Multi-analyte sensor

Abstract

The present disclosure describes a multi-analyte sensor including a working electrode, an enzyme sensing layer disposed on the working electrode, and a polymer film layer covering the enzyme sensing layer, the working electrode including a first working electrode for detecting a first analyte and a second working electrode for detecting a second analyte, the first analyte being glucose, the second analyte being glucose, and the polymer film layer covering the enzyme sensing layer. The second analyte is beta-hydroxybutyric acid; the enzyme sensing layer comprises 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, and the area of the first analyte enzyme layer is larger than or equal to that of the second analyte enzyme layer; the polymer film layer is configured to be permeable to the first analyte and the second analyte at the same time, and has different permeability coefficients for the first analyte and the second analyte. As a result, it is possible to provide a sensor for continuously monitoring blood glucose and blood ketones, which has high sensitivity and a simple process.

Description

[0001] This application is a divisional application of the patent application with the application date of July 31, 2023, the application number of 2023109568404, and the invention title of Multi - analyte Monitoring Sensor. Technical Field

[0002] The present disclosure generally relates to the field of biosensors, and more particularly to a multi - analyte sensor. Background Art

[0003] A biosensor is an analytical device that tightly combines biological materials, bio - derived materials, or bio - mimetic materials with optical, electrochemical, temperature, piezoelectric, magnetic, or micro - mechanical physicochemical sensors or sensing microsystems. It can usually be used to quickly detect certain specific chemical substances in the human body, such as glucose, ketone bodies, uric acid, and a series of amino acid compounds.

[0004] When blood sugar is too high, the body lacks insulin, or a ketogenic diet is carried out, the body will metabolize fat to produce ketone bodies. When the amount of ketone bodies exceeds the metabolic capacity of the liver, it is easy to cause complications of diabetes. Therefore, the monitoring of blood ketone content also has great significance in diabetic patients.

[0005] Traditional blood ketone detection methods mainly include urine ketone test strip method, blood ketone test strip 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, a wearable acetone sensor detects acetone in the exhaled gas of diabetic patients, converts the resistance change into a digital quantity, and sends it to the display module to display the acetone concentration.

[0006] However, the response time of the above - mentioned acetone sensor is usually about 50s, and the evaluation of acetone usually has a lag; moreover, when it is necessary to simultaneously monitor blood sugar and blood ketone, generally, another blood sugar sensor needs to be set up to monitor blood sugar, which is inconvenient to use. Summary of the Invention

[0007] The present disclosure is proposed in view of the above - mentioned situation, and its purpose is to provide a multi - analyte monitoring sensor that can continuously monitor blood sugar and blood ketone simultaneously.

[0008] To this end, the present disclosure provides a multi - analyte monitoring sensor. The sensor includes 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 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 disposed on the first working electrode, and 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 film layer is permeable to the first analyte and the second analyte and covers the first analyte enzyme layer and the second analyte enzyme layer.

[0009] In the multi - analyte monitoring sensor involved in the present disclosure, by providing enzyme layers of two analytes on two working electrodes of the sensor, the concentration levels of two different analytes in the interstitial fluid can be detected respectively; the polymer film layer can simultaneously permeate glucose and ketone. By covering the first analyte enzyme layer and the second analyte enzyme layer with the polymer film layer, it can play a role in restricting the permeation of the two analytes, and at the same time is beneficial to improving the stability of the sensor; in addition, due to the small volume of the biosensor, the precision requirement for coating the surface of the sensor is relatively high. Usually, the dip - coating method is used to cover the surface of the sensor with a film solution to form a film layer covering the sensor. Compared with the scheme of respectively providing semi - permeable membranes permeable to glucose and ketone on the first analyte enzyme layer and the second analyte enzyme layer, the polymer film in the present invention can simultaneously permeate glucose and ketone, 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 continuously monitoring glucose and ketone simultaneously can be provided.

[0010] In addition, in the multi - analyte monitoring sensor involved in the present disclosure, optionally, the polymer film layer includes a vinylpyridine - type polymer, a cross - linker, and a regulator. The mass fraction of the vinylpyridine - type polymer in the polymer film layer is 80% to 100%. Thus, it is beneficial to form a polymer film layer that has a restricting permeation effect on glucose and ketone and has biocompatibility. Through the polymer film, the amounts of glucose and ketone can be controlled within the linear range of the sensor, improving the sensitivity and accuracy of the sensor.

