CN106725470B

CN106725470B - Continuous or discontinuous physiological parameter analysis system

Info

Publication number: CN106725470B
Application number: CN201611035407.3A
Authority: CN
Inventors: 程荣恩; 肖林春; 吴伟华
Original assignee: Nantong Jiunuo Medical Technology Co ltd
Current assignee: Nantong Jiunuo Medical Technology Co ltd
Priority date: 2016-11-22
Filing date: 2016-11-22
Publication date: 2023-12-19
Anticipated expiration: 2036-11-22
Also published as: CN106725470A

Abstract

The invention discloses a continuous or discontinuous physiological parameter analysis system. The system includes an implantable sensor electrode for sensing an analyte, an internal transmitter, and an external processor. The internal transmitter comprises a voltage source capable of providing a continuously variable DC voltage and a module for measuring, analyzing, processing and transmitting signals. The external processor can store, analyze, process the data and include other physiological parameter monitoring functions. A connection device for connecting the sensor electrode to a voltage source and a method for evaluating the resistance of the connection device are also provided. The applied continuous voltage can cause one or more electrochemical reactions of one or more chemicals in tissue fluids within the living being. The system can realize continuous or discontinuous monitoring of the physiological parameters of organisms, has accurate and convenient results and is suitable for industrialization.

Description

Continuous or discontinuous physiological parameter analysis system

Technical Field

The invention relates to the field of in-vivo physiological parameter analysis, in particular to a continuous or discontinuous physiological parameter analysis system.

Background

With the development of technology, the level of the medical detection field is improved, and the use of the biosensor is more and more widespread. Biosensors are devices for measuring components in a living body by converting chemical substances in the living body into detectable signals such as current signals, optical signals, thermal signals, etc. Existing biosensors are of various kinds, such as bio-enzyme electrodes, chemical electrodes, physical electrodes, electrochemical electrodes, spectrophotometer electrodes, electrodes for measuring polarization, electrodes for measuring heat, electrodes for measuring radiation, electrodes for immunochemistry, and the like.

Such as glucose monitoring in humans. Diabetes mellitus is commonly called as rich and honour disease, along with the improvement of living standard, the prevalence rate is also increasing, and at present, diabetics keep stable blood sugar by injecting insulin. However, insulin overdose results in 3 times the probability of hypoglycemia occurring as compared to conventional therapies, and although insulin injection is effective in reducing blood glucose levels, complications can occur. Therefore, the blood glucose level should be monitored multiple times before and after insulin injection by the patient. Accurate blood glucose monitoring is a key for effectively controlling blood glucose, is beneficial to evaluating the metabolic disturbance degree of diabetics, and makes a blood glucose reducing scheme, and simultaneously reflects the blood glucose reducing treatment effect and timely adjusts the treatment scheme.

The U.S. patent application number US20130126349 discloses a method for manufacturing a continuous biosensor, which provides a method for manufacturing an electrode of an electrochemical biosensor, comprising the construction of the electrode and the construction of a polymer membrane.

Chinese patent grant publication No. CN102469966B discloses a continuous analyte measurement system and a system and method for implanting them. The system consists of a base, a sensor, an electronic element and a processing unit.

The Chinese patent application No. CN201492421U discloses a dynamic blood glucose meter capable of realizing dynamic blood glucose detection. The device consists of a sensor, a transmitter and a receiving processor.

The existing sensors are used in a variety of ways to immobilize enzymes, such as in U.S. patent application No. US20050115832, U.S. 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. In addition, during the use process of the sensors, the phenomena of short circuit and open circuit occur due to temperature, humidity and assembly, so that monitoring fails, particularly in the early detection stage, the detection result is inaccurate due to interference caused by small molecular interferents, and the sensor electrode is possibly unstable, so that the improvement of a continuous or discontinuous physiological parameter analysis system is necessary for the reasons.

Disclosure of Invention

Aiming at the defects of the prior art, the invention aims to provide a continuous or discontinuous physiological parameter analysis system which has high accuracy and simple system structure.

In order to solve the problems in the prior art, the invention adopts the following technical scheme:

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, the external processor comprises a processing module and a transmission receiving module, 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 device is higher than the electrode impedance R of the sensor electrode _e At least 2 orders of magnitude smaller; the resistance of the standard circuit is the electrode resistance R of the sensor electrode _e The continuous voltage applied by the voltage source can cause at least one electrochemical reaction of one or more chemicals in tissue fluids within the living being.

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.

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

the electrode carrier of the sensor electrode is a polymer sheet, a reference electrode and a working electrode on the polymer sheet are stacked in the opposite direction, and the reference electrode is positioned on the outermost layer or the innermost layer.

The improvement is that a polymer film layer is arranged outside the sensor electrode.

Further improved is that the polymer film layer is 3-10 mu m, and the time for forming the hydrogel is at least 3min.

