Multi-analyte monitoring sensors

Abstract

The present disclosure describes a multi-analyte monitoring sensor, the sensor includes a substrate, a working electrode arranged on the substrate, an enzyme sensing layer arranged on the working electrode, and a polymer film layer arranged on the enzyme sensing layer, the working electrode includes a first working electrode and a second working electrode; the enzyme sensing layer includes a first analyte enzyme layer and a second analyte enzyme layer, the first analyte enzyme layer is arranged on the first working electrode, the second analyte enzyme layer is arranged on the second working electrode, the first analyte is glucose, and the second analyte is ketone; the polymer film layer can penetrate the first analyte and the second analyte and covers the first analyte enzyme layer and the second analyte enzyme layer. Thus, a sensor for continuous monitoring of blood glucose and blood ketone with high sensitivity and simple process can be provided.

Description

Technical Field

[0001] The present disclosure relates generally to the field of biosensors, and more particularly to a multi-analyte monitoring sensor. Background Art

[0002] Biosensors are analytical devices that tightly combine biological materials, bio-derived materials, or biomimetic materials with optical, electrochemical, temperature, piezoelectric, magnetic, or micromechanical physical and chemical sensors or sensing microsystems. They can usually be used to quickly detect certain specific chemicals in the human body, such as glucose, ketone bodies, uric acid, and a series of amino acid compounds.

[0003] When blood sugar is too high, the body lacks insulin, or a ketogenic diet is followed, the body will metabolize fat to produce ketone bodies. When the amount of ketone bodies exceeds the metabolic capacity of the liver, complications of diabetes are likely to occur. Therefore, monitoring of blood ketone levels is also of great significance in diabetic patients.

[0004] Traditional blood ketone detection methods mainly include urine ketone test paper method, blood ketone test paper method and biochemical analysis method, etc., which detect the content of β-hydroxybutyric acid (abbreviated as β-BHB or hydroxybutyric acid), acetone and acetoacetic acid in blood or urine in a quantitative or qualitative manner to reflect the ketone body level. Usually, wearable acetone sensors detect acetone in the exhaled gas of diabetic patients, convert the resistance change into a digital value and send it to the display module to display the acetone concentration.

[0005] However, the response time of the above-mentioned acetone sensor is usually about 50 seconds, and the evaluation of acetone usually has a lag; and when it is necessary to monitor blood sugar and blood ketones at the same time, it is generally necessary to set up a blood sugar sensor to monitor blood sugar, which is inconvenient to use. Summary of the invention

[0006] The present disclosure is proposed in view of the above situation, and its object is to provide a multi-analyte monitoring sensor that can continuously monitor blood glucose and blood ketones simultaneously.

[0007] To this end, the present disclosure provides a multi-analyte monitoring sensor, the sensor comprising a substrate, a working electrode arranged on the substrate, an enzyme sensing layer arranged on the working electrode, and a polymer film layer arranged on the enzyme sensing layer, the working electrode comprising a first working electrode and a second working electrode; the enzyme sensing layer comprising a first analyte enzyme layer and a second analyte enzyme layer, the first analyte enzyme layer being arranged on the first working electrode, the second analyte enzyme layer being arranged on the second working electrode, the first analyte being glucose, and the second analyte being ketone; the polymer film layer being permeable to the first analyte and the second analyte and covering the first analyte enzyme layer and the second analyte enzyme layer.

[0008] In the multi-analyte monitoring sensor involved in the present disclosure, by setting two analyte enzyme layers on the two working electrodes of the sensor, the concentration levels of two different analytes in the tissue fluid can be detected respectively; the polymer film layer can simultaneously permeate glucose and ketones, and by covering the first analyte enzyme layer and the second analyte enzyme layer with the polymer film layer, the permeation of the two analytes can be restricted, and at the same time, the stability of the sensor can be improved; in addition, due to the small size of the biosensor, the precision of the film coating on the surface of the sensor is required to be high, and the immersion pulling method is usually used to cover the surface of the sensor with a membrane liquid to form a film layer covering the sensor. Compared with the scheme of setting a semipermeable membrane that can permeate glucose and ketones on the first analyte enzyme layer and the second analyte enzyme layer, the polymer film in the present invention can simultaneously permeate glucose and ketones, thereby simplifying the manufacturing process of the sensor while maintaining the high sensitivity and accuracy of the sensor. Thus, a sensor with a simple manufacturing process and capable of simultaneously and continuously monitoring glucose and ketones can be provided.

[0009] In addition, in the multi-analyte monitoring sensor of the present disclosure, optionally, the polymer film layer includes a vinyl pyridine polymer, a cross-linking agent, and a regulator, and the mass fraction of the vinyl pyridine polymer in the polymer film layer is 80% to 100%. Thus, it is conducive to forming a polymer film layer that has a limiting permeation effect on glucose and ketones and has biocompatibility, and the amount of glucose and ketones can be controlled within the linear range of the sensor through the polymer film, thereby improving the sensitivity and accuracy of the sensor.

[0010] In addition, in the multi-analyte monitoring sensor of the present disclosure, optionally, the polymer film layer includes a first film layer sequentially disposed on the enzyme sensing layer, a transition layer formed on the first film layer, and a second film layer formed on the transition layer and having biocompatibility. Thus, the bonding stability of the film layer can be improved, which is beneficial to improving the stability of the sensor.

[0011] In addition, in the multi-analyte monitoring sensor of the present disclosure, optionally, the first film layer is formed of a first type of polymer, the second film layer is formed of a second type of polymer, and the transition layer is formed of a third type of polymer, and the third type of polymer is a copolymer formed of a first monomer that is the same as or similar to the first type of polymer and a second monomer that is the same as or similar to the second type of polymer. In this way, the bonding stability and mechanical strength of the polymer film layer can be improved, thereby improving the stability of the sensor.

[0012] In addition, in the multi-analyte monitoring sensor of the present disclosure, optionally, the first film layer comprises a vinyl pyridine polymer, and the transition layer and the second film layer comprise a vinyl pyridine and styrene copolymer, thereby improving the film-forming performance of the film layer and thus enhancing the sensitivity of the sensor.

[0013] In addition, in the multi-analyte monitoring sensor involved in the present disclosure, optionally, the area of ​​the first analyte enzyme layer is greater than the area of ​​the second analyte enzyme layer. Since the amount of glucose and ketones that can pass through the polymer membrane is different, and the amount of glucose is usually less than the amount of ketones, in this case, by configuring the area of ​​the first analyte enzyme layer to be larger than the area of ​​the second analyte enzyme layer, the contact area between the first analyte enzyme layer and glucose can be increased, thereby minimizing the current difference between the two working electrodes, making the current on the two working electrodes at the same level, reducing the interference of the electric field on the sensor system, and further improving the measurement accuracy of the sensor.

[0014] In addition, in the multi-analyte monitoring sensor of the present disclosure, optionally, the area ratio of the first analyte enzyme layer to the second analyte enzyme layer is 1: 1 to 1.5: 1. Thus, the reaction sensitivity of the first analyte can be improved, thereby more accurately reflecting the concentration level of the analyte in the tissue fluid.

[0015] In addition, in the multi-analyte monitoring sensor of the present disclosure, optionally, the first analyte enzyme layer and / or the second analyte enzyme layer are linear. In this case, compared with the multi-point discrete arrangement scheme, a larger area of ​​enzyme layer can be arranged on the working electrode of limited area, thereby improving the sensitivity of the sensor.

[0016] In addition, in the multi-analyte monitoring sensor involved in the present disclosure, optionally, the second analyte enzyme layer includes hydroxybutyrate dehydrogenase, diaphorase, coenzyme, and electron mediator, and in the second analyte enzyme layer, the mass fraction of hydroxybutyrate dehydrogenase is 10% to 20%, the mass fraction of diaphorase is 5% to 20%, the mass fraction of coenzyme is 10% to 30%, and the mass fraction of the electron mediator is 10% to 30%. In this case, the reaction sensitivity of the second analyte enzyme layer to β-BHB can be improved, and by providing a macromolecular electron mediator in the enzyme layer, hydroxybutyrate dehydrogenase can be covalently bound to the electron mediator, thereby further improving stability.

[0017] In addition, in the multi-analyte monitoring sensor involved in the present disclosure, optionally, the sensor further includes a reference electrode and a counter electrode, the first working electrode and the second working electrode are respectively located on both sides of the substrate, and the reference electrode and the counter electrode are respectively located on both sides of the substrate. In this case, a four-electrode two-way detection sensor can be formed, and arranging the four electrodes in pairs on both sides of the substrate can improve the utilization rate of the substrate, while further reducing the interference of the electric field on the sensor system.

[0018] According to the present disclosure, a multi-analyte monitoring sensor can be provided, which can continuously monitor glucose and ketones simultaneously, has a simple manufacturing process, and has high sensitivity and stability. BRIEF DESCRIPTION OF THE DRAWINGS

[0019] The present disclosure will now be explained in further detail by way of example only with reference to the accompanying drawings, in which:

[0020] Figure 1 is a schematic diagram showing a multi-analyte monitoring sensor according to an example of the present disclosure.

[0021] Figure 2A is a front schematic diagram showing a multi-analyte monitoring sensor according to an example of the present disclosure.

[0022] Figure 2B is a schematic diagram showing the reverse side of the multi-analyte monitoring sensor involved in the examples of the present disclosure.

