CN106725470B - A continuous or discontinuous physiological parameter analysis system - Google Patents

Abstract

The invention discloses a continuous or discontinuous physiological parameter analysis system. The system includes implanted sensor electrodes for sensing analytes, an internal transmitter, and an external processor. The internal transmitter contains a voltage source that provides a continuously varying DC voltage and modules for signal measurement, analysis, processing, and transmission. The external processor can store, analyze, and process data, and includes other physiological parameter monitoring functions. Also included is a connection device for connecting the sensor electrodes to the voltage source and a method for analyzing the resistance of the connection device. The application of a continuous voltage can cause one or more electrochemical reactions to occur with one or more chemicals in the tissue fluid of an organism. The system of the invention can realize continuous or discontinuous monitoring of biological physiological parameters, and the results are accurate, convenient and suitable for industrialization.

Description

A continuous or discontinuous physiological parameter analysis system

Technical field

The invention relates to the field of physiological parameter analysis in living organisms, and specifically to a continuous or discontinuous physiological parameter analysis system.

Background technique

With the development of science and technology and the improvement of the level of medical detection, the use of biosensors is becoming more and more widespread. Biosensors measure components in organisms by converting chemical substances in organisms into detectable signals such as current signals, light signals, heat signals, etc. There are many types of biosensors, such as biological enzyme electrodes, chemical electrodes, physical electrodes, electrochemical electrodes, spectrophotometer electrodes, polarization measuring electrodes, heat measuring electrodes, radiation measuring electrodes, immunochemical electrodes, etc.

For example, glucose monitoring in the human body. Diabetes is commonly known as the disease of affluence. With the improvement of living standards, the prevalence is also increasing. Currently, diabetic patients mostly inject insulin to keep their blood sugar stable. However, excessive insulin can cause hypoglycemia three times more often than traditional treatment. Although insulin injections effectively reduce blood sugar levels, they can cause complications. Therefore, patients should monitor their blood sugar levels multiple times before and after injecting insulin. Accurate blood sugar monitoring is the key to effectively controlling blood sugar. It helps assess the degree of metabolic disorder in diabetic patients, formulate a blood sugar-lowering plan, reflect the effect of blood sugar-lowering treatment, and adjust the treatment plan in a timely manner.

US Patent Application No. US20130126349 discloses a method for preparing a continuous biosensor. The method provides a method for preparing an electrochemical biosensor electrode, including the structure of the electrode and the composition of the polymer membrane.

Chinese patent number CN102469966B discloses a continuous analyte measurement system and a system and method for implanting them. The system consists of a base, sensors, electronic components, and processing units.

The Chinese utility model patent with authorization announcement number CN201492421U discloses a dynamic blood glucose meter that can realize dynamic blood glucose detection. It consists of a sensor, a transmitter, and a receiving processor.

Existing sensors have various ways of immobilizing enzymes, such as US patent application US20050115832, US patent No. US20050027463-A1; US20060020187-A1; US-2005-0143635-A1; US-2007-0213611-A1; US-2008-0083617 -A1 and US6,565,509; US6,579,690; US6,484,046; US6,001,067; US6,702,857, etc. Moreover, during the use of these sensors, short circuits and open circuits may occur due to temperature, humidity, and assembly, causing monitoring failures. Especially in the early stages of detection, interference caused by small molecule interfering substances may lead to inaccurate detection results, and the sensor electrode itself may also be unstable. In view of the above reasons, it is necessary to improve the continuous or discontinuous physiological parameter analysis system.

Contents of the invention

In view of the shortcomings of the existing technology, the purpose of the present invention is to provide a continuous or discontinuous physiological parameter analysis system with high accuracy and simple system structure.

In order to solve the existing technical problems, the technical solutions adopted by the present invention are:

A continuous or discontinuous physiological parameter analysis system, including a sensor electrode implanted in the subcutaneous tissue of a living body, a connection device connecting the sensor electrode and a voltage source, an internal transmitter and an external processor. The internal transmitter includes a voltage source, Standard circuit, measurement circuit and transmission circuit. The standard circuit is connected to the voltage source in parallel with the sensor electrode. The external processor includes a processing module and a transmission and reception module. The sensor electrode includes at least one working electrode and one parameter. Specific electrode, 1 sensing area, the length of the sensing area is (K-2) × cscα, k is 2.5-6mm, α is the angle at which the sensor electrode is implanted under the skin of the organism, and the range is 0-60°; The resistance value of the connecting device is at least 2 orders of magnitude smaller than the electrode impedance _Re of the sensor electrode; the resistance of the standard circuit is a multiple of one hundredth to one ten thousandth of the electrode impedance _Re of the sensor electrode, and the The continuous voltage applied by the voltage source can cause at least one electrochemical reaction of one or more chemicals in the tissue fluid of the organism.

A further improvement is that the electrode carrier of the sensor electrode is a cylindrical metal wire, and the implanted end of the cylindrical metal wire is provided with a protective layer with a length of 0.3-1 mm. The material of the protective layer is parylene, poly Imide or expanded polytetrafluoroethylene insulation material.

Improvedly, the reference electrode material of the sensor electrode is a mixed slurry of silver and silver chloride, and the mass ratio of silver:silver chloride is 50:50-70:30.

An improvement is that the electrode carrier of the sensor electrode is a polymer sheet, the reference electrode and the working electrode on the polymer sheet are stacked in opposite directions, and the reference electrode is located in the outermost layer or the innermost layer.

Improvedly, the sensor electrode is provided with a polymer film layer outside.

A further improvement is that the polymer film layer is 3-10 μm in diameter, and the hydrogel formation time of the polymer film layer is at least 3 minutes.

A further improvement is that the electrode impedance _Re of the sensor electrode is not greater than 7×10 ⁹ Ω, and the ratio to the voltage V of the voltage source is:

When the detection time is 3 days, the sensor electrode sensitivity is not greater than 100nA/mm, and R _e /V ≥ 10 ⁷ ;

When the detection time is 7 days, the sensor sensitivity is not greater than 50nA/mm, and R _e /V ≥ 2×10 ⁷ ;

When the detection time is 14 days, the sensor sensitivity is not greater than 25nA/mm, and _Re /V≥4×10 ⁷ .