Further improved is the electrode impedance R of the sensor electrode _e Not more than 7X 10 ⁹ Omega, and the ratio of the omega to the voltage V of the voltage source is:

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

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

When the detection time is 14 days, the sensitivity of the sensor is not more than 25nA/mm, R _e /V≥4×10 ⁷ 。

The improvement is that the connecting piece of the connecting device is conductive rubber, a metal spring or a metal clip.

Further improved is that the high compression amount of the conductive rubber is 12-30%, the Shore A hardness is 60-80 degrees, the tensile strength is more than 1.3MPa, the compression set is less than 30 (70 h multiplied by l00 ℃) and the conductivity is less than 0.1 Ω & cm.

The working electrode and the reference electrode of the sensor electrode are connected with the conductive rubber in a way of penetrating the conductive rubber through the electrode or extruding the electrode and the conductive rubber.

The improvement is that the electrode impedance R of the sensor electrode _e Surface electron transfer resistor R comprising a working electrode and a reference electrode _et And electrolyte resistance R _S 。

The external processor further comprises an acceleration detection module, a heart rate monitoring module or a calibration input module.

The improvement is that the voltage source generates a pulse voltage or a step voltage.

Further improved is that the voltage source applies N continuous voltages, the first voltage value is the smallest, the Nth voltage value is smaller than the N-1 th voltage, and the two voltages are in the same direction.

It is a further improvement that the voltage source applies three consecutive voltages, the second voltage and the second voltage being in the same direction, and the second voltage being greater than the first voltage and the third voltage.

It is further improved that the first voltage causes at least an electrochemical reaction to occur with the analyte, the second voltage causes an electrochemical reaction to occur with the interferent in addition to the electrochemical reaction to occur with the analyte, and the third voltage causes an electrochemical reaction to occur with the analyte.

Further improved is that 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; third voltage value V ₃ 0.2-0.7V.

Advantageous effects

The physiological parameter analysis system provided by the invention has the advantages that the structure is simple, the operation is convenient, the continuous or discontinuous data monitoring can be realized, the accuracy of the monitored data is improved, and the reliable data support is provided for the treatment of patients.

Drawings

FIG. 1 is an illustration of the effect of interferent ascorbic acid on the results of an exemplary amperometric glucose sensor of the present design;

FIG. 2 is a diagram of an example of sensor electrode implantation for monitoring glucose in subcutaneous tissue fluid;

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

FIG. 4 shows a planar sensor electrode, wherein 1 is a reference electrode, 2 is a working electrode, 3 is a blank electrode, 4 is a counter electrode, 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;

FIG. 5 is a schematic diagram of a single needle cylindrical sensor electrode passing through conductive rubber, wherein 1 is a reference electrode, 2 is a working electrode, 9 is a first conductive rubber, and 10 is a second conductive rubber;

FIG. 6 is a schematic diagram of a single needle cylindrical sensor electrode connected by a metal conducting clip, wherein 1 is a reference electrode, 2 is a working electrode, 11 is a first metal clip, and 12 is a second metal clip;

FIG. 7 is a schematic diagram of a single needle planar sensor electrode connected by a conductive metal spring, wherein (a) is a top view, 1 is a reference electrode, 2 is a working electrode, 3 is a blank electrode, 5 is a first conductive contact point, 6 is a second conductive contact point, 7 is a third conductive contact point, 8 is a fourth conductive contact point, (b) is a side view, 13 is a first spring turn, 14 is a second spring turn, 15 is a third spring turn, 16 is a fourth spring turn;

FIG. 8 is a schematic diagram of a single needle planar shape sensor electrode conductive rubber embodiment, (a) top view, 1 reference electrode, 2 working electrode, 3 blank electrode, (b) side view, 17 third conductive rubber, 18 fourth conductive rubber;

FIG. 9 is a schematic diagram of a fixed connection of a multi-needle electrode conductive paste, wherein (a) a top view, 1 is a reference electrode, 2 is a working electrode, 3 is a blank electrode, 5 is a first conductive contact point, 6 is a second conductive contact point, 7 is a third conductive contact point, (b) a side view, 19 is a thimble, and 27 is a wire;

FIG. 10 is a diagram of an equivalent circuit of Randes of the resistive composition of a dual electrode sensor system, where 20 is the resistance R,21 is the electrolyte resistance, 22 is the electrode surface electron transfer resistance, 23 is the Wang Bang (Warburg) impedance, and 24 is the electrode surface double layer capacitance;

FIG. 11 is a standard circuit equivalent schematic diagram of a dual electrode sensor system, 25 being a standard circuit resistor, 26 being a voltmeter, 28 being a first ammeter, 29 being a second ammeter;

FIG. 12 is a schematic diagram of impedance of an equivalent circuit of a amperometric sensor Randes;

FIG. 13 is an impedance plot of a two-electrode in vitro simulated human interstitial fluid environment;

FIG. 14 is a schematic diagram illustrating a method of applying voltage for a glucose concentration detection system for continuous analysis of tissue fluid in one example of the invention;