[0023] Figure 3 Schematic diagram showing a polymer film layer involved in an example of the present disclosure.

[0024] Figure 4 1 is a graph showing a response current-test concentration curve of Example 1 of the blood glucose+blood ketone sensor according to the present disclosure.

[0025] Figure 5 is a picture showing a monitoring probe of Example 2 of the blood glucose + blood ketone sensor according to the present disclosure.

[0026] Figure 6 is a graph showing a response current-test time curve of Example 2 of the blood glucose+blood ketone sensor according to the present disclosure.

[0027] Figure 7 1 is a graph showing a response current-test concentration curve of Example 2 of the blood glucose+blood ketone sensor according to the present disclosure.

[0028] Figure 8 1 is a glucose / sodium β-hydroxybutyrate response current-time curve of Example 2 of the blood glucose+blood ketone sensor according to the present disclosure.

[0029] Fig. 9 1 is a graph showing a response current-test concentration curve of Example 3 of the blood glucose+blood ketone sensor according to the present disclosure.

[0030] Fig.10 is a picture showing a monitoring probe of Comparative Example 1 of the blood glucose + blood ketone sensor according to the present disclosure.

[0031] Fig.11 is a response current-test concentration curve graph of Comparative Example 1 of the blood glucose+blood ketone sensor according to the present disclosure.

[0032] Fig.12 is a picture showing the monitoring probe of Comparative Example 2 of the blood glucose + blood ketone sensor according to the present disclosure.

[0033] Fig.13 is a response current-test time curve graph of Comparative Example 2 of the blood glucose+blood ketone sensor according to the present disclosure.

[0034] Fig.14 is a picture showing the monitoring probe of Comparative Example 3 of the blood glucose + blood ketone sensor according to the present disclosure.

[0035] Fig.15 2 is a picture showing the monitoring probe of Comparative Example 3 of the blood glucose + blood ketone sensor according to the present disclosure from another perspective.

[0036] Fig.16 1 is a picture showing the monitoring probes of the sensors of Example 1 and Comparative Examples 1-2 of the blood ketone sensor according to the present disclosure.

[0037] Fig.17 1 is a picture showing monitoring probes coated with a polymer film layer in Example 1 and Comparative Examples 1-2 of the blood ketone sensor according to the present disclosure.

[0038] Fig.18 is a graph showing a response current-test concentration of a linearity test of Example 1 of the blood ketone sensor according to the present disclosure.

[0039] Fig.19 1 is a graph showing a response current-test concentration curve of a linearity test of Example 1 of the blood ketone sensor according to the present disclosure.

[0040] Fig. 20 1 is a graph showing a response current-test concentration curve of a stability test of Example 1 of the blood ketone sensor according to the present disclosure.

[0041] Fig.21 1 is a graph showing a response current-test concentration curve of a stability test of Comparative Example 1 of the blood ketone sensor according to the present disclosure.

[0042] Fig. 22 3 is a graph showing a response current-test concentration curve of a stability test of Comparative Example 2 of the blood ketone sensor according to the present disclosure.

[0043] Fig.23 3 is a graph showing a response current-test concentration of a linearity test of Comparative Example 3 of the blood ketone sensor according to the present disclosure.

[0044] Fig.24 is a picture showing a working electrode of Comparative Example 4 of the blood ketone sensor according to the present disclosure.

[0045] Fig.25 is a picture showing a monitoring probe of a sensor of Comparative Example 4 of the blood ketone sensor according to the present disclosure.

[0046] Fig.26 4 is a graph showing a response current-test time curve of the sensor of Comparative Example 4 of the blood ketone sensor according to the present disclosure.

[0047] Fig. 27 is a picture showing a working electrode of the sensor of Comparative Example 5 of the blood ketone sensor according to the present disclosure.

[0048] Fig.28 3 is a graph showing a response current-test concentration of a linearity test of a blood ketone sensor comparative example 5 according to the present disclosure.

[0049] Description of reference numerals:

[0050] 10 ... monitoring probe, 11 ... substrate, 12 ... first working electrode, 121 ... first analyte enzyme layer, 13 ... second working electrode, 131 ... second analyte enzyme layer, 14 ... semipermeable membrane, 141 ... first membrane layer, 142 ... transition layer, 143 ... second membrane layer, 15 ... reference electrode, 16 ... counter electrode DETAILED DESCRIPTION

[0051] Hereinafter, the preferred embodiments of the present disclosure will be described in detail with reference to the accompanying drawings. In the following description, the same symbols are assigned to the same components, and repeated descriptions are omitted. In addition, the accompanying drawings are only schematic diagrams, and the ratio of the dimensions of the components or the shapes of the components may be different from the actual ones.

[0052] It should be noted that the terms "including" and "having" and any variations thereof in the present disclosure, such as a process, method, system, product or device that includes or has a series of steps or units, are not necessarily limited to those steps or units clearly listed, but may include or have other steps or units that are not clearly listed or inherent to these processes, methods, products or devices.

[0053] In addition, the subheadings and the like involved in the following description of the present disclosure are not intended to limit the content or scope of the present disclosure, but are merely used as a reminder for reading. Such subheadings should neither be understood as being used to divide the content of the article, nor should the content under the subheading be limited to the scope of the subheading.

[0054] The present disclosure relates to a multi-analyte monitoring sensor in one aspect, a semipermeable membrane of a multi-analyte monitoring sensor in a second aspect, and a blood ketone sensor in a third aspect.

[0055] In the present disclosure, the multi-analyte monitoring sensor may also be referred to as a "multi-analyte sensor", "blood glucose and blood ketone continuous monitoring sensor", "blood glucose and blood ketone sensor", "blood glucose + blood ketone sensor", "physiological parameter sensor" or "sensor", "sensing probe", "sensing probe", "implanted probe", etc.

[0056] The sensor involved in the present disclosure can be used to monitor physiological parameters of a host. The physiological parameters can be parameters such as glucose, urea, uric acid, ketone bodies, and a series of amino acid compounds in the host.

[0057] In the present disclosure, the sensor can also be used to detect the object to be detected in the host. The object to be detected can be a chemical substance in body fluids. For example, the object to be detected can be one or more of glucose, acetylcholine, amylase, bilirubin, cholesterol, chorionic gonadotropin, creatine kinase, creatine, creatinine, DNA, fructosamine, glucose, glutamine, growth hormone, hormone, ketone body, lactate, oxygen, peroxide, prostate specific antigen, prothrombin, RNA, thyroid stimulating hormone or troponin. In addition, the object to be detected can also be a drug in body fluids. For example, the object to be detected can be digitoxin, digoxin, theophylline, warfarin or an antibiotic (such as gentamicin, vancomycin, etc.).

[0058] Hereinafter, the multi-analyte monitoring sensor involved in the first aspect of the present disclosure will be described in conjunction with the accompanying drawings.

[0059] Figure 1 is a schematic diagram showing a multi-analyte monitoring sensor according to an example of the present disclosure.

[0060] In this embodiment, the sensor 10 can be used to obtain the physiological parameter information of the host. In some examples, the sensor 10 can be applied to the host, and the host can obtain its own physiological parameter information by applying the monitoring device 1 to itself. In some examples, the sensor 10 can be mounted on the body surface parts such as the host's arms, back, abdomen, waist, legs, etc. where the sensor 10 can be partially implanted to monitor physiological parameters. In some examples, the sensor 10 can be used to be implanted subcutaneously in the host to obtain a sensing signal (e.g., a current signal). In some examples, when the sensor 10 is applied to the host, it can contact the tissue fluid or blood in the host to measure the analyte level in the tissue fluid or blood.

[0061] In some examples, the sensor 10 includes a substrate 11 and a working electrode disposed on the substrate. In some examples, there may be multiple working electrodes, and each working electrode may be used to detect a different analyte. Thus, the sensor 10 can monitor multiple analytes.

[0062] In some examples, the working electrode may include a first working electrode 12 and a second working electrode 13. The first working electrode 12 may be used to detect a first analyte, and the second working electrode 13 may be used to detect a second analyte. Thus, the sensor 10 can monitor two analytes.

[0063] In some examples, the first analyte may be any one of glucose, urea, uric acid, ketone bodies, creatinine, ethanol and lactic acid. In some examples, the second analyte may be any one of glucose, urea, uric acid, ketone bodies, creatinine, ethanol and lactic acid. Among them, ketone bodies (which may be referred to as ketones) are collectively referred to as acetoacetic acid, β-hydroxybutyric acid and acetone, the intermediate products of fatty acid oxidation and decomposition in the liver. In the present invention, the content of ketone bodies can be monitored by detecting any one of acetoacetic acid, β-hydroxybutyric acid and acetone. Preferably, the β-hydroxybutyric acid in the host can be monitored to obtain parameters related to the ketone body level.

[0064] In some examples, the first analyte may be glucose. Thus, the glucose concentration in the body can be monitored. In some examples, the second analyte may be ketone bodies. Thus, the ketone body content in the body can be monitored. It should be noted that, as mentioned above, ketone bodies are a general term for acetoacetic acid, β-hydroxybutyric acid and acetone, the intermediate products of fatty acid oxidation and decomposition in the liver. Therefore, the second analyte may be any one of acetoacetic acid, β-hydroxybutyric acid and acetone. Preferably, the second analyte may be β-hydroxybutyric acid.