Improvedly, the connecting member of the connecting device is conductive rubber, conductive glue, metal spring or metal clip.

Further improvement is that the conductive rubber has a high degree of compression of 12-30%, a Shore A hardness of 60-80 degrees, a tensile strength of greater than 1.3MPa, and a compression permanent deformation of less than 30 (70 h × 100°C) /%, the conductivity is less than 0.1Ω·cm.

A further improvement is that the working electrode and the reference electrode of the sensor electrode are connected to the conductive rubber by the electrodes piercing the conductive rubber or the electrodes and the conductive rubber being extruded.

Improvedly, the electrode impedance R _e of the sensor electrode includes the surface electron transfer resistance R _et of the working electrode and the reference electrode and the electrolyte resistance R _S .

Improvedly, the external processor also includes an acceleration detection module, a heart rate monitoring module or a calibration input module.

Improvedly, the voltage source generates pulse voltages or step voltages.

A further improvement is that the voltage source applies N consecutive voltages, the first voltage value is the smallest, the Nth voltage value is smaller than the N-1th voltage, and the two voltages are in the same direction.

A further improvement is that the voltage source applies three continuous voltages, the second voltage and the second voltage have the same direction, and the second voltage value is greater than the first voltage value and the third voltage value.

A further improvement is that the first voltage at least causes the analyte to produce an electrochemical reaction, the second voltage not only causes the analyte to produce an electrochemical reaction, but also causes the interfering substance to produce an electrochemical reaction, and the third voltage causes the analyte to produce an electrochemical reaction. substances produce electrochemical reactions.

A further improvement is that the first voltage value V ₁ is 0.2-0.7V, lasting 5-10min; the second voltage value V ₂ is 0.7-1.2V, lasting 3-10min; the third voltage value V ₃ is 0.2-0.7V.

beneficial effects

The physiological parameter analysis system of the present invention has a simple structure and is easy to operate, can realize continuous or discontinuous data monitoring, improves the accuracy of monitoring data, and provides reliable data support for patient treatment.

Description of the drawings

Figure 1 is the impact of the interfering substance ascorbic acid on the test results of an example current-type glucose sensor designed in the present invention;

Figure 2 is an example of sensor electrode implantation for monitoring glucose in subcutaneous tissue fluid;

Figure 3 is an embodiment of a cylindrical sensor electrode. (a) is the longitudinal interface diagram of the sensor electrode, (b) is the transverse interface diagram of the sensing area of the sensor electrode, 31 is the conductive layer, 28 is the inner membrane layer, 29 is Enzyme layer, 30 is the polymer film layer, (c) is the lateral interface diagram of the sensor reference electrode area, 31 is the conductive layer, 32 is the insulating layer, 33 is the silver/silver chloride layer, 30 is the polymer film layer;

Figure 4 shows the planar shape of the sensor electrode, in which 1 is the reference electrode, 2 is the working electrode, 3 is the blank electrode, 4 is the counter electrode, 5 is the first conductive contact point, 6 is the second conductive contact point, and 7 is the third Conductive contact point, 8 is the fourth conductive contact point;

Figure 5 is a schematic diagram of a single-pin cylindrical sensor electrode passing through conductive rubber, where 1 is the reference electrode, 2 is the working electrode, 9 is the first conductive rubber, and 10 is the second conductive rubber;

Figure 6 is a schematic diagram of a single-pin cylindrical sensor electrode connected through a conductive metal clip, in which 1 is the reference electrode, 2 is the working electrode, 11 is the first metal clip, and 12 is the second metal clip;

Figure 7 is a schematic diagram of a single-pin planar sensor electrode connected through a conductive metal spring. (a) is a top view, 1 is the reference electrode, 2 is the working electrode, 3 is the blank electrode, 5 is the first conductive contact point, and 6 is The second conductive contact point, 7 is the third conductive contact point, 8 is the fourth conductive contact point, (b) is the side view, 13 is the first spring coil, 14 is the second spring coil, 15 is the third spring coil, 16 is the fourth spring coil;

Figure 8 is a single-pin planar shape sensor electrode conductive rubber implementation. (a) is a top view, 1 is the reference electrode, 2 is the working electrode, 3 is the blank electrode, (b) is a side view, 17 is the third conductive rubber, 18 is the fourth conductive rubber;

Figure 9 is a schematic diagram of the fixed connection of multi-pin electrodes with conductive glue. (a) Top view, 1 is the reference electrode, 2 is the working electrode, 3 is the blank electrode, 5 is the first conductive contact point, and 6 is the second conductive contact point. , 7 is the third conductive contact point, (b) side view, 19 is the thimble, 27 is the wire;

Figure 10 is the Randles equivalent circuit diagram composed of the resistance of the two-electrode sensor system, in which 20 is the resistance R, 21 is the electrolyte resistance, 22 is the electron transfer resistance of the electrode surface, 23 is the Warburg impedance, and 24 is the resistance of the electrode surface. Electric double layer capacitor;

Figure 11 is the standard circuit equivalent schematic diagram of the dual-electrode sensor system. 25 is the standard circuit resistance, 26 is the voltmeter, 28 is the first ammeter, and 29 is the second ammeter;

Figure 12 is the impedance schematic diagram of the current sensor Randles equivalent circuit;

Figure 13 shows the impedance diagram of dual electrodes in an in vitro simulated human tissue fluid environment;

Figure 14 illustrates a voltage application method of a detection system for continuously analyzing glucose concentration in interstitial fluid in one example of the present invention;

Figure 15 is a graph illustrating the in vitro test current curves of multiple implantable glucose sensors prepared by an example solution of the present invention under a series of glucose concentrations;

Figure 16 is a graph illustrating the 11-day test results of a glucose sensor prepared by an example solution of the present invention and implanted in the subcutaneous tissue of a living body. During the test process, it was required to record daily events such as calibration values and diet;

Figure 17 is a graph illustrating a comparison of biochemical blood test results between a glucose sensor prepared by an example solution of the present invention and implanted in the subcutaneous tissue of the body within a certain period of time. During the test process, it is required to record a test across diet, Biochemical analysis process in stationary steady state;

Figure 18 is a schematic diagram of the analysis system of the present invention.