FIG. 15 is a graph illustrating in vitro test currents for a plurality of implantable glucose sensors prepared according to an example embodiment of the present invention at a range of glucose concentrations;

FIG. 16 is a graph illustrating the results of a test of 11 days of implantation of a glucose sensor in subcutaneous tissue in a living being, during which routine events such as correction values, diet, etc. are required to be recorded, prepared in accordance with an example embodiment of the present invention;

FIG. 17 is a graph illustrating the results of a glucose sensor implanted in subcutaneous tissue in a living being versus biochemical blood test during a period of time, as prepared in accordance with an embodiment of the present invention, during which the test is required to be recorded during a biochemical analysis that spans the diet, stationary plateau;

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

Detailed Description

Example 1

1. Sensor electrode

The sensor electrodes of the present invention provide a raw signal for measurement, including but not limited to biological enzyme electrodes, chemical electrodes, physical electrodes, electrochemical electrodes, spectrophotometer electrodes, polarization determining electrodes, heat determining electrodes, radiation determining electrodes, immunochemical electrodes, or the like.

A single needle sensor refers to a working electrode and a reference electrode on the same substrate carrier, separated from each other by an insulating layer.

A multi-needle sensor refers to only one electrode per substrate carrier, the working electrode and the reference electrode being separated from each other by different substrate carriers.

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

The current implantable sensor commonly used for measuring glucose works as follows:

the glucose oxidase is used for catalyzing glucose to generate gluconic acid and hydrogen peroxide, and the hydrogen peroxide is subjected to electrochemical reaction under the action of platinum serving as a catalyst to generate an electrochemical current signal, and the current signal can be measured by electronic components. Although glucose oxidase is specific for glucose, this specificity is reduced by poor choice of electrode, such as at +650 mV voltage, not only is hydrogen peroxide responsive to current flow, but also other point-active interferents in body fluids that interfere with sensor function through oxidation of the electrode or reaction with hydrogen peroxide.

Chinese patent No. CN200910097842 discloses needle-like current measuring glucose sensor for subcutaneous tissue real-time monitoring and its manufacturing method, and the glucose sensor of the present invention is especially suitable for detecting glucose in subcutaneous tissue fluid. The glucose sensor is a sandwich type sensor, wherein metal is used as a conductive layer, platinum black is used as a catalytic layer, high molecular siloxane is used as an inner membrane layer, and the influence of interferents such as ascorbic acid, amino acid uric acid, bilirubin, catecholamine, medicines and various other metabolites is eliminated through the high molecular material of the inner membrane layer. The high molecular polymer serves as an outer membrane layer for isolating macromolecules and cells.

A variety of sensor inner and outer membrane preparation techniques are well known to those skilled in the art, and film formation may be accomplished by solution techniques such as spraying, dipping, casting, spin coating, and the like. This technique is achieved by evaporating a polymer film that remains after evaporation from a volatile liquid, such as water and organic solvents. The evaporation mode can be heat, high-energy radiation, ultraviolet light or negative pressure.

The inner and outer polymer films realized by the technology are very dry after the solution is evaporated, which is helpful for protecting the activity of enzyme. However, the pore size of the dried polymer membrane is large enough without infiltration that small molecules such as ascorbic acid, amino acid uric acid, bilirubin, catecholamine, drugs and various other metabolites are very likely to enter the sensor catalytic layer, and electrochemical reactions are generated at certain potentials to form interfering signals. The sensor signal measured at this time contains not only an electrochemical signal generated by glucose but also a signal generated by an interfering substance, and has a large influence on the test result. Figure 1 illustrates the effect of the interferent ascorbic acid on the test results.

After the sensor is implanted into organism tissue for a certain time, the polymer film is gradually expanded under the infiltration of tissue fluid to form hydrogel, the pore diameter of the film is gradually reduced, and a stable space three-dimensional structure is formed, so that small molecular interferents are not easy to enter the sensor electrode. The time for the polymer film to gradually expand to stabilize the hydrogel increases as the thickness of the film increases. Through the study, it was confirmed that when the thickness of the polymer film was 3 to 10 μm, it took not less than 3 minutes to form a stable hydrogel. In the process of stabilizing the polymer membrane hydrogel, body fluid continuously passes through the membrane layer to diffuse into the sensor conductive layer, and besides some macromolecules and proteins, glucose, oxygen, ions and small molecular interferents enter the enzyme layer and the catalytic layer together, so that electrochemical interference signals are generated under the action of certain potential and platinum and other catalysts. When the polymer film forms a stable hydrogel, the outer hydrogel film blocks most of the macromolecules, proteins, and a portion of the small molecule interferents, and the inner hydrogel film blocks most of the small molecule interferents from entering the catalytic layer. Thus, small molecule interfering signals are greatly reduced when the polymer forms a stable hydrogel.

The material of the polymer film layer is a hydrophilic film prepared from polytetrafluoroethylene, polyolefin, polyamide, polydimethylsiloxane, polyurethane, polycarbonate, polyurea cellulose acetate, nafion, polyester sulfonic acid and the like by a solution volatilization technology.