[0065] Hereinafter, the sensor 10 involved in the present disclosure will be described in detail by taking the case where the first analyte is glucose and the second analyte is ketone as an example.

[0066] In some examples, the sensor 10 may include an enzyme sensing layer disposed on a working electrode. In an example where the working electrode includes a first working electrode 12 and a second working electrode 13, the enzyme sensing layer may include a first analyte enzyme layer 121 and a second analyte enzyme layer 131. The first analyte enzyme layer 121 may be disposed on the first working electrode 12, and the second analyte enzyme layer 131 may be disposed on the second working electrode 13. In this case, by disposing enzyme layers of two analytes on the two working electrodes of the sensor 10, the concentration levels of two different analytes in the tissue fluid can be detected respectively.

[0067] In some examples, the sensor 10 may include a polymer film layer 14. The polymer film layer 14 may be configured to be permeable to the first analyte and the second analyte and cover the first analyte enzyme layer 121 and the second analyte enzyme layer 131. In this case, the polymer film layer 14 can limit the permeability of the two analytes, and at the same time, it is beneficial to improve the stability of the sensor. Compared with the solution of respectively setting a semipermeable membrane permeable to the first analyte and the second analyte on the first analyte enzyme layer 121 and the second analyte enzyme layer 131, the polymer film 14 in the present invention can simultaneously permeate the first analyte and the second analyte, thereby simplifying the manufacturing process of the sensor 10 while maintaining the high sensitivity and accuracy of the sensor 10. In some examples, the polymer film layer 14 can simultaneously permeate glucose and ketones. In this case, by covering the first analyte enzyme layer 121 and the second analyte enzyme layer 131 with the polymer film layer 14, the permeability of the two analytes can be restricted, and the stability of the sensor can be improved. In addition, since the volume of the biosensor 10 is small, the precision of the coating on the surface of the sensor 10 is required to be high. Usually, the surface of the sensor 10 is covered with a film liquid by an immersion pull method to form a film layer covering the sensor 10. Compared with the solution of setting a semipermeable membrane that can pass glucose and ketone on the first analyte enzyme layer 121 and the second analyte enzyme layer 131, the polymer film 14 in the present invention can pass glucose and ketone at the same time, thereby simplifying the manufacturing process of the sensor 10 while maintaining the high sensitivity and accuracy of the sensor 10. Therefore, a sensor 10 with a simple manufacturing process and capable of continuously monitoring glucose and ketone at the same time can be provided.

[0068] In some examples, the polymer film layer 14 may have a permeability coefficient of 1 to 1000 for glucose. Preferably, the polymer film layer 14 may have a permeability coefficient of 300 to 800 for glucose. For example, the polymer film layer 14 may have a permeability coefficient of 300, 400, 500, 600, 700 or 800 for glucose. In some examples, the polymer film layer 14 may have a permeability coefficient of 1 to 1000 for ketone. Preferably, the polymer film layer 14 may have a permeability coefficient of 200 to 600 for ketone. For example, the polymer film layer 14 may have a permeability coefficient of 200, 300, 400, 500 or 600 for ketone. The permeability coefficient refers to the concentration difference between the inside and outside of the polymer film layer. For example, if the permeability coefficient of the polymer film layer for glucose is 500, the concentration of glucose outside the polymer film layer is 500 times the concentration of glucose inside the polymer film layer.

[0069] In some examples, the polymer film layer 14 can have different permeability coefficients by adjusting the thickness ratio of the polymer film layer 14. In this case, by adjusting the permeability coefficient of the polymer film layer 14, the amount of the analyte that passes through can be controlled, so that the current generated by the working electrode is more accurate, thereby making the sensor 10 more sensitive.

[0070] Figure 2A is a schematic diagram showing the front side of a multi-analyte monitoring sensor according to an example of the present disclosure. Figure 2B is a schematic diagram showing the reverse side of the multi-analyte monitoring sensor involved in the examples of the present disclosure.

[0071] In some examples, the sensor 10 can be used in conjunction with an electronic system, and the sensor 10 and the electronic system cooperate to form a sensing assembly. Specifically, the sensor 10 can be electrically connected to the electronic system, and the sensor 10 can be partially implanted under the host's skin to obtain a sensing signal indicating the level of the analyte, and then transmit the sensing signal to the electronic system; the electronic system can process and / or retransmit the sensing signal. In this case, the analyte concentration in the body can be accurately obtained.

[0072] In some examples, the sensor 10 may also include a reference electrode 15 and a counter electrode 16 (see Figure 2A and Figure 2B In this case, a four-electrode dual-path detection sensor can be formed, that is, the first working electrode 12, the reference electrode 15 and the counter electrode 16 can form a set of three-electrode circuits, and the second working electrode 13, the reference electrode 15 and the counter electrode 16 can also form a set of three-electrode circuits.

[0073] In some examples, the four-electrode two-way detection sensor may have a first working electrode 12, a second working electrode 13, a reference electrode 15, and a counter electrode 16, and the first working electrode 12 forms a first loop with the reference electrode 15 and the counter electrode 16, and the second working electrode 13 forms a second loop with the reference electrode 15 and the counter electrode 16. In this case, the number of electrodes in the sensor can be reduced, so that the length of the substrate can be reduced, and the implantation depth of the sensor 10 can be reduced. Since the greater the implantation depth, the more obvious the pain generated, thus, the pain of implanting the sensor 10 can be reduced, and the wearing comfort can be improved.

[0074] In some examples, the first working electrode 12 and the second working electrode 13 can be located on both sides of the substrate 11, respectively, and the reference electrode 15 and the counter electrode 16 are located on both sides of the substrate 11, respectively. In this case, arranging the four electrodes in pairs on both sides of the substrate can improve the utilization rate of the substrate, while optimizing the voltage drop of the sensor electrode, reducing errors, and further reducing the interference of the electric field on the sensor system. In some examples, since the wearing comfort is related to the implantation depth, the longer the substrate, the greater the implantation depth, and the wearing pain may also increase. Therefore, arranging the four electrodes in pairs on both sides of the substrate can improve the utilization rate of the substrate, and arranging four electrodes in a limited substrate space can reduce the length of the substrate, thereby improving the wearing comfort.

[0075] In some examples, the reference electrode 15 can form a known and fixed potential difference with the tissue fluid or blood. In this case, the potential difference between the working electrode and the tissue fluid or blood can be measured by the potential difference formed by the reference electrode 15 and the working electrode. As a result, the voltage generated by the working electrode can be obtained more accurately, and the electronic system can automatically adjust and maintain the stability of the voltage at the working electrode according to the preset voltage value, so that the measured current signal can more accurately reflect the analyte concentration level information in the tissue fluid or blood. In some examples, the number of reference electrodes 15 can be one or more, for example, two. In this case, the number of reference electrodes can be adjusted according to the number of working electrodes or the current generated by the working electrodes, so that an accurate current signal can be obtained, thereby improving the sensitivity of the sensor.

[0076] In some examples, the sensor 10 can be divided into an implant portion and a connection portion. The implant portion can be placed under the skin of the host and connected to the electronic system through the connection portion. In this case, by placing the implant portion under the skin of the host and electrically connecting the implant portion to the electronic system through the connection portion, it is possible to facilitate the transmission of the sensing signal obtained by the implant portion to the electronic system. In some examples, the implant portion can be rigid. Thus, it can help to be placed under the skin of the host. In other examples, the implant portion can also be flexible, in which case the host's foreign body sensation can be reduced.

[0077] In some examples, as described above, the sensor 10 may include a first analyte enzyme layer 121. In this case, the first analyte can react in the first analyte enzyme layer 121 to generate a current, and a concentration signal of the first analyte can be obtained.

[0078] In some examples, the first analyte enzyme layer 121 may include glucosylase. Thus, glucose can react in the first analyte enzyme layer, thereby generating a glucose concentration signal. In some examples, in the first analyte enzyme layer 121, the content (mass fraction) of glucosylase can be 40% to 70%. For example, the content (mass fraction) of glucosylase can be 40%, 41%, 42%, 43%, 44%, 45%, 46%, 47%, 48%, 49%, 50%, 51%, 52%, 53%, 54%, 55%, 56%, 57%, 58%, 59%, 60%, 61%, 61%, 62%, 63%, 64%, 65%, 66%, 67%, 68%, 69%, or 70%. Thus, it is possible to facilitate glucose to react.

[0079] In some examples, the first analyte enzyme layer 121 can include an electron mediator. In some examples, the electron mediator can be a redox polymer. In some examples, the redox polymer can be a metal redox polymer. In some examples, the metal redox polymer can be selected from at least one of poly(vinylferrocene), quaternized poly(4-vinylpyridine) of ferrocyanide, quaternized poly(1-vinylimidazole) of ferrocyanide, quaternized poly(4-vinylpyridine) of ferrocyanide, quaternized poly(1-vinylimidazole) of ferrocyanide, quaternized poly(1-vinylimidazole) of ferrocyanide, osmium 2,2'-bipyridine complex coordinated to poly(1-vinylimidazole), osmium 2,2'-bipyridine complex coordinated to poly(4-vinylpyridine), ruthenium 2,2'-bipyridine complex coordinated to poly(1-vinylimidazole), ruthenium 2,2'-bipyridine complex coordinated to poly(4-vinylpyridine), cobalt 2,2'-bipyridine complex coordinated to poly(1-vinylimidazole) or cobalt 2,2'-bipyridine complex coordinated to poly(4-vinylpyridine). Thus, metal redox polymers can participate in redox reactions via covalent, coordinate or ionic bonds.