Detailed ways

Example 1

1. Sensor electrode

The sensor electrode of the present invention provides the original signal for measurement. The electrode includes but is not limited to a biological enzyme electrode, a chemical electrode, a physical electrode, an electrochemical electrode, a spectrophotometer electrode, an electrode for measuring polarization, an electrode for measuring heat, an electrode for measuring radiation, Immunochemical electrodes, or similar electrodes.

A single-needle sensor means that the working electrode and the reference electrode are on the same substrate carrier and are separated from each other by an insulating layer.

A multi-pin sensor means that there is only one electrode on each substrate carrier, and the working electrode and the reference electrode are separated from each other through different substrate carriers.

Whether it is a multi-pin sensor or a single-pin sensor, during the preparation process, due to inconsistent control of environmental conditions such as temperature and humidity, especially traditional manual manufacturing, the impact on sensor performance is slightly different.

The working principle of current implantable sensors commonly used to measure glucose is as follows:

Glucose oxidase is used to catalyze glucose to generate gluconic acid and hydrogen peroxide. Under the action of platinum as a catalyst, the hydrogen peroxide undergoes an electrochemical reaction to produce an electrochemical current signal, which can be measured by electronic components. Although glucose oxidase is specific for glucose, this specificity is reduced by poor choice of electrodes. For example, at +650 mV, not only hydrogen peroxide responds with current, but also some other point-active interferents in body fluids. Electric current, these electroactive species interfere with the function of the sensor through oxidation of the electrodes or reaction with hydrogen peroxide.

Chinese Patent No. CN200910097842 discloses a needle-shaped amperometric glucose sensor for real-time monitoring of subcutaneous tissue and a manufacturing method thereof. The glucose sensor of the invention is particularly suitable for detecting glucose in subcutaneous tissue fluid. The glucose sensor is a sandwich sensor that uses metal as the conductive layer, platinum black as the catalytic layer, and polymer siloxane as the inner membrane layer. The inner membrane layer polymer material eliminates the influence of interfering substances, such as ascorbic acid and amino acid uric acid. , bilirubin, catecholamines, drugs and various other metabolites. The high molecular weight polymer serves as the outer membrane layer, which is used to isolate macromolecules and cells.

Those skilled in the art are familiar with a variety of sensor inner and outer membrane preparation technologies, and film formation can be achieved through solution technologies such as spraying, dipping, casting, spin coating, coating and other similar technologies. This technology is achieved by evaporating volatile liquids, such as water and organic solvents, into polymer films that remain behind. The evaporation method can be heat, high-energy radiation, ultraviolet light or negative pressure.

The internal and external polymer membranes achieved by this technology are very dry after the solution evaporates, which helps protect the activity of the enzyme. However, in the absence of infiltration, the pore size of a dry polymer membrane is large enough that small molecules such as ascorbic acid, amino acid uric acid, bilirubin, catecholamines, drugs, and various other metabolites can easily enter The sensor catalytic layer produces an electrochemical reaction at a certain potential to form an interference signal. The sensor signal measured at this time not only contains the electrochemical signal generated by glucose, but also contains the signal generated by interfering substances, which has a great impact on the test results. Figure 1 illustrates the impact of the interfering substance ascorbic acid on test results.

When the sensor is implanted into the biological tissue for a certain period of time, under the infiltration of tissue fluid, the polymer membrane gradually expands to form a hydrogel. The pore size of the membrane gradually decreases, and a stable spatial three-dimensional structure is formed. At this time, small molecule interference substances It is not easy to enter the sensor electrode. The time it takes for the polymer film to gradually expand to a stable hydrogel becomes longer as the thickness of the film increases. After research, it was confirmed that when the thickness of the polymer film is 3 to 10 μm, it takes no less than 3 minutes to form a stable hydrogel. In the process of achieving the stability of the polymer film hydrogel, body fluids continuously diffuse through the film layer into the conductive layer of the sensor. In addition to some macromolecules and proteins, glucose, oxygen, ions and small molecule interference substances also enter the enzyme layer and The catalytic layer generates electrochemical interference signals under the action of a certain potential and platinum and other catalysts. When the polymer film forms a stable hydrogel, the outer hydrogel film blocks most macromolecules, proteins and some small molecule interferences, and the inner hydrogel film blocks most small molecule interferences from entering. to the catalytic layer. Therefore, when the polymer forms a stable hydrogel, the small molecule interference signal is greatly reduced.

The material of the polymer film layer in the present invention is polytetrafluoroethylene, polyolefin, polyamide, polydimethylsiloxane, polyurethane, polycarbonate, polyurea cellulose acetate, Nafion, polyester sulfonic acid, etc. Hydrophilic membranes prepared from such materials through solution evaporation technology.

In addition, in terms of electrode structure, the sensor electrode of the present invention at least includes a working electrode and a reference electrode. When the sensor electrode is a single-pin sensor, the working electrode and the reference electrode can be realized by using a metal wire as a carrier, such as a platinum wire carrier, and separating multiple electrodes with an insulating layer. When the sensor electrode is a multi-pin sensor, it can be realized by using multiple metal wires such as platinum-coated stainless steel wires as carriers, and each separate metal wire as an electrode. The working electrode and reference electrode can also be coated with a conductive layer on a flat polymer material or ceramic material as the carrier of the electrode. The conductive layer coating method includes electroplating, vapor deposition, spray coating, dip coating and other methods.

The working electrode of the sensor of the present invention is designed to have at least one sensing area, in which the analyte can generate an original signal such as an optical signal, an electrical signal or other raw signals that can be detected by electronic components. As shown in Figure 2, the sensing area A (enzyme area) of the sensor electrode used to detect glucose in subcutaneous tissue fluid is required to be implanted into the subcutaneous tissue. Therefore, the sensing area is required to neither stay in the dermis nor Implanted into muscle. In this example, whether the sensor electrode is a cylindrical electrode similar to a metal wire as a carrier, a planar electrode with a polymer sheet substrate, or an electrode of other shapes, the maximum length L of the implanted sensing area is required. The value can be calculated using the following formula:

L=（K-2）×cscα

The length L1 of the sensor electrode required to be implanted can be calculated by the following formula:

L1=K×cscα

Among them, α is the angle at which the sensor electrode is implanted, ranging from 0 to 60 degrees, and the K value ranges from 2.5 to 6mm. It can be understood that L and L1 are required to be calculated from the implantation end.