In addition, the electrode structure of the sensor at least comprises a working electrode and a reference electrode. When the sensor electrode is a single needle sensor, the working electrode and the reference electrode may be realized by a single wire as a carrier, such as a platinum wire carrier, with the electrodes separated by an insulating layer. When the sensor electrode is a multi-needle sensor, this may be achieved by using a plurality of wires, such as platinized stainless steel wires, as the carrier, each wire being individually separated as one electrode. The working electrode and the reference electrode can also be prepared by coating a conductive layer on a planar polymer material or a ceramic material to be used as a carrier of the electrode, wherein the coating mode of the conductive layer comprises electroplating, vapor deposition, spray coating, dip coating and the like.

In the sensor working electrode of the present invention, at least one sensing region is designed in which the analyte is capable of producing a primary signal such as an optical signal, an electrical signal, or other signal that can be detected by an electronic component. As shown in fig. 2, the sensing region a (enzyme region) of the sensor electrode for detecting glucose in subcutaneous tissue fluid is required to be implanted into subcutaneous tissue, and thus the sensing region is required to be neither to stay in dermis nor to be implanted into muscle. In this example, the maximum value of the sensing area length L required to be implanted, whether the sensor electrode is a cylindrical electrode like a wire-supported or a planar electrode with a polymer sheet substrate, or other shaped electrode, can be calculated by the following equation:

L=（K-2）×cscα

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

L1=K×cscα

wherein alpha is the implantation angle of the sensor electrode, the range is 0-60 degrees, and the range of K value is 2.5-6 mm. It will be appreciated that L and L1 are required to be calculated from the implant end.

Or as shown in fig. 3, a cylindrical sensor electrode for detecting glucose in subcutaneous tissue fluid. Fig. 3A is a longitudinal sectional view of the sensor electrode. The metal wire A is an electrode carrier, and can be a platinum wire, a platinum iridium wire, a stainless steel wire coated with platinum, a nickel titanium wire, a silver wire, a titanium wire, a tantalum wire, a palladium wire, a gold wire and a mixed metal wire thereof with good biocompatibility. C is a reference electrode and D is a polymer membrane hydrogel layer. The end close to the enzyme reaction area (A1) is a subcutaneous implantation end. The working electrode sensing region is designed as an enzyme reaction region with a length L in the range of 0.5mm to 5mm. The other end A2 is a circuit connecting end, B2 is an insulating layer, the circuit connecting end A2 cannot be completely covered by the insulating layer, a section of exposed metal wire with the length of L3 is reserved to be connected to the measuring circuit, and the length of L3 is determined by the size of a contact of the measuring circuit.

The protective layer B1 is an implantation protective end, and the length is 0.3 mm-1 mm. The insulating layer B2 separates the reference electrode from the working electrode, and the protective layer B1 and the insulating layer B2 can be made of the same high polymer material with good biocompatibility or different high polymer materials with good biocompatibility. The insulating layer is made of parylene, polyimide, expanded polytetrafluoroethylene and the like. The metal wire is coated on the metal wire carrier by adopting methods of vapor deposition, spray coating, dip coating and the like.

The reference electrode C may or may not be entirely covered with the insulating layer B2. The reference electrode C is typically a silver/silver chloride electrode. Silver wires can be wound on the insulating layer B2 and then chlorinated in a solution containing chloride ions. The silver/silver chloride paste may be coated on the insulating layer B2 by dip coating, spray coating, deposition, or the like. Silver is required for the use of silver/silver chloride coated slurries: the mass ratio of the silver chloride is 50: 50-70: within 30.

D is a polymer film layer, and the protection end B1, the reaction area A1, the insulating layer B2 near the implantation end and the reference electrode C are continuously covered. The covering length L2 is at least larger than the length L1 of the electrode implanted in subcutaneous tissue, but can not completely cover the reference electrode C close to the terminal A2, the silver/silver chloride layer with the reserved length L4 is connected to the measuring circuit, and the size of the L4 is determined by the contact point of the measuring circuit.

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

As shown in fig. 4, a sensor electrode for detecting glucose in subcutaneous tissue fluid as an analyte comprises a reference electrode 1, a working electrode 2, a blank electrode 3, a counter electrode 4, and a polymer outer membrane layer entirely covering the electrodes in this order from bottom to top. The reference electrode is stacked in the opposite direction to the working electrode to ensure that the reference electrode is maximally utilized, the A region being the sensing region, which is covered by glucose-specific glucose oxidase. 5 is a first conductive contact, 6 is a second conductive contact, 7 is a third conductive contact, and 8 is a fourth conductive contact. It will be appreciated by those skilled in the art that stacking may also be performed in a different order. Each electrode comprises a substrate sheet, which is a thin film sheet made of a water-impermeable insulating material. Typically a ceramic substrate (such as alumina) or a polymeric substrate (such as polyimide) or the like. The sensor electrode electrocatalytic layer in this example is generally referred to as a layer capable of measuring the signal of the analyte and transmitting the signal to a measurement circuit and simultaneously catalyzing the electrochemical reaction of the analyte or the product of the catalytic reaction of the analyte, typically by coating a metal layer or carbon layer with platinum or palladium, the metal layer being deposited on a substrate sheet and etched to a geometry, the deposition being by means of magnetron sputtering, electroplating, vapor deposition, printing, etc.

