzene with stirring at 80°C, 300 mg/mL) was spin-coated on a smooth silicon wafer cleaned by UV-ozone treatment for 10 minutes at 6000 rpm for 30 seconds. , the spin-coated PMMA was heated on a heating plate at 180°C for 10 seconds, and then the polyamic acid was rotated at 6000 rpm for 30 seconds and spin-coated on the PMMA layer, and heated at 80°C, 120°C, and 140°C for three times. The temperature was maintained for one hour each to obtain a flexible substrate 10 .

The tensile curve of the flexible substrate 10 is shown in Fig. 9, and the Young's modulus is calculated to be 4.75 MPa, which can indeed bend, stretch and fit the skin very well.

Step 2: depositing gold with a thickness of 50-300 nm on the prepared flexible substrate 10 by thermal evaporation at an evaporation rate of 0.1-1 angstrom/second to obtain a counter electrode layer 20 and a working electrode layer 30 . In this embodiment, the thicknesses of the counter electrode layer 20 and the working electrode layer 30 are 200 nm.

Step 3: under vacuum conditions, sputter gold nanoparticles on the surface of the working electrode layer 30 (gold film formed by evaporation) for 5-150 seconds. The particle size of the gold nanoparticles formed by sputtering is about 5 nm, or even very small (see Figures 4 and 6), so their catalytic activity is higher. See Figures 3-4 for the SEM images detected at 0 seconds and 60 seconds from the start of sputtering. It can be seen that when sputtering for 60 seconds, the SEM can observe the surface of the accumulation of gold nanoparticles.

See Figures 5-6 for AFM images measured at 0 seconds and 60 seconds after sputtering. It can be seen that when sputtering for 60 seconds, a very uniform and fine (obviously finer than Fig. 5) nano-gold particle accumulation layer is obtained, and the surface is uneven. These nano-gold particle stacking layers play a very good catalytic role, and correspondingly constitute the above-mentioned nano-precious metal particle stacking layers 40 .

Step 4: Take 40.6 mg of ferric chloride (FeCl ₃ ), 82.31 mg of potassium ferricyanide (K ₃ Fe(CN) ₆ ) and 745.5 mg of potassium chloride (KCl) in a 100 mL volumetric flask, add 829 μL of hydrochloric acid (concentration 99%) using deionized water as solvent to prepare ferric ferrocyanide solution. In the three-electrode system, the voltage was set to 0.4V, and PB was electrochemically deposited on the nano-gold particle stacking layer (ie, the nano-precious metal particle stacking layer 40 ) using the chronoamperometry method, and then dried at 60° C. for 5 minutes.

Step 5: Dissolve 20 mg of chitosan into 2% acetic acid solution, add 8 μL of glycerol, and stir evenly at 80° C. until it is clear without bubbles. Glucose oxidase solution (dissolved in deionized water, 30 mg/mL) was prepared, and the chitosan solution was mixed with the glucose oxidase solution and stirred until it was clear without bubbles. The mixed solution of glucose oxidase and chitosan is drop-coated on the PB layer and placed at room temperature for 2-7 hours until the solvent evaporates to dryness.

Step 6: Lift off the above completed structure from the silicon wafer to obtain a flexible sensor, and obtain a highly sensitive wearable non-invasive glucose sensor with the structure shown in FIG. 1 .

Figure 7 is the time-current curve of the glucose sensor without the gold nanoparticle stacking layer in response to glucose, and Figure 8 is the time-current curve of the glucose sensor containing the gold nanoparticle stacking layer prepared by ion beam sputtering for 60 seconds. , the responses of the two glucose sensors were tested by electrochemical chronoamperometry in a three-electrode system. Under the condition that the thickness and preparation conditions of the counter electrode layer 20, the working electrode layer 30, the electron mediator layer 50, and the glucose oxidase layer 60 of the two glucose sensors are the same, it can be seen from the comparison of the two figures that the average of the former per millimol The current was 88.11 μA/mM, and the average current per mmol of the latter was 217.78 μA/mM, which was 2.47 times that of the former. It can be seen that the response of the glucose sensor of the PB layer modified by the gold nanoparticle stacking layer is more than twice that of the glucose sensor not modified by the gold nanoparticle stacking layer.