Or as shown in Figure 3, a cylindrical sensor electrode for detecting glucose in subcutaneous tissue fluid. Figure 3A is a longitudinal cross-sectional view of the sensor electrode. Among them, metal wire A is the electrode carrier. The metal wire can be platinum wire, platinum-iridium wire, platinum-coated stainless steel wire, nickel-titanium wire, silver wire, titanium wire, tantalum wire, or palladium wire with good biocompatibility. , gold wire and its mixed metal wires. C is the reference electrode, and D is the polymer film hydrogel layer. The end close to the enzyme reaction area (A1) is the subcutaneous implant end. The working electrode sensing area is designed as an enzyme reaction area, and its length L ranges from 0.5mm to 5mm. The other end A2 is the circuit connection end, and B2 is the insulation layer. The circuit connection end A2 cannot be completely covered by the insulation layer. A length of exposed metal wire of L3 is reserved to connect to the measurement circuit. The length of L3 is determined by the contact size of the measurement circuit. Decide.

The protective layer B1 is the implanted protective end, with a length of 0.3mm to 1mm. The insulating layer B2 separates the reference electrode and the working electrode. The protective layer B1 and the insulating layer B2 can be the same polymer material with good biocompatibility, or they can be different polymer materials with good biocompatibility. The materials of the insulating layer are polyparaxylene, polyimide, expanded polytetrafluoroethylene, etc. It is coated on the metal wire carrier using vapor deposition, spray coating, dip coating and other methods.

The reference electrode C may completely cover the insulating layer B2, or may not completely cover it. The reference electrode C is usually a silver/silver chloride electrode. It can be made by winding silver wire around the insulating layer B2 and then chlorinating it in a solution containing chloride ions. The silver/silver chloride paste can also be coated on the insulating layer B2, and the coating method can be dipping, spraying, deposition, etc. The use of coated silver/silver chloride slurry requires the mass ratio of silver:silver chloride to be in the range of 50:50 to 70:30.

D is a polymer film layer that continuously covers the protective end B1, the reaction area A1, the insulating layer B2 close to the implanted end, and the reference electrode C. Its coverage length L2 value is at least greater than the length L1 of the electrode implanted in the subcutaneous tissue, but it cannot completely cover the reference electrode C near terminal A2. A silver/silver chloride layer of length L4 should be reserved to connect to the measurement circuit. , the size of L4 is determined by the contacts of the measuring circuit.

B1 is a transverse cross-sectional view of the sensing area of the sensor electrode. The design 1 is a conductive catalytic layer, which can be a platinum wire, a platinum-iridium wire, a platinum-coated metal wire or a metal alloy wire with good biocompatibility; the design 2 is the inner membrane layer, which can be organic silane or a polymer containing silane components; the design 3 is an enzyme layer, which is glucose-specific glucose oxidase; the design 4 is a hydrogel polymer outer membrane layer, It can be polytetrafluoroethylene, polyolefin, polyamide, polydimethylsiloxane, polyurethane, polycarbonate, polyurea cellulose acetate, Nafion, polyester sulfonic acid and other materials.

As shown in Figure 4, a sensor electrode used to detect the analyte, which is glucose in subcutaneous tissue fluid, includes reference electrode 1, working electrode 2, blank electrode 3, counter electrode 4, and electrodes in order from bottom to top. The whole is covered with a polymer outer film layer. The reference electrode and the working electrode are stacked in opposite directions to ensure maximum utilization of the reference electrode. Area A is the sensing area, which is covered by glucose-specific glucose oxidase. 5 is a first conductive contact point, 6 is a second conductive contact point, 7 is a third conductive contact point, and 8 is a fourth conductive contact point. Those skilled in the art will understand that stacking can also be performed in different orders. Each electrode includes a base material sheet, which is a film sheet made of a water-impermeable insulating material. It can usually be a ceramic substrate (such as alumina) or a polymer substrate (such as polyimide). The electrocatalytic layer of the sensor electrode in this example usually refers to the ability to measure the signal of the analyte and transmit the signal to the measurement circuit, and at the same time catalyze the electrochemical reaction of the analyte or the product after the catalytic reaction of the analyte. The function is usually to coat platinum or palladium on the metal layer or the carbon layer. The metal layer is deposited on the substrate sheet and etched into a certain geometric shape. The deposition method can be magnetron sputtering, electroplating, vapor deposition, printing and other methods.

Each electrode is stacked by gluing a silane compound such as aminopropyltrimethoxysilane. The polymer film layer is made of polytetrafluoroethylene, polyolefin, polyamide, polydimethylsiloxane, polyurethane, polycarbonate, polyurea cellulose acetate, Nafion, polyester sulfonic acid and other materials.

The electrode impedance, including electron migration impedance, loop impedance, electrolyte impedance, etc., is simulated in the Randles equivalent circuit to simulate the sensor impedance, and the electron migration impedance occupies a dominant position. The electrochemical current signal of the sensor is mainly determined by the electron migration resistance. Electron migration impedance is determined by the nature of the substrate material participating in the electrochemical reaction, substrate concentration, sensor electrode surface, applied potential, etc.

The total amount of electrons released by the reference electrode (usually a silver/silver chloride electrode) in the sensor electrode of the present invention should be able to maintain the entire electrochemical reaction at a certain potential. Therefore, it is required that the impedance of the sensor electrode cannot be too small, so that the reference electrode will not fail quickly.

When the detection time is 3 days, the current sensitivity of the electrochemical reaction is not greater than 100nA/mm, and the corresponding electrode impedance is 2.0 MΩ ~ 7.0 MΩ. When the detection time is 7 days, the current sensitivity of the electrochemical reaction is not greater than 50nA/mm, and the corresponding electrode impedance is 4MΩ~14MΩ. When the detection time is 14 days, the current sensitivity of the electrochemical reaction is not greater than 25nA/mm, and the corresponding electrode impedance is 5MΩ~28MΩ.