Each electrode is stacked by gluing, which is a silane compound such as aminopropyl trimethoxysilane. 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 comprises electromigration impedance, loop impedance, electrolyte impedance and the like, and is used for simulating sensor impedance in a Randes equivalent circuit, wherein the electromigration impedance is dominant. The electrochemical current signal of the sensor is mainly determined by the electromigration resistance. The electromigration resistance is determined by the nature of the substrate material involved in the electrochemical reaction, the substrate concentration, the sensor electrode surface, the magnitude of the applied potential, etc.

The total amount of electrons released by the reference electrode (typically a silver/silver chloride electrode) in the sensor electrode of the present invention should be such that the entire electrochemical reaction is maintained at a certain potential. It is therefore required that the impedance of the sensor electrode not be too small so that the reference electrode will fail quickly.

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

Meanwhile, in the physiological parameter detection of biological tissue fluid, particularly in a glucose amperometric sensor, the sampling precision requirement of a measuring circuit cannot be too high, so that the equipment is not too large in size and too high in cost. Typically not higher than 0.1nA. With this accuracy requirement, the impedance of the sensor is no greater than 7000M Ω.

2. Connection circuit

Whether the manufacture of the sensor electrode with the needle-shaped metal as the matrix is published in Chinese patent No. CN200910097842 or the manufacture of the sheet-shaped biosensor electrode with the polymer as the matrix is published in U.S. Pat. No. 20130126349, the influence of the connection method and the connection mode of the sensor electrode and the measuring circuit as well as the power supply on the performance of the sensor is not described in detail. However, in practical application, the activity of the enzyme and the stability of the polymer membrane material of the sensor must be maintained, so as to ensure the detection effect, and conventional methods such as thermal welding, ultrasonic welding, laser welding and the like cannot be adopted.

As shown in fig. 5, the connection member is a first conductive rubber 9 and a second conductive rubber 10. The electrode is connected with the measuring circuit and the power supply by adopting conductive rubber, and the electrode is fixed by utilizing the elasticity of the conductive rubber. The electrode is pierced from the middle of the conductive rubber by using an auxiliary tool, a contact (not shown) of the measuring circuit applies a certain pressure to the conductive rubber, the conductive rubber deforms, and the electrode is in close contact with the conductive rubber. Or the electrodes are directly pressed on the surface of the conductive rubber by pressing the conductive rubber to realize connection.

As shown in fig. 6, a single needle cylindrical sensor electrode is connected with a measuring circuit and a power supply by clamping the electrode by a first metal clamp 11 and a second metal clamp 12. Or a piece of metal is used to press at 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 FIG. 7, in the example of a single needle planar sensor electrode, four electrodes (one reference electrode on the opposite side is not labeled) are all in communication with their corresponding conductive contact points, and one end of the spring coil is in contact with the measuring circuit. The other end of the spring coil presses against the electrode contact. The first spring ring 13, the second spring ring 14, the third spring ring 15 and the fourth spring ring are pressed by the measuring circuit contact points to deform 16 so as to be respectively in close contact with the conductive contact points, and thus, the electrodes are well connected with the measuring circuit and a power supply. The spring can be changed into conductive metal such as thimble, metal sheet, metal column, etc. When the spring is changed to the conductive rubber, the single needle planar shape sensor is as shown in fig. 8 (third conductive rubber 17, fourth conductive rubber 18).

As shown in fig. 9, reference electrode 1 is connected to wire 27 via conductive contact 5, and pin 19 compresses the wire. Wherein the wire is connected with the electrode and fixed on the contact plate by curing the conductive adhesive. The conductive adhesive not only plays a role in connection, but also plays a role in fixing the electrode.

The electrodes may be double electrodes or multiple electrodes. Whether a conductive rubber clamping mode or a metal clamping mode such as a spring, a metal sheet and the like or a point conductive rubber fixing mode is adopted, the resistance between the electrode and the contact of the measuring circuit after clamping connection and the stable fixing of the electrode are considered. The instability of the fixing of the electrode affects both the use of the electrode and the resistance between the electrode and the measuring circuit contact.