[0080] In some examples, in the first analyte enzyme layer 121, the content (mass fraction) of the electron mediator can be 10% to 40%. For example, the content (mass fraction) of the electron mediator can be 10%, 11%, 12%, 13%, 14%, 15%, 16%, 17%, 18%, 19%, 20%, 21%, 22%, 25%, 26%, 27%, 28%, 29%, 30%, 31%, 32%, 33%, 34%, 35%, 36%, 37%, 38%, 39%, or 40%. Thus, the first analyte can be promoted to react in the first analyte enzyme layer.

[0081] In some examples, the first analyte enzyme layer 121 may include a crosslinker. In some examples, the crosslinker may be PEGDGE. In some examples, the crosslinker may be PEGDGE400. In some examples, in the first analyte enzyme layer 121, the content (mass fraction) of the crosslinker may be 0% to 20%. For example, the content (mass fraction) of the crosslinker may be 0%, 1%, 2%, 3%, 4%, 5%, 6%, 7%, 8%, 9%, 10%, 11%, 12%, 13%, 14%, 15%, 16%, 17%, 18%, 19%, or 20%. Thus, it is possible to facilitate long-term monitoring of the sensor.

[0082] In some examples, the area of ​​the first analyte enzyme layer 121 may be larger than the area of ​​the second analyte enzyme layer 131. Since the amount of glucose and ketones that can pass through the polymer membrane 14 is different, and the amount of glucose is usually less than the amount of ketones, in this case, by configuring the area of ​​the first analyte enzyme layer 121 to be larger than the area of ​​the second analyte enzyme layer 131, the contact area between the first analyte enzyme layer 121 and glucose can be increased, thereby reducing the current difference between the two working electrodes as much as possible, making the current magnitudes on the two working electrodes at the same level, reducing the interference of the electric field on the sensor 10 system, and further improving the measurement accuracy of the sensor 10.

[0083] In some examples, the area ratio of the first analyte enzyme layer 121 to the second analyte enzyme layer 131 may be greater than 1:1. Thus, it is possible to improve the sensitivity of the sensor 10. In some examples, the area ratio of the first analyte enzyme layer 121 to the second analyte enzyme layer 131 may be less than 1.5:1. Thus, the accuracy of the sensor 10 can be improved. In some examples, the area ratio of the first analyte enzyme layer 121 to the second analyte enzyme layer 131 may be 1:1 to 1.5:1. Thus, the reaction sensitivity of the first analyte can be improved, and the concentration level of the analyte in the tissue fluid can be accurately reflected. In some examples, the area ratio of the second analyte enzyme layer 131 to the first analyte enzyme layer 121 may be 1:1.5, 1:1.4, 1:1.3, 1:1.2, 1:1.1. Thus, the concentration level of the analyte in the tissue fluid can be more accurately reflected.

[0084] In some examples, the first analyte enzyme layer 121 and / or the second analyte enzyme layer 131 may be linear. Thus, the utilization rate of the substrate can be improved, and the sensitivity of the sensor can be improved at the same time. Limited by the existing processing level, due to the small volume of the sensor, when a very small analyte enzyme layer is coated on the working electrode of the sensor, it is difficult to directly coat a uniform linear enzyme layer on the electrode surface with the existing processing accuracy. In some examples of the present disclosure, the solution of the enzyme layer can be applied by dripping on the working electrode, and the point solution dripped on the working electrode can be diffused and connected to form a whole linear solution. In this case, compared with the scheme of multi-point discrete arrangement, a larger area of ​​enzyme layer can be arranged on the working electrode of limited area, thereby improving the sensitivity of the sensor 10. In some examples, two adjacent drops of solution can be intersected during dripping, so that a linear analyte enzyme layer can be formed on the working electrode in a way that multiple points are connected in a line. In this case, the processing difficulty of the sensor 10 can be reduced by connecting multiple points in a line, so that a linear analyte enzyme layer can be formed, thereby improving the utilization rate of the substrate.

[0085] In the present disclosure, "the first analyte enzyme layer 121 and / or the second analyte enzyme layer 131 are linear" may include at least three embodiments: the first analyte enzyme layer 121 is linear, the second analyte enzyme layer 131 is linear, and the first analyte enzyme layer 121 and the second analyte enzyme layer 131 are linear. In the present disclosure, other expressions of "A and / or B" have the same meaning as the above-mentioned illustration. In this case, by adjusting the area of ​​the analyte enzyme layer, it can be helpful to improve the sensitivity and accuracy of the sensor 10 when monitoring two analytes.

[0086] Figure 3 is a schematic diagram showing the polymer film layer 14 involved in the example of the present disclosure. Figure 3 In the figure, the illustration of the analyte enzyme layer is omitted.

[0087] In some examples, the first working electrode 12 may have a substrate 140. Thus, the stability of the sensor 10 can be improved. In some examples, the substrate 140 has electrical conductivity. Thus, the current signal generated by the analyte reaction can be conducted. In some examples, the substrate 140 can be made of at least one material selected from gold, glassy carbon, graphite, silver, silver chloride, palladium, titanium, and iridium. In this case, the substrate 140 can have good electrical conductivity and can inhibit the occurrence of electrochemical reactions in the substrate 140, thereby improving the stability of the substrate 140.

[0088] In some examples, the polymer film layer 14 may cover the base layer 140 of the working electrode (see Figure 3 ). Thus, the film manufacturing process of the sensor 10 can be simplified, and the biocompatibility of the sensor 10 can be improved, which can be beneficial to the long-term use of the sensor 10.

[0089] In some examples, the polymer film layer 14 may include a vinylpyridine polymer, thereby facilitating the formation of a polymer film layer 14 that has a limiting effect on the permeation of glucose and ketones.

[0090] In some examples, the mass fraction of vinylpyridine polymers in the polymer film layer 14 is 80% to 100%. As a result, it is possible to form a polymer film layer 14 that has a limiting permeation effect on glucose and ketones and is biocompatible. The amount of glucose and ketones can be controlled within the linear range of the sensor through the polymer film, thereby improving the sensitivity and accuracy of the sensor. In some examples, the mass fraction of vinylpyridine polymers in the polymer film layer 14 is 80%, 81%, 82%, 83%, 84%, 85%, 86%, 87%, 88%, 89%, 90%, 91%, 92%, 93%, 94%, 95%, 96%, 97%, 98%, 99%, 100%. As a result, it is possible to form a stable polymer film layer 14, while also being conducive to enhancing the limiting permeation effect of the polymer film layer 14 on glucose and ketones.

[0091] In some examples, the polymer film layer 14 may further include a cross-linking agent and a regulator. The cross-linking agent may be polyethylene glycol diglycidyl ether. The regulator may be polydimethylsiloxane regulator. Thus, the stability of the polymer film layer 14 may be improved, thereby improving the performance of the sensor 10.

[0092] In some examples, the polymer film layer 14 may include a first film layer 141. The first film layer 141 may be disposed on the enzyme sensing layer. Thus, the penetration of the analyte in the tissue fluid can be limited, while the penetration of other interferents can be prevented, thereby improving the reaction efficiency of the enzyme sensing layer, and further improving the sensitivity and stability of the sensor.

[0093] In some examples, the polymer film layer 14 may include a second film layer 143. The second film layer 143 may be biocompatible. Thus, the occurrence of inflammation caused by implantation of the sensor 10 in the body can be reduced. In some examples, the second film layer 143 may be disposed on the first film layer 141. In this case, a polymer film layer 14 having a permeation limiting effect and a biocompatible effect can be formed, which is beneficial to improving the sensitivity and stability of the sensor 10.

[0094] In some examples, the polymer film layer 14 may further include a transition layer 142. The transition layer 142 may be disposed between the first film layer 141 and the second film layer 143. Thus, a stable polymer film layer 14 may be formed. That is, in some examples, the polymer film layer 14 may include a first film layer 141 disposed sequentially on the enzyme sensing layer, a transition layer 142 formed on the first film layer 141, and a second film layer 143 formed on the transition layer 142 and having biocompatibility. Thus, the bonding stability of the polymer film layer 14 may be improved, thereby facilitating the improvement of the stability of the sensor 10.

[0095] In some examples, the first film layer 141 is formed of a first type of polymer. In some examples, the first type of polymer may include a water-swellable homopolymer. Thus, it can be beneficial to the first film layer. In some examples, the water-swellable homopolymer may be one selected from polystyrene, polyurethane, polyethoxyethyl acrylate, polyethoxypropyl acrylate, poly-2-vinyl pyridine, poly-4-vinyl pyridine, polyhydroxyethyl methacrylate, and polyhydroxyethyl acrylate. Thus, it is possible to form a diffusion-controlled property, so that the analyte concentrations on both sides of the polymer film layer have a fixed concentration ratio, thereby expanding the linear range of the biosensor response.