At the same time, in the detection of physiological parameters of biological tissue fluids, especially in glucose current-type sensors, the sampling accuracy requirements of the measurement circuit cannot be too high, so as to avoid the equipment being too large and the cost being too high. Typically no higher than 0.1nA. Under this accuracy requirement, the impedance of the sensor is not greater than 7000 MΩ.

2. Connect the circuit

Whether it is the production of needle-shaped metal-based sensor electrodes disclosed in Chinese patent number CN200910097842 or the preparation of polymer-based sheet biosensor electrodes disclosed in US application number US20130126349, there is no detailed explanation of the sensor electrodes, measurement circuits and power supplies. Connection methods and their impact on sensor performance. However, in practical applications, it is necessary to maintain the activity of the enzyme and the stability of the sensor polymer membrane material to ensure the detection effect. Conventional thermal welding, ultrasonic welding, laser welding and other methods cannot be used.

As shown in FIG. 5 , the connecting member is a first conductive rubber 9 and a second conductive rubber 10 . Conductive rubber is used to connect the electrodes to the measurement circuit and power supply, and the elasticity of the conductive rubber is used to fix the electrodes. Use an auxiliary tool to pierce the electrode through the middle of the conductive rubber. The contacts of the measuring circuit (omitted in the figure) exert a certain pressure on the conductive rubber. The conductive rubber deforms and the electrode is in close contact with the conductive rubber. Alternatively, the connection can be achieved by pressing the conductive rubber and pressing the electrode directly on the surface of the conductive rubber.

As shown in Figure 6, a single-pin cylindrical sensor electrode is clamped by a first metal clip 11 and a second metal clip 12 to connect the electrode to the measurement circuit and power supply. Or you can use a piece of metal and apply pressure to the contact point of the measuring circuit so that the metal piece is in close contact with the electrode, and the connection can also be completed.

As shown in Figure 7, an example of a single-pin planar sensor electrode has four electrodes (a reference electrode is not marked on the back of the picture). All electrodes are connected to their corresponding conductive contact points. One end of the spring coil is connected to the contact of the measurement circuit. Click to connect. The other end of the spring coil presses against the electrode contact. By pressing the first spring coil 13, the second spring coil 14, the third spring coil 15 and the fourth spring coil 16 under the pressure of the measuring circuit contacts, the first spring coil 13, the second spring coil 14, the third spring coil 15 and the fourth spring coil 16 are deformed so that they are in close contact with the conductive contact points respectively, thereby realizing the connection between the electrodes, the measuring circuit and the power supply. Good connection. The spring can be changed into a conductive metal such as an ejector pin, a metal sheet, or a metal column. If the spring is changed to conductive rubber, the single-pin planar shape sensor is as shown in Figure 8 (third conductive rubber 17, fourth conductive rubber 18).

As shown in Figure 9, the reference electrode 1 is connected to the wire 27 through the conductive contact point 5, and the ejector pin 19 presses the wire. The wires are connected to the electrodes and fixed on the contact plate using conductive glue solidification. Conductive glue not only plays a role in connection, but also provides a role in fixing the electrodes.

The electrodes can be double electrodes or multiple electrodes. Regardless of whether the conductive rubber clamping method is used, the metal clamping method such as springs, metal sheets, or the conductive glue fixation method is used, the resistance between the electrode and the contact point of the measurement circuit after the clamping connection and the stability of the electrode need to be considered fixed. The unstable fixation of the electrode not only affects the use of the electrode, but also affects the resistance between the electrode and the measuring circuit contacts.

The following gives the Randles equivalent circuit principle of a two-electrode sensor:

As shown in Figure 10, the entire two-electrode electrochemical system can be simulated with the Randles equivalent circuit, where R _et is the electron transfer resistance of the working electrode surface, Z _w is the Warburg impedance, C _dl is the double layer capacitance of the electrode surface, R _S is the electrolyte resistance, and R is the equivalent resistance of the loop including the wire connection resistance. In the case of wires only, the wire connection resistance R is ignored. However, as mentioned before, in the preparation technology of implantable biosensor electrodes, the sensor electrodes must be connected to the contacts of the measurement circuit through a connecting device, which can be through conductive rubber or metal clips (springs) , conductive glue and other similar methods to achieve. The equivalent circuit resistance R is mainly determined by the resistance between the electrode and the contact point of the measuring circuit. In particular, when using conductive glue or conductive rubber, the connection resistance may be very high due to special conditions, such as insufficient contact pressure on the wire rubber, too much resin content in the conductive glue, etc. Then if the loop resistance R value in Randles' equivalent circuit is not different from the sum of R _et and R _S by an order of magnitude, the R value cannot be ignored. The result is that in the series circuit in which the entire sensor works, the connecting resistor shares part of the voltage given by the power supply. The working voltage of the sensor electrode is not the theoretical voltage given in the design process, resulting in a large difference between the test results and the actual results.

The inventor's research has confirmed that the resistance of the connection device and the impedance between the sensor electrodes (including the working electrode surface electron transfer resistance R _et and the electrolyte resistance R _S ) are at least 2 orders of magnitude smaller or more, so that the resistance of the connection device will not affect the test. The results have an impact. Therefore, when designing a sensor, especially when designing an implantable glucose sensor, the connection method of the sensor electrodes to the contacts of the measurement circuit must be taken into account, such as the resistance of the selected material of the connection device, and the method of connection.

The conductive rubber can be one of nickel-copper silicone rubber, silver-nickel silicone rubber, silver-aluminum silicone rubber, silver-glass silicone rubber, silver-coated copper silicone rubber, and carbon black silicone rubber. The shape can be a long strip or a cylinder, or other geometric shapes.

In the connection method shown in Figure 7 and Figure 8, the sensor electrode uses an insulating polymer sheet as a carrier for a planar electrode sensor. A layer of metal such as gold or platinum is deposited on the insulating polymer sheet substrate through electroplating, sputtering, etc. Conductive catalytic layer. The connection to the measuring circuit is through contact with the electrode contacts through thimbles, springs, conductive rubber, etc. In this connection method, the connection resistance will also be too high due to reasons such as the density of the deposited metal layer such as gold or platinum is not dense and the bonding force is insufficient.