In addition, by comparing the cyclic voltammetry curves of the PB layer of the gold nanoparticle accumulation layer sputtered for 60 seconds and the PB layer of the gold nanoparticle accumulation layer sputtered for 0 seconds, it can be seen that the cyclic voltammetry curve of the former PB layer is better than that of the unsputtered layer. The peak current of gold nanoparticles was more than doubled through (sputtering 0 s). Experiments show that the impedance of the PB layer containing the stacked layer of nano-gold particles becomes smaller, and it is easier to detect the current generated by the oxidation of glucose, which greatly improves the sensitivity and detection stability of the glucose sensor.

The above are only preferred embodiments of the present invention, and are not intended to limit the present invention. Although the present invention has been described in detail with reference to the foregoing embodiments, those skilled in the art can still understand the foregoing. The technical solutions described in the embodiments are modified, or some technical features thereof are equivalently replaced. Any modification, equivalent replacement, improvement, etc. made within the spirit and principle of the present invention shall be included within the protection scope of the present invention.

Claims (10)

1. a highly sensitive wearable non-invasive glucose sensor, is characterized in that, it comprises: a base layer (10); a counter electrode layer (20) and a working electrode layer (30) formed on the base layer (10); A nano-precious metal particle accumulation layer (40) is formed above the working electrode layer (30); an electron mediator layer (50) is formed on the nano-precious metal particle accumulation layer (40), and an electron mediator layer (50) is formed on the nano-precious metal particle accumulation layer (40). A glucose oxidase layer (60) is provided on (50). 2. The highly sensitive wearable non-invasive glucose sensor according to claim 1, wherein the working electrode layer (30) and the counter electrode layer (20) are conductive carbon films, metallic silver films, metallic copper films or Gold film. 3. The highly sensitive wearable non-invasive glucose sensor according to claim 2, wherein the working electrode layer (30) and the counter electrode layer (20) are deposited on the base layer (10) by an evaporation method Alternatively, the working electrode layer (30) and the counter electrode layer (20) are formed on the base layer (10) by spin coating, brush coating, spray coating, screen printing or printing. The high-sensitivity wearable non-invasive glucose sensor according to claim 1, 2 or 3, characterized in that, the nano-precious metal particle accumulation layer (40) is a nano-platinum, nano-silver or nano-gold particle accumulation layer. 5. The high-sensitivity wearable non-invasive glucose sensor according to claim 4, characterized in that, the nano-precious metal particle accumulation layer (40) is formed by ion beam sputtering, magnetron sputtering or physical vapor deposition on the on the working electrode layer (30). 6 . The high-sensitivity wearable non-invasive glucose sensor according to claim 5 , wherein the particle size of the nano-precious metal particles in the nano-precious metal particle stacking layer ( 40 ) is 0.1 nm-100 nm. 7 . The high-sensitivity wearable non-invasive glucose sensor according to claim 5 or 6, characterized in that, the thickness of the nano-precious metal particle stacking layer (40) is 1 nm-1 mm. 8. The high-sensitivity wearable non-invasive glucose sensor according to claim 1, wherein the base layer (10) is a flexible base layer or a hard base layer; the flexible base layer is a flexible polymer resin material Substrate, the hard base layer is a hard plastic substrate or hard paper. 9. The high-sensitivity wearable non-invasive glucose sensor according to claim 1, wherein the base layer (10) is a flexible base layer, which comprises a polymethyl methacrylate layer (11) and The upper polyimide film layer (12), the counter electrode layer (20) and the working electrode layer (30) are arranged above the polyimide film layer (12). 10 . The high-sensitivity wearable non-invasive glucose sensor according to claim 9 , wherein the Young's modulus of the flexible base layer is 1-200 MPa. 11 .

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