[0096] In some examples, the second film layer 143 is formed by a second type of polymer. In some examples, the second type of polymer may include a water-soluble polymer. In some examples, the water-soluble polymer may be selected from polyvinyl pyrrolidone, polyvinyl alcohol, chitosan, carboxymethyl chitosan, chitosan salt, alginic acid, alginate, hyaluronic acid, hyaluronate, cellulose ethers, cellulose esters, polyvinyl pyrrolidone, polyacrylamide, polyacrylic acid, polypropylene alcohol, sodium polystyrene sulfonate, polyethylene glycol, polyethylene glycol polypropylene glycol copolymer. Thus, the biocompatibility of the biosensor can be improved.

[0097] In some examples, the transition layer 142 is formed of a third type of polymer. The third type of polymer is a copolymer formed by a first monomer that is the same or similar to the first type of polymer and a second monomer that is the same or similar to the second type of polymer. Thus, the bonding stability and mechanical strength of the polymer film layer 14 can be improved, thereby improving the stability of the sensor 10. Thus, a transition layer with enhanced bonding to the first film layer 141 and the second film layer 143 can be formed, which can be beneficial to the formation of the polymer film layer 14, thereby helping to expand the response linear range of the sensor 10 and improve the biocompatibility of the sensor 10.

[0098] In some examples, the first membrane layer 141 may include a vinylpyridine polymer, thereby forming a membrane layer having a permeation-limiting effect on two analytes.

[0099] In some examples, the transition layer 142 may include a copolymer of vinyl pyridine and styrene, thereby facilitating the formation of a uniform and stable polymer film layer 14, thereby improving the stability of the polymer film layer.

[0100] In some examples, the second film layer 143 may include a copolymer of vinyl pyridine and styrene, which can help improve the film-forming performance of the polymer film layer 14 and further enhance the sensitivity of the sensor 10 .

[0101] In some examples, the second analyte enzyme layer 131 may include hydroxybutyrate dehydrogenase. In this case, hydroxybutyrate in the tissue fluid can react under the action of hydroxybutyrate dehydrogenase, so that the concentration level of hydroxybutyrate can be obtained through the signal generated by the reaction, and then the concentration of ketone bodies in the tissue fluid is mapped with the concentration level of hydroxybutyrate. In some examples, in the second analyte enzyme layer 131, the content of hydroxybutyrate dehydrogenase can be 10% to 20%. For example, the content of hydroxybutyrate dehydrogenase can be 10%, 11%, 12%, 13%, 14%, 15%, 16%, 17%, 18%, 19%, or 20%. Thereby, the oxidation-reduction reaction of ketones in the second analyte enzyme layer 131 can be promoted.

[0102] In some examples, the second analyte enzyme layer 131 may include diaphorase. In this case, the electrons generated by the β-BHB reaction can be transferred by diaphorase. In some examples, in the second analyte enzyme layer 131, the content of diaphorase can be 5% to 20%. For example, the content of diaphorase can be 5%, 6%, 7%, 8%, 9%, 10%, 11%, 12%, 13%, 14%, 15%, 16%, 17%, 18%, 19%, or 20%. Thus, the reaction of the second analyte enzyme layer 131 can be promoted, and the transfer efficiency of electrons in the working electrode can be improved.

[0103] In some examples, the second analyte enzyme layer 131 may include a coenzyme. Thus, the oxidation of hydroxybutyrate dehydrogenase can be promoted, thereby transferring the electrons generated by the β-BHB reaction. In some examples, in the second analyte enzyme layer 131, the content of the coenzyme can be 10% to 30%. For example, the content of the coenzyme can be 10%, 12%, 14%, 15%, 16%, 18%, 19%, 20%, 21%, 22%, 25%, 26%, 27%, 28%, 29%, or 30%. Thus, the redox reaction of hydroxybutyrate dehydrogenase and β-BHB can be promoted.

[0104] In some examples, the second analyte enzyme layer 131 can include an electron mediator. In some examples, the electron mediator can be a redox polymer. In some examples, the redox polymer can be a metal redox polymer. In some examples, the metal redox polymer can be selected from at least one of poly(vinylferrocene), quaternized poly(4-vinylpyridine) of ferrocyanide, quaternized poly(1-vinylimidazole) of ferrocyanide, quaternized poly(4-vinylpyridine) of ferrocyanide, quaternized poly(1-vinylimidazole) of ferrocyanide, quaternized poly(1-vinylimidazole) of ferrocyanide, osmium 2,2'-bipyridine complex coordinated to poly(1-vinylimidazole), osmium 2,2'-bipyridine complex coordinated to poly(4-vinylpyridine), ruthenium 2,2'-bipyridine complex coordinated to poly(1-vinylimidazole), ruthenium 2,2'-bipyridine complex coordinated to poly(4-vinylpyridine), cobalt 2,2'-bipyridine complex coordinated to poly(1-vinylimidazole) or cobalt 2,2'-bipyridine complex coordinated to poly(4-vinylpyridine). Thus, the metal redox polymer can participate in the redox reaction through a covalent bond, a coordination bond or an ionic bond, and can promote the transfer of electrons to the working electrode, thereby generating a current signal of the second analyte concentration. In some examples, in the second analyte enzyme layer 131, the content of the electron mediator can be 10% to 30%. For example, the content of the electron mediator can be 10%, 12%, 14%, 15%, 16%, 18%, 19%, 20%, 21%, 22%, 25%, 26%, 27%, 28%, 29%, or 30%. In this case, the reaction sensitivity of the second analyte enzyme layer 131 to β-BHB can be improved, and by providing a macromolecular electron mediator in the enzyme layer, the hydroxybutyrate dehydrogenase can be covalently bound to the electron mediator, thereby further improving the stability.

[0105] According to the first aspect of the present disclosure, a multi-analyte monitoring sensor can be provided, which can continuously monitor glucose and ketones simultaneously, has a simple manufacturing process, and has high sensitivity and stability.

[0106] The second aspect of the present disclosure relates to a semipermeable membrane of a multi-analyte monitoring sensor. The semipermeable membrane of a multi-analyte monitoring sensor involved in the second aspect of the present disclosure is consistent with the polymer membrane layer 14 in the sensor 10 involved in the first aspect of the present disclosure. The structure of the polymer membrane layer 14 and the setting of its various components and parameters as well as the preparation method can be found in the above description and will not be repeated here.

[0107] In a second aspect of the present disclosure, a semipermeable membrane with high stability and capable of limiting the permeation of both glucose and ketones can be provided. The semipermeable membrane covering the electrodes of the multi-analyte sensor 10 can help improve the sensitivity of the sensor 10 .

[0108] The third aspect of the present disclosure relates to a blood ketone sensor. The blood ketone sensor involved in the third aspect of the present disclosure is basically the same as the sensor 10 involved in the first aspect of the present disclosure, except that the first working electrode and the first analyte enzyme layer disposed on the first working electrode are removed on the basis of the sensor 10 involved in the first aspect, and the second working electrode and the second analyte enzyme layer (i.e., the enzyme layer for detecting ketones) for analyzing ketones are retained. The other parts are consistent with the sensor 10 involved in the first aspect and are not repeated here. In some examples, in the blood ketone sensor, the second working electrode 13 and the reference electrode 15 can be disposed on the same side of the substrate 11, and the counter electrode 16 can be disposed on the other side of the substrate 11. Thus, the utilization rate of the substrate can be improved. The ketone body level in the host can be measured by the blood ketone sensor involved in the third aspect of the present disclosure.

[0109] Hereinafter, the multi-analyte monitoring sensor and the blood ketone sensor provided by the present disclosure will be described in detail in combination with the embodiments and comparative examples, but they should not be construed as limiting the scope of protection of the present disclosure.

[0110] [Blood glucose + blood ketone sensor]

[0111] Example 1

[0112] First, a monitoring probe having front and rear dual working electrodes is prepared, which are a first working electrode and a second working electrode.

[0113] Next, prepare the ketone body sensitive layer reagent components (osmium coordinated metal polymer, β-hydroxybutyrate dehydrogenase, diaphorase, NAD + , stabilizer, cross-linking agent), dissolve each substance in a buffer solvent in a HEPES buffer according to the concentration parameters in the table to obtain a β-hydroxybutyric acid sensitive layer solution.

[0114] Subsequently, a glucose sensitive layer solution was prepared according to the concentration of osmium coordinated metal polymer of 10 mg / mL, glucose oxidase (GOD) of 12.5 mg / mL, and cross-linking agent PEGDGE of 2.5 mg / mL.

[0115] The glucose sensitive layer solution was deposited on the first working electrode, and the β-hydroxybutyric acid sensitive layer solution was deposited on the second working electrode, both forming a linear pattern, in which the glucose sensing layer area was about 0.27 mm 2 The ketone body sensing layer area is about 0.27mm 2 ; That is, the area ratio of blood ketone to blood glucose sensing layer is 1:1; solidification is performed to obtain an electrode containing a dual-component response sensing layer of glucose and ketone bodies.

[0116] Table 1. Ketone body sensitive layer reagent ingredients

[0117] substance Concentration (mg / mL) Metal polymer 10 β-Hydroxybutyrate dehydrogenase 6 Diaphorase 6 NAD+ 6 Human serum albumin 6 PEGDGE 5

[0118] Then, ethanol was prepared as a solvent, and polyvinyl pyridine and PEGDGE were prepared according to Table 2 to obtain a diffusion limiting membrane solution. The diffusion limiting membrane solution was coated on the electrode deposited with the sensitive layer reagent to prepare a dual-component monitoring working electrode containing a membrane layer of about 25 microns.