3. Internal transmitter

The internal transmitter is used to store the measurement data, analyze the measurement data, correct the measurement data and transmit the data. The current and voltage measured in the measurement circuit are continuous data, so the processing circuit of the internal transmitter is required to average the collected data.

The internal transmitter includes the voltage source, standard circuit, measurement circuit and transmission circuit.

1.Voltage source

When the thickness of the polymer film outside the sensor electrode is 3 to 10 μm, it takes no less than 3 minutes to form a stable hydrogel. Before the polymer film layer is stabilized, body fluids continue to diffuse through the film layer into the conductive layer of the sensor. In addition to some macromolecules and proteins, glucose, oxygen, ions and small molecule interference substances also enter the enzyme layer in the sensing area, and Produce electrochemical reactions and form interference signals. The following describes a method for eliminating the influence of the above-mentioned small molecule interfering substances by applying a continuously changing voltage.

In electrochemical current-type sensors, both the analyte and the interfering substance may produce electrochemical reactions. Generally, the analyte and the interfering substance are collectively called substrates, and the oxidation potential and reduction potential of each substrate are different. The voltage source is designed to produce a continuous pulse voltage or step voltage that changes in the same direction or in different directions. The designed voltage application method is used to cause an electrochemical reaction of the analyte and at the same time to completely or partially consume the small molecule interfering substances that diffuse into the sensor before forming a stable polymer film hydrogel. If N consecutive voltages are applied, the value of the first voltage is required to be the smallest, and at least the Nth voltage and the N-1th voltage are in the same direction, and the value of the Nth voltage is less than or equal to the N-1th voltage. voltage value. In addition, the voltage source of the present invention has at least three consecutive voltages, in which the first voltage acts on the time period before the polymer film hydrogel is stabilized, and the second voltage acts on the small molecules that are believed to be able to fully or partially diffuse into the sensor. The interference consumes time period, and the third voltage is used until the end of the detection period. The voltage in the second stage and the voltage in the third stage are required to be in the same direction. The voltage in the first stage and the voltage in the second stage can be in the same direction or in different directions. And at the same time, the value of the second stage voltage is required to be greater than the value of the first stage voltage and the third voltage. The applied voltage in the first stage causes at least an electrochemical reaction of the analyte. The applied second-stage voltage, in addition to the electrochemical reaction of the analyte, also causes the electrochemical reaction of the main interfering substances entering the enzyme layer and conductivity catalysis. The applied third-stage voltage mainly causes an electrochemical reaction of the analyte.

2.Standard circuit

The way the sensor electrodes are connected affects the test of the sensor. However, when the sensor is implanted into the body, the accuracy of the test results cannot be known until they are compared with the standard values. Moreover, even if the accuracy of the test results is doubted, it cannot be judged whether the error is caused by the sensor electrodes. The reason is also caused by the connection method of the sensor electrode, which causes confusion when analyzing the factors of test failure. Therefore, it is necessary to propose an early warning plan for test failures caused by sensor electrode connection problems.

The standard circuit is a fixed-value resistor, which is connected in parallel with the sensor electrode to the DC voltage source circuit. Figure 11 is the equivalent schematic diagram, and R _c is the resistance of the standard circuit. The voltage and current at both ends can be tested by the measuring circuit. The function of the standard resistor is to determine whether the resistance of the connecting device affects the test results of the sensor. The value of the standard resistor can be determined based on the impedance between the sensor electrodes.

The following describes a method by which a standard circuit can be used to determine whether the resistance of the connection between the electrodes is causing a failure of the test.

First determine the resistance between the sensor electrodes. The impedance R _e between the sensor electrodes is composed of the electrode surface electron transfer resistance R _et and the electrolyte resistance R _S. This step can determine the electrode surface electron transfer resistance R _et and the electrolyte resistance R _S through the impedance electrochemical method in an environment that simulates human tissue fluid in vitro. The impedance schematic diagram of Randles equivalent circuit is shown in Figure 12. The shape of the Nyquist diagram of this circuit is a connected semicircle and a diagonal line. The semicircle part measured in the high frequency area is controlled by the electron transfer process, and its diameter is equal to Ret value, the front part of the semicircle is the electrolyte resistance _RS , then the resistance _Re of the entire sensor when implanted in the living body can be considered to be approximately equal to the sum of the value of Ret and the value of _RS . Figure 13 shows the impedance diagram of a dual-electrode of the present invention in an in vitro simulated human tissue fluid environment. The electron transfer resistance R _et value of the working electrode surface of the sensor is 1850KΩ, the electrolyte resistance R _S value is about 50KΩ, and the value between the sensor electrodes is The impedance _Re is approximately 1800KΩ.

The second step is to set the standard fixed resistor resistance R _c . The resistance value of R _c ranges from one hundredth to one ten thousandth of the electrode resistance _Re . Considering power consumption, it cannot be a smaller multiple.

The third step is to determine whether the resistance R of the connecting device affects the test results of the sensor. As mentioned above, the inventor confirmed that the test results are closer to the true value when the impedance between the resistance of the connecting device and the sensor electrode is at least 2 orders of magnitude smaller or more. The current of a standard fixed resistor can be calculated as I _s =V/R _c . V is the voltage applied across the standard circuit and is also the operating voltage applied across the electrodes. The measuring circuit tests the current between the sensor electrodes as I _e . If I _s is greater than 101 times I _e , it means that the resistance R of the connecting device is 2 orders of magnitude greater than the impedance between the sensor electrodes (including the working electrode surface electron transfer resistance R _et and the electrolyte resistance R _S ). The resistance R of the connecting device If the resistance value is too large, it will affect the test results of the sensor. If I _s is less than 101 times I _e , it means that the resistance R of the connecting device is 2 orders of magnitude smaller than the impedance between the sensor electrodes (including the working electrode surface electron transfer resistance R _et and the electrolyte resistance R _S ). The R resistance value will not affect the test results of the sensor.