The principle of a double electrode sensor operating randes equivalent circuit is given below:

as shown in FIG. 10, the entire bipolar electrochemical system can be modeled using a Randes equivalent circuit, where R _et Z is the electron transfer resistance of the surface of the working electrode _w For Warburg impedance, C _dl R is the electric double layer capacitance of the electrode surface _S The electrolyte resistance is R is the loop equivalent resistance including the lead connection resistance. In the case of a wire only, the wire connection resistance R is negligible. However, as already mentioned, in the implantable biosensor electrode preparation technology, the sensor electrode must be connected to the contacts of the measuring circuit by a connection means, which may be implemented by conductive rubber, metal clips (springs), conductive glue, or the like. The equivalent resistance R of the circuit is mainly determined by the resistance between the electrode and the contacts of the measuring circuit. In particular, in the case of using a conductive paste or conductive rubber, the connection resistance may be large due to special conditions such as insufficient pressure of the contact pressing wire rubber, excessive resin component of the conductive paste, and the like. Then if the loop resistance R and R in the Randes equivalent circuit _et And R is _S The sum of R is not different by an order of magnitude and the value of R is not negligible. The result is that in the series loop where the whole sensor works, the connecting resistor shares a part of the voltage given by the power supply, and the working voltage of the sensor electrode is not the theoretical voltage given in the design process, so that the test result is greatly different from the actual result.

The inventors have studied to confirm that the impedance between the resistance of the connection means and the sensor electrode (including the working electrode surface electron transfer resistance R _et And electrolyte resistance R _S ) At least 2 orders of magnitude or more less, the resistance of the connection means will not affect the test results. The manner in which the sensor electrodes are connected to the contacts of the measuring circuit, such as the selected material resistance of the connection means, and the manner of connection must thus be taken into account when designing the sensor, in particular in designing an implantable glucose sensor.

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 may be a strip or a cylinder, or other geometric shape.

In the connection mode of fig. 7 and fig. 8, the sensor electrode is a planar electrode sensor using an insulating polymer sheet as a carrier, and a metal layer such as gold or platinum is deposited on the insulating polymer sheet substrate as a conductive catalytic layer by electroplating, sputtering or the like. The connection mode with the measuring circuit is that the measuring circuit is contacted with the electrode contact through a thimble, a spring, conductive rubber and the like. In this connection method, the connection resistance is too high due to the fact that the deposited metal layer such as gold or platinum is not dense in density and insufficient in bonding force.

3. Internal emitter

The internal transmitter is used to store measurement data, analyze the measurement data, correct the measurement data, and transmit the data. The current and voltage measured at the measurement circuit is a continuous datum, thus requiring the processing circuit of the internal transmitter to be able to average the data acquired.

The internal transmitter includes a voltage source, standard circuitry, measurement circuitry, and transmission circuitry.

1. Voltage source

When the thickness of the polymer film layer outside the sensor electrode is 3-10 mu m, the time for forming the stable hydrogel is not less than 3min. Before the polymer film layer is stable, body fluid continuously diffuses through the film layer into the sensor conductive layer, and besides some macromolecules and proteins, glucose, oxygen, ions and small molecular interferents enter the enzyme layer of the sensing area together and generate electrochemical reaction to form interference signals. A method of eliminating the influence of the above-described small molecule interferents by applying a continuously variable voltage is described below.

Electrochemical amperometric sensors, in which both the analyte and the interferent may react electrochemically, typically refer to the analyte and interferent as substrates, each of which has a different oxidation and reduction potential. The voltage source is designed to generate a continuous pulse voltage or step voltage that varies in the same direction or in different directions. The voltage application means is designed to allow electrochemical reaction of the analyte while simultaneously allowing all or part of the consumption of small molecule interferents that will diffuse into the sensor before forming a stable polymer film hydrogel. If N consecutive voltages are applied, then it is required that the value of the first voltage is minimal and at least the nth voltage and the N-1 th voltage are in the same direction and the value of the nth voltage is less than or equal to the value of the N-1 th voltage. In addition, the voltage source of the present invention comprises at least three consecutive voltages, wherein the first voltage is applied to the period of time before the polymer film hydrogel stabilizes, the second voltage is applied to the period of time that is considered to be capable of all or part of the consumption of small molecule interferents that will diffuse into the sensor, and the third voltage is applied until the end of the assay. The voltage of the second stage and the voltage of the third stage are required to be in the same direction, and the voltage of the first stage and the voltage of 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 values of the first stage voltage and the third voltage. The applied first stage voltage 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 primary interference into the enzyme layer and conductance catalysis to produce an electrochemical reaction. The applied third stage voltage mainly causes the analyte to react electrochemically.

2. Standard circuit

The manner in which the sensor electrodes are connected can affect the testing of the sensor. However, when a sensor is implanted in a living body, the accuracy of the obtained test result cannot be known before comparison with a standard value, and even if the accuracy of the test result is suspected, it cannot be judged whether the error is caused by the reason of the sensor electrode or the reason of the connection manner of the sensor electrode, which causes confusion to the factors failing in the analysis test. Therefore, a warning scheme is required to be provided for the failure of the test caused by the connection problem of the sensor electrode.