[0119] Table 2. Diffusion-limiting membrane solution composition

[0120] substance Concentration (mg / mL) Polyvinylpyridine 100 Dimethicone 4 PEGDGE 35

[0121] Prepare 80% ethanol aqueous solution, and mix poly (4-vinylpyridine)-co-polystyrene and PEGDGE in corresponding concentrations according to Table 3 to obtain a compatible membrane solution; apply the compatible membrane solution to the diffusion limiting membrane of the working electrode to prepare a working electrode containing a compatible membrane layer of about 15 microns.

[0122] Table 3. Compatible membrane solution components

[0123]

[0124] Prepare 80% ethanol aqueous solution, and configure (poly 4-vinylpyridine-g-polyethylene glycol)-co-polystyrene and PEGDGE to corresponding concentration solutions according to Table 4 to obtain a biocompatible membrane solution; apply the biocompatible membrane solution to the compatible membrane of the working electrode to prepare a biocompatible membrane layer containing about 15 microns, and obtain a working electrode with a total film thickness of about 56 microns.

[0125] Table 4. Biocompatible membrane solution composition

[0126]

[0127] The sensor of Example 1 was immersed in a standard PBS buffer (pH 7.4) at 37°C. A working voltage was applied until a constant background current was reached, and the response current was measured at the same time containing corresponding concentrations of glucose / sodium β-hydroxybutyrate. The concentrations were the following values, 2.2mM / 1mM, 5mM / 2mM, 10mM / 3mM, 15mM / 4mM, 20mM / 6mM, and 25mM / 8mM, respectively, with the former being the glucose concentration and the latter being the sodium β-hydroxybutyrate concentration.

[0128] Figure 4 is a graph showing the response current-test concentration curve of Example 1 of the blood glucose+blood ketone sensor of the present disclosure. Figure 4 As shown, the response of the dual-component sensor of Example 1 to glucose and sodium β-hydroxybutyrate shows good linearity. Specifically, it can be seen that the response current linear curve R of the sensor to glucose and sodium β-hydroxybutyrate is 2Both are greater than 0.996, indicating a good linear relationship. The sensitivity of the glucose working electrode is 0.79nA / mM, and the sensitivity of the ketone body working electrode is 2.35nA / mM.

[0129] Example 2

[0130] First, a monitoring probe having front and rear dual working electrodes is prepared, which are a first working electrode and a second working electrode.

[0131] Secondly, the β-hydroxybutyric acid sensitive layer solution and the glucose sensitive layer solution were prepared in the same formula as in Example 1, wherein the metal polymer was a ruthenium coordinated metal polymer. The glucose sensitive layer solution was deposited on the first working electrode, and the β-hydroxybutyric acid sensitive layer solution was deposited on the second working electrode, both forming a linear pattern, wherein the glucose sensing layer area was about 0.3 mm 2 The ketone body sensing layer area is about 0.27mm 2 ; Solidification is performed to obtain an electrode containing a dual-component responsive sensing layer of glucose and ketone bodies.

[0132] The diffusion limiting membrane, the compatibility membrane and the biocompatible membrane were coated in the same manner as in Example 1 to obtain the sensor of Example 2 having a blood glucose to blood ketone sensing layer area ratio of 1.11:1.

[0133] Figure 5 is a picture showing a monitoring probe of Example 2 of the blood glucose + blood ketone sensor according to the present disclosure.

[0134] The sensor of Example 2 was immersed in a standard PBS buffer (pH 7.4) at 37°C. A working voltage was applied until a constant background current was reached, and the response current was measured simultaneously containing corresponding concentrations of glucose / sodium β-hydroxybutyrate. The concentrations were the following values, 2.2mM / 1mM, 5mM / 2mM, 10mM / 3mM, 15mM / 4mM, 20mM / 6mM, and 25mM / 8mM, respectively, with the former being the glucose concentration and the latter being the sodium β-hydroxybutyrate concentration.

[0135] Figure 6 is a graph showing a response current-test time curve of Example 2 of the blood glucose+blood ketone sensor according to the present disclosure. Figure 7 2 is a graph showing the response current-test concentration curve of Example 2 of the blood glucose+blood ketone sensor of the present disclosure. Figure 6 , Figure 7 As shown, the dual-component sensor of Example 2 exhibits good linearity in response to glucose and sodium β-hydroxybutyrate. Specifically, Figure 7 As shown, the linear response current curve R of the sensor to glucose and sodium β-hydroxybutyrate 2Both are greater than 0.99, indicating a good linear relationship. The sensitivity of the glucose working electrode is 1.10 nA / mM, and the sensitivity of the ketone body working electrode is 2.43 nA / mM.

[0136] The sensor of Example 2 was immersed in standard PBS buffer (pH 7.4) at 37°C for 336 hours (14 days). Figure 8 2 is a graph showing the glucose / sodium β-hydroxybutyrate response current-time curve of Example 2 of the blood glucose+blood ketone sensor of the present disclosure. Figure 8 As shown, during the test period from day 0 to day 14, both current signals of the sensor remained stable.

[0137] Example 3

[0138] A monitoring probe with dual working electrodes was prepared in the same manner as in Example 1. The glucose sensitive layer solution was deposited on the first working electrode, and the β-hydroxybutyric acid sensitive layer solution was deposited on the second working electrode, both forming a linear pattern, wherein the glucose sensing layer area was approximately 0.39 mm 2 The ketone body sensing layer area is about 0.26mm 2 ; That is, the area ratio of blood ketone to blood glucose sensing layer is 1:1.5; solidification is performed to obtain an electrode containing a dual-component response sensing layer of glucose and ketone bodies.

[0139] Next, the diffusion limiting membrane, the compatibility membrane and the biocompatible membrane were coated in the same manner as in Example 1 to obtain the sensor of Example 3 having a blood ketone to blood glucose sensing layer area ratio of 1:1.5.

[0140] The sensor of Example 3 was immersed in a standard PBS buffer (pH 7.4) at 37°C. A working voltage was applied until a constant background current was reached, and the response current was measured at the same time containing corresponding concentrations of glucose / sodium β-hydroxybutyrate. The concentrations were the following values, 2.2mM / 1mM, 5mM / 2mM, 10mM / 3mM, 15mM / 4mM, 20mM / 6mM, and 25mM / 8mM, respectively, with the former being the glucose concentration and the latter being the sodium β-hydroxybutyrate concentration.

[0141] Fig. 9 is a graph showing the response current-test concentration curve of Example 3 of the blood glucose+blood ketone sensor of the present disclosure. Fig. 9 As shown, the response of the dual-component sensor of Example 3 to glucose and sodium β-hydroxybutyrate shows good linearity. Specifically, the response current linear curve R 2 Both are greater than 0.997, indicating a good linear relationship. The sensitivity of the glucose working electrode is 1.22nA / mM, and the sensitivity of the ketone body working electrode is 2.52nA / mM.

[0142] Comparative Example 1

[0143] First, a monitoring probe having front and rear dual working electrodes is prepared, which are a first working electrode and a second working electrode.

[0144] Secondly, a β-hydroxybutyric acid sensitive layer solution and a glucose sensitive layer solution were prepared in the same formula as in Example 1, and the glucose sensitive layer solution was deposited on the first working electrode, and the β-hydroxybutyric acid sensitive layer solution was deposited on the second working electrode, both forming a linear pattern, wherein the glucose sensing layer area was about 0.3 mm 2 The ketone body sensing layer area is about 0.45mm 2 , that is, the area ratio of the glucose enzyme layer to the ketone body enzyme layer is 1:1.5; solidification is performed to obtain an electrode containing a dual-component response sensing layer of glucose and ketone bodies.

[0145] Then, the two-component response electrode obtained above was coated according to the same three-layer film recipe as in Example 1, and finally a two-component sensor with a film thickness of about 57 microns was obtained.

[0146] Fig.10 is a picture showing a monitoring probe of Comparative Example 1 of the blood glucose + blood ketone sensor according to the present disclosure.

[0147] The sensor of Comparative Example 1 was immersed in a standard PBS buffer (pH 7.4) at 37° C.; a working voltage was applied relative to the reference electrode until a constant background current was reached, and the response current was measured at the corresponding concentrations of glucose / sodium β-hydroxybutyrate, and the concentrations were the following values, 2.2mM / 1mM, 5mM / 2mM, 10mM / 3mM, 15mM / 4mM, 20mM / 6mM, 25mM / 8mM, the former being the glucose concentration and the latter being the sodium β-hydroxybutyrate concentration;

[0148] Fig.11 is a response current-test concentration curve graph of Comparative Example 1 of the blood glucose+blood ketone sensor according to the present disclosure.

[0149] The dual-component sensor of Comparative Example 1 exhibited good linearity in the current response to the concentration of sodium β-hydroxybutyrate, but poor in the current response to the concentration of glucose. Fig.11 As shown, the linear response current curve R of the sensor to glucose and sodium β-hydroxybutyrate 2 They are 0.877 and 0.995 respectively, which indicates that the glucose sensor is not linear in the range of 0-25mM. This may be because when responding to high concentrations, the amount of ketone body reaction is large, and the product will affect the response of the working electrode on the glucose side, resulting in the non-linear problem of high glucose concentration.

[0150] Comparative Example 2

[0151] First, a monitoring probe having front and rear dual working electrodes is prepared, which are a first working electrode and a second working electrode.