In addition, in the electrode structure design of Figure 5 and Figure 6, an example of the present invention, the electrode working electrode and the reference electrode are separated by an intermediate insulating layer. During the preparation process of the insulating layer or the reference preparation process , the insulating layer may cause defects such as exposure and dents. For example, uneven coating of the insulating polymer may cause the polymer to not completely cover the conductive layer, or the insulating polymer may fall off due to wear, etc., causing the working electrode and The reference electrode is directly connected and is not isolated. In this case, the working electrode and the reference electrode directly form a circuit, and there is no need for electron transfer through the medium of tissue fluid. The sensor electrode does not produce an electrochemical reaction, and the sensor fails. At this time, the resistance R _e of the sensor electrode is almost zero. At this time, I _s and I _e are nearly equal. In another example of the present invention, as shown in the electrode structure design in Figures 7 and 8, when the working electrode and the reference electrode are in the preparation process, causing the working electrode and the reference electrode to be directly connected, at this time I _s It is also nearly equal to I _e .

3.Measurement circuit

The measurement circuit is designed to be controlled by the internal transmitter processing circuit, including sampling modulation circuit, ADC conversion circuit, etc., and can collect the current data of the sensor and the current and voltage data of the standard circuit. The sampling modulation circuit is required to be designed to sample at a preset sampling frequency. The ADC conversion circuit is designed to convert the collected current and voltage data into digital signals after analog-to-digital conversion by the ADC circuit and send them to the internal transmitter processor for analysis. The measurement circuit also includes a digital filter. Usually, the digital filter is required to filter the collected data to smooth the original data collected.

4.Transmission circuit

The transmission circuit is controlled by the processor and is designed to transmit the data packets established by the processor to the external controller and establish unique communication with the external controller. The transmitted data includes the sensor identification, the unique identifier corresponding to the sensor and the external controller, the sensor raw data, and the data processed by the processor. The transmission circuit is also designed to accept instructions from the external controller.

4. External processor

The external processor is used to store the data transmitted from the internal transmitter, evaluate the analysis results, compare the results of the analyte in various time periods, give the change trend, and evaluate the analysis results. It can also be used to download the data of the analyte. Daily time input and abnormal warning, etc.

The external processor also includes an acceleration module. It can complete functions such as exercise monitoring, sleep quality, sedentary reminders, and fall alarms. The step counting information and speed information provided by exercise monitoring, combined with the ECG/heart rate module, can provide an accurate quantitative reference for exercise intensity, and then guide users on how to control blood sugar through appropriate exercise. As for the fall alarm, because diabetic patients are more likely to suffer from syncope and fall caused by hypoglycemia, a remote alarm is issued for falls when no one is around (such as passing the fall alarm and GPS location through the phone function module). Sent to another relative) is necessary.

The external controller processor circuit also includes a heart rate monitoring module. The heart rate module mainly refers to the dynamic heart rate, which is monitored using the photoelectric principle. Together with the motion information provided by the acceleration module, it can provide a reference for the user's exercise duration and intensity. The amount of exercise will directly affect the changes in blood sugar. Users can predict the amount and intensity of exercise needed next time through the effect of the duration and intensity of exercise on blood sugar, and can more accurately know blood sugar control. The electrocardiogram module is an ECG circuit, which mainly reads electrocardiograms. Some data show that even if diabetic patients have no symptoms of heart disease, their electrocardiograms still have abnormal changes to varying degrees. Deaths caused by diabetic heart disease account for about 70% of diabetes mortality, among which arrhythmia is one of the important causes of death. Therefore, it is also very meaningful for diabetic patients to monitor dynamic electrocardiogram. An ECG sensor electrode is installed on the outside of the processor. When performing glucose detection, the ECG sensor electrode is started to work and can directly monitor the heart rate changes of the person being analyzed.

The external processor also includes a calibration input module. Test strips are mainly used to collect finger blood glucose concentration readings. This reading is used directly to calibrate the continuous glucose monitoring algorithm. When the test strip is inserted, the watch displays the blood glucose test strip reading interface. At this time, blood can be collected to test the finger blood concentration. After pressing Confirm to save, the processor will directly save the value as a reference blood sugar for algorithm calibration.

Example 2

In this example, the electrode carrier, that is, the conductive catalytic layer, is platinum wire. After the platinum wire is ionically activated, a layer of approximately 8 micron parylene is deposited by vapor deposition as an insulating layer for the working electrode and reference electrode. Using dip coating, a silver/silver chloride slurry with a ratio of 50:50 is applied on the surface of the insulating layer, and is cured and dried by ultraviolet light. A 2.5mm area is etched with a laser to expose the platinum wire as the sensing area, where the length of the protective end is 0.75mm. The glucose oxidase solution containing silane is coated on the sensing area, cross-linked with a cross-linking agent, and dried and solidified to form a 2.5 mm long glucose sensing area. A layer of polyurethane hydrogel is coated to isolate external macromolecular interference. The hydrogel is also used to limit the diffusion of glucose and oxygen in the interstitial fluid.

The sensor electrode is fixed on the electrode fixing piece of the sensor base through conductive rubber. The electrode passes through the middle of the conductive rubber, and the electrode forms an angle of 45 degrees with the horizontal. The sensor electrode is implanted into the subcutaneous tissue of the organism with the help of an auxiliary puncture tool, and the implantation length is 8mm. The thimble of the measuring circuit of the internal transmitter presses the conductive rubber tightly, and the resistance between the electrode, conductive rubber and the measuring circuit is 8.31KΩ ohms. The impedance between the electrodes is 1850KΩ and the standard circuit has a resistance of 10KΩ.

The voltage source of the internal transmitter applies the first-stage voltage V1 to a value of 0.60V for 9 minutes. At the 3rd minute of applying the voltage, find out if the ratio of I _s to I _e is 1056, which is much greater than 101. This shows that the resistance of the connection method using conductive rubber for the electrodes is much smaller than the impedance between the sensor electrodes, and this connection scheme is feasible. The applied second-stage voltage V2 is 1.07V, which is in the same direction as the first-stage voltage and lasts for 5 minutes. The applied third-stage voltage is 0.60V, which lasts until the end of the detection. The voltage application method is shown in Figure 14.