The standard circuit is a resistor with a fixed value, is connected with the sensor electrode in parallel to the direct-current voltage source circuit, and fig. 11 is an equivalent schematic diagram,R _c is the resistance of a standard circuit. Both the voltage and the current at both ends can be tested by the measuring circuit. The function of the standard resistor is used for judging whether the resistance of the connecting device influences the test result of the sensor. The resistance of the standard resistor may be determined based on the impedance between the sensor electrodes.

A method of judging whether or not the resistance of the connection means between the electrodes fails the test by using a standard circuit is described below.

The resistance between the sensor electrodes is first determined. Impedance R between sensor electrodes _e Is formed by electrode surface electron transfer resistance R _et And electrolyte resistance R _S Composition is prepared. The step can determine the electron transfer resistance R of the electrode surface by an impedance electrochemical method in the environment of the simulated human body tissue fluid of an organism _et And electrolyte resistance R _S Size of the product. As shown in FIG. 12, the impedance schematic diagram of the Randes equivalent circuit is that the Nyquist diagram of the circuit is formed by two parts of a semicircle and an oblique line which are connected, wherein the semicircle part measured in the high frequency area is controlled by an electron transfer process, the diameter of the semicircle part is equal to the value of Ret, and the front part of the semicircle is an electrolyte resistor R _S Then the resistance R of the whole sensor when implanted in the living body _e It can be considered that the sum R of values approximately equal to Ret _S Sum of values. FIG. 13 shows an impedance diagram of a dual electrode in vitro simulated human tissue fluid environment of the present invention, the sensor working electrode surface electron transfer resistance R _et The value is 1850KΩ, the electrolyte resistance R _S With a value of about 50K omega, the impedance R between the sensor electrodes _e About 1800K omega.

Second step, setting standard fixed resistance R _c ，R _c The resistance range belongs to the electrode resistance R _e From one percent to one tenth of a million. A smaller multiple is not possible for power consumption reasons.

And thirdly, judging whether the resistance R of the connecting device influences the test result of the sensor. From the foregoing, the inventors have confirmed that when the impedance between the resistance of the connection means and the sensor electrode is at least 2 orders of magnitude smaller or more, the test result is closer to the true value. The current of the 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. Measuring circuit tests the current between the sensor electrodes as I _e . If I _s Greater than 101 times I _e Indicating that the resistance R of the connecting means is greater than the resistance between the sensor electrodes (including the working electrode surface electron transfer resistance R _et And electrolyte resistance R _S ) 2 orders of magnitude, the resistance R of the connecting device is overlarge, and the test result of the sensor is affected. If I _s Less than 101 times of I _e Indicating that the resistance R of the connecting means is smaller than the resistance between the sensor electrodes (including the working electrode surface electron transfer resistance R _et And electrolyte resistance R _S ) The resistance R of the connecting device cannot influence the test result of the sensor by 2 orders of magnitude.

In addition, in the electrode structure designs of fig. 5 and 6 according to an embodiment of the present invention, the electrode working electrode and the reference electrode are separated by an insulating layer therebetween, and the insulating layer may cause defects such as open bottom and dishing during the preparation of the insulating layer or during the reference preparation, for example, the non-uniformity of the coating of the insulating polymer may cause the polymer to not completely cover the conductive layer, or the insulating polymer may be detached due to abrasion, etc., so that the working electrode and the reference electrode are directly conducted and not isolated. In this case, the working electrode and the reference electrode directly form a loop, no electron transfer through the medium of the tissue fluid is required, no electrochemical reaction is generated by the sensor electrode, and the sensor fails. At this time, the resistance R of the sensor electrode _e The resistance of (2) is almost zero. I at this time _s And I _e Approximately equal. In another example of the present invention, as shown in the electrode structure designs of FIGS. 7 and 8, when the working electrode and the reference electrode are in the process of preparation, the working electrode and the reference electrode are directly conducted, I _s And I _e And are also nearly equal.

3. Measuring circuit

The measuring circuit is designed to be controlled by an internal transmitter processing circuit, comprises a sampling modulation circuit, an ADC conversion circuit and the like, and can collect current data of a sensor and current and voltage data of a 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 acquired current and voltage data into digital signals after analog-to-digital conversion by the ADC circuit and send the digital signals to the internal transmitter processor for analysis. The measurement circuit also includes a digital filter, which is typically required to filter the acquired data to smooth the acquired raw data.

4. Transmission circuit

The transmission circuit is controlled by the processor and is designed to transmit the data packet established by the processor to the external controller and establish unique communication with the external controller. The transmitted data comprise sensor identification, a unique identifier corresponding to the sensor and an external controller, sensor original data and data processed by a processor. The transmission circuit is also designed to receive instructions issued by the external controller.

4. External processor

The external processor is used for storing the data transmitted by the internal transmitter, evaluating the analysis result, comparing the results of each time period of the analyte, giving a change trend, evaluating the analysis result, and simultaneously being used for data downloading, daily time input, abnormal early warning and the like of the analyte.