[0152] Secondly, the β-hydroxybutyric acid sensitive layer solution and the glucose sensitive layer solution were prepared in the same formula as in Example 1. The glucose sensitive layer solution was deposited on the first working electrode, and the β-hydroxybutyric acid sensitive layer solution was deposited on the second working electrode, both forming a linear pattern, wherein the glucose sensing layer area was about 0.3 mm 2 The ketone body sensing layer area is about 0.27mm 2 ; Solidification is performed to obtain an electrode containing a dual-component responsive sensing layer of glucose and ketone bodies.

[0153] Then, 80% ethanol aqueous solution was prepared as solvent, and polyvinylpyridine-b-polystyrene and PEGDGE were prepared into alcohol solutions with corresponding concentrations according to Table 5 to obtain a membrane solution. The membrane solution was coated on the electrode deposited with the sensitive layer reagent to prepare a dual-component monitoring working electrode containing a membrane layer of about 60 microns.

[0154] Table 5. Diffusion-limiting membrane solution composition

[0155] substance Concentration (mg / mL) Polyvinylpyridine 100 Dimethicone 4 PEGDGE 35

[0156] Fig.12 is a picture showing the monitoring probe of Comparative Example 2 of the blood glucose + blood ketone sensor according to the present disclosure.

[0157] The sensor of comparative example 2 was immersed in a standard PBS buffer (pH 7.4) at 37°C; a working voltage was applied until a constant background current was reached, and the response current was measured at the corresponding concentrations of glucose / sodium β-hydroxybutyrate. The concentrations were as follows: 2.2mM / 1mM, 5mM / 2mM, 10mM / 3mM, 15mM / 4mM, 20mM / 6mM, 25mM / 8mM, the former being the glucose concentration and the latter being the sodium β-hydroxybutyrate concentration.

[0158] Fig.13 2 is a graph showing the response current-test time curve of Comparative Example 2 of the blood glucose+blood ketone sensor of the present disclosure. Fig.13 As shown, the response of the dual-component sensor of Comparative Example 2 to glucose and sodium β-hydroxybutyrate shows good linearity. Specifically, the response current linear curve R 2All of them are greater than 0.99, indicating a good linear relationship; among them, the sensitivity of the glucose working electrode is 0.30nA / mM, and the sensitivity of the ketone body working electrode is 0.73nA / mM. However, compared with Example 2, the sensitivity of the two electrodes in Comparative Example 2 is lower, indicating that adding a compatible film layer and a biocompatible film layer can significantly improve the sensitivity of the electrode.

[0159] Comparative Example 3

[0160] First, a monitoring probe having front and rear dual working electrodes is prepared, which are a first working electrode and a second working electrode.

[0161] Secondly, the β-hydroxybutyric acid sensitive layer solution and the glucose sensitive layer solution were prepared in the same formula as in Example 1. The glucose sensitive layer solution was deposited on the first working electrode, and the β-hydroxybutyric acid sensitive layer solution was deposited on the second working electrode, both forming a linear pattern, wherein the glucose sensing layer area was about 0.3 mm 2 The ketone body sensing layer area is about 0.27mm 2 ; Solidification is performed to obtain an electrode containing a dual-component responsive sensing layer of glucose and ketone bodies.

[0162] Then, ethanol was prepared as a solvent, and polyvinyl pyridine and PEGDGE were prepared into alcohol solutions of corresponding concentrations according to Table 6 to obtain a diffusion limiting membrane solution. The diffusion limiting membrane solution was coated on the electrode deposited with the sensitive layer reagent to prepare a dual-component monitoring working electrode containing a membrane layer of about 35 microns.

[0163] Table 6 Composition of diffusion limiting membrane solution

[0164] substance Concentration (mg / mL) Polyvinylpyridine 100 Dimethicone 4 PEGDGE 35

[0165] Finally, prepare an 80% ethanol aqueous solution, and configure (poly-4-vinylpyridine-g-polyethylene glycol)-co-polystyrene and PEGDGE to corresponding concentration solutions according to Table 7 to obtain a biocompatible membrane solution; apply the biocompatible membrane solution to the diffusion limiting membrane of the working electrode to prepare a working electrode containing a biocompatible membrane layer of about 25 microns and a total film thickness of about 60 microns.

[0166] Table 7 Biocompatible membrane solution composition

[0167]

[0168] Fig.14 is a picture showing the monitoring probe of Comparative Example 3 of the blood glucose + blood ketone sensor according to the present disclosure. Fig.15 2 is a picture showing the monitoring probe of Comparative Example 3 of the blood glucose + blood ketone sensor according to the present disclosure from another perspective.

[0169] like Fig.14 , Fig.15 As shown, it can be seen that in the absence of a compatible membrane layer, the semipermeable membrane of the prepared sensor is wrinkled and has uneven thickness from top to bottom, which is not conducive to the preparation of a highly consistent sensor.

[0170] [Blood Ketone Sensor]

[0171] Example 1, Comparative Examples 1-2

[0172] First, a monitoring probe with a working electrode is prepared.

[0173] Next, the sensitive layer reagent raw materials of each blood ketone sensor embodiment and comparative example were prepared according to Table 8 (ruthenium coordinated metal polymer, β-hydroxybutyrate dehydrogenase, diaphorase, NAD + , stabilizer, cross-linking agent, etc., dissolve each substance in the corresponding solvent buffer solvent according to the concentration parameters in the table, and ultrasonically vibrate to completely dissolve them to obtain the β-hydroxybutyric acid sensitive layer solution of each embodiment.

[0174] The β-hydroxybutyric acid sensitive layer solution is deposited on the working electrode to form a linear pattern with a width of about 160 microns and a length of about 2.1 mm; the solution is solidified to obtain an enzyme sensing layer.

[0175] Table 8. Examples of blood ketone sensors

[0176]

[0177] Fig.16 1 is a picture showing the monitoring probes of the sensors of Example 1 and Comparative Examples 1-2 of the blood ketone sensor according to the present disclosure.

[0178] Then, pure ethanol was prepared as a solvent, and polyvinylpyridine, polydimethylsiloxane, and PEGDGE were prepared into alcohol solutions with corresponding concentrations according to Table 9 to obtain a diffusion limiting membrane solution. The diffusion limiting membrane solution was coated on the working electrode on which the β-hydroxybutyric acid sensitive layer solution was deposited to prepare a working electrode containing a diffusion limiting membrane layer of about 20 microns.

[0179] Table 9. Diffusion-limiting membrane solution composition

[0180] substance Concentration (mg / mL) Polyvinylpyridine 92 Dimethicone 3 PEGDGE 30

[0181] Prepare 80% ethanol aqueous solution, and mix poly (4-vinylpyridine)-co-polystyrene and PEGDGE into corresponding concentration solutions according to Table 10 to obtain a compatible membrane solution; coat the working electrode coated with the diffusion limiting layer membrane with the compatible membrane solution to prepare a working electrode containing a compatible membrane layer of about 15 microns.

[0182] Table 10. Compatible membrane solution components

[0183]

[0184] Prepare 80% ethanol aqueous solution, and configure (poly 4-vinylpyridine-g-polyethylene glycol)-co-polystyrene and PEGDGE to corresponding concentration solutions according to Table 11 to obtain a biocompatible membrane solution; coat the working electrode coated with the compatible membrane layer with the biocompatible membrane solution to prepare a working electrode containing a biocompatible membrane layer of about 15 microns and a total film thickness of 50 microns.

[0185] Table 11. Biocompatible membrane solution composition

[0186]

[0187] Fig.17 1 is a picture showing monitoring probes coated with a polymer film layer in Example 1 and Comparative Examples 1-2 of the blood ketone sensor according to the present disclosure.

[0188] Finally, the linearity and stability tests of the ketone body biosensors obtained in Example 1 and Comparative Examples 1-2 were performed:

[0189] (1) Linearity test:

[0190] The ketone body biosensor was immersed in a standard PBS buffer (pH 7.4) at 37°C; a working voltage was applied until the ketone body biosensor reached a constant background current, and sodium β-hydroxybutyrate at concentrations of 0mM, 1mM, 2mM, 3mM, 4mM, 6mM, and 8mM was added to the solution to measure the linearity of the reaction. After each addition of sodium β-hydroxybutyrate, the solution was allowed to equilibrate for 10 minutes, and the solution was continuously stirred during the measurement process to make its concentration uniform.

[0191] Fig.18 is a graph showing a response current-test concentration of a linearity test of Example 1 of the blood ketone sensor according to the present disclosure. Fig.19 1 is a graph showing a response current-test concentration curve of a linearity test of Example 1 of the blood ketone sensor according to the present disclosure.

[0192] like Fig.18 , 19 As shown, the ketone body biosensor of blood ketone sensor Example 1 exhibits good linearity in response to glucose. Specifically, the glucose sensor in Example 1 has a good linear correlation between the concentration of sodium β-hydroxybutyrate and the response current in the range of 0 mM to 8 mM.

[0193] (2) Stability test:

[0194] The ketone body biosensor was immersed in 3 mM sodium β-hydroxybutyrate (PBS buffer pH 7.4 as solvent) at 37°C and measured continuously for 14 days.