The current and voltage signals tested by the measurement circuit are transmitted to the external controller through the transmission circuit, and after data processing, the test results are displayed.

Example 3

The electrode of the analysis system for measuring glucose prepared in Example 2 was placed in a phosphate buffer containing glucose. The concentrations of glucose were 5 mmol/L, 10 mmol/L, 15 mmol/L, and 20 mmol/L.

As can be seen from Figure 15, the in vitro detection sensitivity of the analysis system of the present invention is 27 nA/mmol ~ 32 nA/mmol, and the linear correlation is 0.999.

Example 4

The electrode of the analysis system for measuring glucose in Example 2 was sterilized and implanted into the organism for 11 days, and the detection results were recorded.

As shown in Figure 16, the continuous curve represents the continuous test results of this embodiment, and the dots represent the biochemical blood test results. As shown in Figure 17, it is a comparison of the 6-hour curve of Example 2 of the present invention and the biochemical blood test results, where , the detection time span spans eating activities and resting activities, and the biochemical results are recorded every 15 minutes. The solid line is the result of Example 2, the dots are the results of the biochemical blood test, and the points of the biochemical blood test are simulated. The resulting dashed curve. After analyzing the two curves, the standard error for the entire test period was 11.2%, the standard error for the high blood sugar period was 13.7%, and the low blood sugar error was 12.5%.

Claims (12)

1. A continuous or discontinuous physiological parameter analysis system comprising a sensor electrode implanted in subcutaneous tissue of a living body, a connecting device for connecting the sensor electrode and a voltage source, an internal transmitter and an external processor, wherein the internal transmitter comprises the voltage source, a standard circuit, a measuring circuit and a transmission circuit, the standard circuit and the sensor electrode are connected to the voltage source in parallel, and the external processor comprises a processing module and a transmission receiving module, and is characterized in that the sensor electrode at least comprises 1 working electrode, 1 reference electrode and 1 sensing area, the length of the sensing area is (K-2) multiplied by csc alpha, K is 2.5-6mm, alpha is the angle of implantation of the sensor electrode into the subcutaneous tissue of the living body, and the range is 0-60 degrees; the resistance value of the connecting means is at least 2 orders of magnitude smaller than the electrode impedance Re of the sensor electrode; the resistance of the standard circuit is one percent to one ten thousandth times of the electrode impedance Re of the sensor electrode, the voltage source applies three continuous voltages, the second voltage and the third voltage are the same in direction, the second voltage is larger than the first voltage and the third voltage, the second voltage can cause at least one electrochemical reaction of chemical substances in one or more tissue fluids in the organism, the first voltage at least causes the analyte to generate the electrochemical reaction, the second voltage causes the interference to generate the electrochemical reaction besides the analyte, and the third voltage causes the analyte to generate the electrochemical reaction; the electrode resistance Re of the sensor electrode is not more than 7×10 ⁹ Omega, in particular the first voltage value V ₁ 0.2-0.7V for 5-10min; second voltage value V ₂ 0.7-1.2V for 3-10min; first, theThree voltage values V ₃ Is 0.2-0.7V, and the ratio of the voltage V to the voltage source is as follows: when the detection time is 3 days, the sensitivity of the sensor electrode is not more than 100nA/mm, and Re/V is more than or equal to 10 ⁷ The method comprises the steps of carrying out a first treatment on the surface of the When the detection time is 7 days, the sensitivity of the sensor is not more than 50nA/mm, and Re/V is more than or equal to 2 multiplied by 10 ⁷ The method comprises the steps of carrying out a first treatment on the surface of the When the detection time is 14 days, the sensitivity of the sensor is not more than 25nA/mm, and Re/V is not less than 4 multiplied by 10 ⁷ 。

2. The continuous or discontinuous physiological parameter analysis system according to claim 1, wherein the electrode carrier of the sensor electrode is a cylindrical metal wire, the implantation end of the cylindrical metal wire is provided with a protective layer with the length of 0.3-1mm, and the protective layer is made of parylene, polyimide or expanded polytetrafluoroethylene insulating material.

3. The continuous or discontinuous physiological parameter analysis system of claim 1, wherein the reference electrode material of the sensor electrode is a mixed slurry of silver and silver chloride, the silver: the mass ratio of the silver chloride is 50:50-70:30.

4. the system of claim 1, wherein the electrode carrier of the sensor electrode is a polymer sheet, and the reference electrode and the working electrode on the polymer sheet are stacked in opposite directions, and the reference electrode is located on the outermost layer or the innermost layer.

5. A continuous or discontinuous physiological parameter analysis system according to claim 1, wherein the sensor electrode is provided with a polymer membrane layer.

6. The continuous or discontinuous physiological parameter analysis system of claim 5, wherein the polymer film layer is 3-10 μm, and the polymer film layer has a hydrogel formation time of at least 3 minutes.

7. A continuous or discontinuous physiological parameter analysis system according to claim 1, wherein the connector of the connector device is a conductive rubber, a conductive gel, a metal spring or a metal clip.

8. The system according to claim 7, wherein the conductive rubber has a high compression of 12-30%, a shore a hardness of 60-80 degrees, a tensile strength of greater than 1.3MPa, a compression set of less than 30 (70 h x l00 ℃) and a conductivity of less than 0.1 Ω -cm.

9. The continuous or discontinuous physiological parameter analysis system of claim 7, wherein the working electrode and the reference electrode of the sensor electrode are connected to the conductive rubber by electrode piercing the conductive rubber or by electrode extrusion with the conductive rubber.

10. A continuous or discontinuous physiological parameter analysis system according to claim 1, wherein the electrode impedance Re of the sensor electrode comprises the surface electron transfer resistance Ret and the electrolyte resistance RS of the working electrode and the reference electrode.

11. The continuous or discontinuous physiological parameter analysis system of claim 1, wherein the external processor further comprises an acceleration detection module, a heart rate monitoring module, or a calibration input module.

12. A continuous or discontinuous physiological parameter analysis system according to claim 1, wherein the voltage source generates a pulsed or stepped voltage.