The external processor also includes an acceleration module. The exercise monitoring, sleep quality, sedentary reminding and prompting, fall alarming and other functions can be completed. The step counting information and the speed information provided by the exercise monitoring are matched with the electrocardio/heart rate module to provide accurate quantitative reference for the exercise intensity, so that a user is guided to control blood sugar through proper exercise. In addition, since the diabetes patient is more likely to suffer from syncope and fall risk caused by hypoglycemia during the fall alarm, it is necessary to send a remote alarm to the fall (e.g., send the fall alarm and GPS position to another person through a telephone function module) in an unmanned situation.

The external controller processor circuit also comprises a heart rate monitoring module, wherein the heart rate module mainly refers to dynamic heart rate, adopts a photoelectric principle to monitor, and can provide references for the movement duration and the movement intensity of a user in cooperation with movement information provided by the acceleration module. The quantity of motion directly influences the change of blood sugar, and a user can predict the quantity of motion and the intensity needed to be carried out next time through the action of the motion time length and the motion intensity on the blood sugar, so that the blood sugar control can be known more accurately. The electrocardio module is an ECG circuit and mainly reads an electrocardiogram. There are data showing that diabetics have no heart disease and have different degrees of abnormal changes in the electrocardiogram. Mortality from the diabetic heart accounts for about 70% of the mortality rate from diabetes, with arrhythmia being one of the leading causes of mortality. It is also of interest for diabetics to monitor dynamic electrocardiography. An electrocardio sensor electrode is arranged at the outer side of the processor, and is started to work when the glucose is detected, so that the heart rate change of an analyte can be directly monitored

The external processor also includes a calibration input module. The test paper is mainly used for collecting blood glucose concentration readings. This reading is directly used to calibrate the continuous blood glucose monitoring algorithm. When the test strip is inserted, the watch displays a blood glucose test strip reading interface, at the moment, the blood concentration of the finger can be sampled and tested, and after the finger is confirmed and stored, the processor directly stores the value as the reference blood glucose for algorithm calibration.

Example 2

In this example, the electrode carrier, i.e. the electrically conductive catalytic layer, is a platinum wire. After the platinum wire is ion activated, a layer of parylene of about 8 microns is deposited by vapor deposition as an insulating layer for the working electrode and the reference electrode. The dip coating method is adopted, and the proportion is 50:50 of silver/silver chloride paste is coated on the surface of the insulating layer and is cured and dried by ultraviolet light. A 2.5mm section was etched with a laser to expose a platinum wire as the sensing region, with a guard length of 0.75mm. The glucose oxidase solution containing silane is coated on the sensing area, and is crosslinked by a crosslinking agent, dried and solidified to form a section of glucose sensing area with the length of 2.5 mm. A layer of polyurethane hydrogel is coated as an external macromolecular disruptor, which hydrogel serves to limit the diffusion of glucose and oxygen in the interstitial fluid at the same time.

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 an included angle of 45 degrees is formed between the electrode and the horizontal. The sensor electrode is implanted into subcutaneous tissue of the living being by means of an auxiliary piercing tool, the implantation length being 8mm. The thimble of the measuring circuit of the internal transmitter compresses the conductive rubber, and the resistance among the electrode, the conductive rubber and the measuring circuit is 8.31KΩ ohm. The impedance between the electrodes is 1850kΩ, and the resistance of the standard circuit is 10kΩ.

The voltage source of the internal emitter applies the first stage voltage V1 at a value of 0.60V for 9min, and at the application voltage of 3min, the if I is found _s And I _e Is 1056, which is much greater than 101. The connection scheme of the electrodes by using conductive rubber is shown to be feasible, and the resistance is far smaller than the impedance between the electrodes of the sensor. The second stage voltage V2 was applied at 1.07V, in the same direction as the first stage voltage, for a duration of 5min. The third stage voltage was applied at 0.60V for a period of time until the end of the test, in the manner shown in fig. 14.

The current and voltage signals tested by the measuring circuit are transmitted to an external controller through a transmission circuit, and the test result is displayed through data processing.

Example 3

The electrode of the analysis system for measuring glucose prepared in example 2 was placed in a phosphate buffer containing glucose at a concentration of 5mmol/L, 10mmol/L, 15mmol/L, 20mmol/L.

As can be seen from FIG. 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 analytical system for glucose measurement of example 2 was sterilized and then implanted into a living body for 11 days, and the measurement result was recorded.

As shown in fig. 16, the continuous curve represents the continuous test result of this example, the dot represents the biochemical blood test result, and as shown in fig. 17, the curve is a 6-hour comparison of the biochemical blood test result and the example 2 of the present invention, wherein the detection period spans the feeding activity and the rest activity, the biochemical result is recorded every 15 minutes, the solid line is the result of example 2, the dot is the result of biochemical blood test, and the dot of biochemical blood test is fitted to the curve of the broken line. After analysis of the two curves, the standard error for the whole test phase was 11.2%, the standard error for the hyperglycemic period was 13.7% and the hypoglycemic error was 12.5%.

Claims

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.