[0195] Fig. 20 1 is a graph showing a response current-test concentration curve of a stability test of Example 1 of the blood ketone sensor according to the present disclosure. Fig.21 1 is a graph showing a response current-test concentration curve of a stability test of Comparative Example 1 of the blood ketone sensor according to the present disclosure. Fig. 22 3 is a graph showing a response current-test concentration curve of a stability test of Comparative Example 2 of the blood ketone sensor according to the present disclosure.

[0196] like Figure 20-22 As shown, comparative example 1 cannot maintain stability for 14 days at 37°C, and performance degradation begins on the 5th day, while both example 1 and comparative example 2 can maintain stability for 14 days because stabilizers are added to the enzyme sensing layer, achieving longer stability of the ketone body sensor. The response value of example 1 is higher than that of comparative example 2, indicating that using HEPES buffer as the solvent of the sensitive layer reagent is better than using PBS buffer as the solvent of the sensitive layer reagent.

[0197] Comparative Example 3

[0198] A monitoring probe having a working electrode was prepared in the same manner as in Example 1 for blood ketone, and an enzyme sensing layer was disposed on the working electrode to obtain a sensor without a semipermeable membrane.

[0199] The sensor of Comparative Example 3 was exposed to a solution containing 1 mM sodium β-hydroxybutyrate in a standard PBS buffer (pH 7.4) at 37° C. under energy for 1 hour.

[0200] Fig.23 3 is a graph showing a response current-test concentration of a linearity test of Comparative Example 3 of the blood ketone sensor according to the present disclosure.

[0201] like Fig.23 As shown in the figure, within the 1-hour test time, the current signal generated by the sensor gradually decreased from 240nA to below 155nA, and the stability was poor. The test results show that the sensor without a semipermeable membrane has poor stability when testing the concentration of sodium β-hydroxybutyrate solution.

[0202] Comparative Example 4

[0203] A monitoring probe having a working electrode was prepared in the same manner as in the blood ketone comparative example 4, and an enzyme sensing layer was disposed on the working electrode to obtain a sensor without a semipermeable membrane.

[0204] Prepare 80% ethanol aqueous solution, and mix poly (4-vinylpyridine)-co-polystyrene and PEGDGE according to the table below to prepare corresponding concentration solutions to obtain an inner membrane solution; coat the inner membrane solution on the working electrode provided with the enzyme sensing layer to prepare a working electrode containing about 35 microns of poly (4-vinylpyridine)-b-polystyrene membrane layer.

[0205] Table 12 Inner membrane solution composition

[0206]

[0207]

[0208] Fig.24 is a picture showing a working electrode of Comparative Example 4 of the blood ketone sensor according to the present disclosure.

[0209] Prepare 80% ethanol aqueous solution, and configure (poly 4-vinylpyridine-g-polyethylene glycol)-co-polystyrene and PEGDGE into corresponding concentration solutions according to Table 13 to obtain an outer layer membrane solution; coat the working electrode coated with the inner layer membrane with the outer layer membrane solution to prepare a working electrode containing an outer layer membrane of about 15 microns and a total membrane thickness of 50 microns.

[0210] Table 13 Outer membrane solution composition

[0211]

[0212] Fig.25 is a picture showing a monitoring probe of a sensor of Comparative Example 4 of the blood ketone sensor according to the present disclosure.

[0213] Finally, the obtained ketone body sensor was subjected to a stability test; that is, the comparative example 4 ketone body biosensor was immersed in 3 mM sodium β-hydroxybutyrate (PBS buffer pH 7.4 as solvent) at 37° C.; and the measurement was continued for 3 days.

[0214] Fig.26 4 is a graph showing a response current-test time curve of the sensor of Comparative Example 4 of the blood ketone sensor according to the present disclosure.

[0215] like Fig.26 As shown, the initial response current of the sensor of Comparative Example 4 is 15.6 nA. After three days, the sensor response current decreases to 11.2 nA, a total decrease of 28.2%, and an average daily decrease of 9.4% / day. This indicates that the sensor coated only with poly (4-vinylpyridine)-b-polystyrene and (poly (4-vinylpyridine)-g-polyethylene glycol)-co-polystyrene film has poor stability.

[0216] The main reason is that compared with blood ketone example 1, the permeability of sodium β-hydroxybutyrate of this membrane layer combination is larger, indicating that the diffusion restriction effect of the membrane is not strong, and it is unable to reduce the permeability of sodium β-hydroxybutyrate, and the small molecule escape effect of the restriction enzyme layer leads to its poor stability.

[0217] Comparative Example 5

[0218] First, a monitoring probe with a working electrode is prepared.

[0219] The corresponding reagents were prepared according to the same enzyme layer formula as in Example 1 of blood ketone, and the β-hydroxybutyric acid sensitive layer solution was drop-coated on the working electrode using an ultra-micro pipetting robot to form 6 independent dot patterns with a dot diameter of 150 μm; the solution was cured to obtain an enzyme sensor having 6 dot patterns.

[0220] Fig. 27 is a picture showing a working electrode of the sensor of Comparative Example 5 of the blood ketone sensor according to the present disclosure.

[0221] Then, the 6-dot pattern enzyme sensor was coated with a film layer according to the same film layer coating scheme as in Example 1 of blood ketones to obtain the ketone body sensor of Comparative Example 5.

[0222] Finally, the obtained ketone body sensor of comparative example 5 was subjected to a linearity test: the ketone body biosensor was immersed in a standard PBS buffer (pH 7.4) at 37°C; a working voltage was applied until the ketone body biosensor reached a constant background current, and sodium β-hydroxybutyrate at concentrations of 1mM, 2mM, 3mM, 4mM, 5mM, 6mM, 7mM, and 8mM was added to the solution to measure the linearity of the reaction. After each addition of sodium β-hydroxybutyrate, the solution was continuously stirred during the measurement process to make its concentration uniform.

[0223] Fig.28 3 is a graph showing a response current-test concentration of a linearity test of a blood ketone sensor comparative example 5 according to the present disclosure.

[0224] like Fig.28 As shown, it can be seen that its response to sodium β-hydroxybutyrate shows good linearity in the concentration range of 0-8mM, but its response sensitivity is low, that is, the slope of the linear fitting curve is 0.677nA / mM, but the sensitivity of the sensor in Example 1 is 2.08nA / mM, and that of Comparative Example 5 is only 32.5% of that in Example 1. This is because the area of ​​the enzyme layer in Comparative Example 5 is small, resulting in a low reaction area with the sodium β-hydroxybutyrate that permeates the membrane layer, which indicates that increasing the area of ​​the enzyme layer can effectively increase the sensitivity of the sensor.

[0225] Various examples of the present disclosure are described above in the detailed description. Although these descriptions directly describe the above examples, it should be understood that modifications and / or variations of the specific examples shown and described herein may be conceived by those skilled in the art. Any such modifications or variations that fall within the scope of this specification are also intended to be included therein. Unless otherwise specified, it is the inventor's intention that the terms in the specification and claims are given the ordinary and customary meanings of ordinary technicians.

Claims

1. A multi-analyte monitoring sensor, characterized in that: The sensor comprises a substrate, a working electrode disposed on the substrate, an enzyme sensing layer disposed on the working electrode, and a polymer film layer disposed on the enzyme sensing layer. The working electrode comprises a first working electrode and a second working electrode; The enzyme sensing layer comprises a first analyte enzyme layer and a second analyte enzyme layer, the first analyte enzyme layer is disposed on the first working electrode, the second analyte enzyme layer is disposed on the second working electrode, the first analyte is glucose, and the second analyte is ketone; The polymer membrane layer is permeable to the first analyte and the second analyte and covers the first analyte enzyme layer and the second analyte enzyme layer. The polymer membrane layer is configured to simultaneously permeate glucose and ketones, and the thickness ratio of the polymer membrane layer is adjusted so that the polymer membrane layer has different permeability coefficients. The area of ​​the first analyte enzyme layer is greater than the area of ​​the second analyte enzyme layer. The polymer membrane layer includes a first membrane layer sequentially arranged on the enzyme sensing layer, a transition layer formed on the first membrane layer, and a second membrane layer formed on the transition layer and having biocompatibility. The first membrane layer includes a vinyl pyridine polymer, and the mass fraction of the vinyl pyridine polymer in the polymer membrane layer is 80% to 100%. The transition layer and the second membrane layer include a copolymer of vinyl pyridine and styrene.

2. The multi-analyte monitoring sensor according to claim 1, characterized in that: An area ratio of the first analyte enzyme layer to the second analyte enzyme layer is 1.11:1 to 1.5:

1.

3. The multi-analyte monitoring sensor according to claim 1, characterized in that: The first analyte enzyme layer and / or the second analyte enzyme layer are in a linear shape.

4. The multi-analyte monitoring sensor according to claim 1, characterized in that: The second analyte enzyme layer includes hydroxybutyrate dehydrogenase, diaphorase, coenzyme, and electron mediator. In the second analyte enzyme layer, the mass fraction of hydroxybutyrate dehydrogenase is 10% to 20%, the mass fraction of diaphorase is 5% to 20%, the mass fraction of coenzyme is 10% to 30%, and the mass fraction of the electron mediator is 10% to 30%.

5. The multi-analyte monitoring sensor according to claim 1, characterized in that: The sensor further includes a reference electrode and a counter electrode. The first working electrode and the second working electrode are respectively located on two sides of the substrate. The reference electrode and the counter electrode are respectively located on two sides of the substrate.

Images