[0011] In addition, in the multi - analyte monitoring sensor involved in the present disclosure, optionally, the polymer film layer includes a first film layer disposed on the enzyme sensing layer in sequence, 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 enhancing the stability of the sensor.

[0012] In addition, in the multi-analyte monitoring sensor according to 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. The third type of polymer is a copolymer formed from 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. Thereby, the adhesion stability and mechanical strength of the polymer film layer can be improved, and further the stability of the sensor can be improved.

[0013] In addition, in the multi-analyte monitoring sensor according to the present disclosure, optionally, the first film layer includes a vinylpyridine polymer, and the transition layer and the second film layer include a copolymer of vinylpyridine and styrene. Thereby, it is beneficial to improve the film-forming performance of the film layer, and further improve the sensitivity of the sensor.

[0014] In addition, in the multi-analyte monitoring sensor according to the present disclosure, optionally, the area of the first analyte enzyme layer is larger than the area of the second analyte enzyme layer. Since the amounts of glucose and ketone that can permeate through the polymer film are different, and usually the amount of glucose is less than that of ketone, 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 as much as possible, making the currents 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.

[0015] In addition, in the multi-analyte monitoring sensor according to the present disclosure, optionally, the area ratio of the first analyte enzyme layer to the second analyte enzyme layer is from 1:1 to 1.5:1. Thereby, the reaction sensitivity of the first analyte can be improved, and further the concentration level of the analyte in the interstitial fluid can be more accurately reflected.

[0016] In addition, in the multi-analyte monitoring sensor according to the present disclosure, optionally, the first analyte enzyme layer and / or the second analyte enzyme layer is linear. In this case, compared with the multi-point discrete arrangement scheme, a larger area of the enzyme layer can be arranged on the working electrode with a limited area, thereby improving the sensitivity of the sensor.

[0017] In addition, in the multi - analyte monitoring sensor according to the present disclosure, optionally, the second analyte enzyme layer includes hydroxybutyrate dehydrogenase, diaphorase, coenzyme, and an 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%. In this case, the reaction sensitivity of the second analyte enzyme layer to β - BHB can be improved. At the same time, by setting a macromolecular electron mediator in the enzyme layer, hydroxybutyrate dehydrogenase can be covalently bonded to the electron mediator, thereby further improving stability.

[0018] In addition, in the multi - analyte monitoring sensor according to 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 dual - path detection sensor can be formed, and setting the four electrodes in pairs on both sides of the substrate can improve the utilization rate of the substrate and further reduce the interference of the electric field on the sensor system.

[0019] According to the present disclosure, a multi - analyte monitoring sensor can be provided, which can continuously monitor glucose and ketone simultaneously, has a simple manufacturing process, and has high sensitivity and stability. BRIEF DESCRIPTION OF THE DRAWINGS

[0020] The present disclosure will now be further explained in detail only by way of examples with reference to the accompanying drawings, where:

[0021] Figure 1 is a schematic diagram showing the multi - analyte monitoring sensor according to the example of the present disclosure.

[0022] Figure 2A is a front - view schematic diagram showing the multi - analyte monitoring sensor according to the example of the present disclosure.

[0023] Figure 2B is a back - view schematic diagram showing the multi - analyte monitoring sensor according to the example of the present disclosure.

[0024] Figure 3 is a schematic diagram showing the polymer film layer according to the example of the present disclosure.

[0025] Figure 4 is a graph showing the response current - test concentration curve of Example 1 of the blood glucose + blood ketone sensor according to the present disclosure.

[0026] Figure 5 is a picture showing the monitoring probe of Example 2 of the blood glucose + blood ketone sensor according to the present disclosure.

[0027] Figure 6 It is a graph showing the response current - test time of Example 2 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0028] Figure 7 It is a graph showing the response current - test concentration of Example 2 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0029] Figure 8 It is a graph showing the glucose / sodium β - hydroxybutyrate response current - time of Example 2 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0030] Figure 9 It is a graph showing the response current - test concentration of Example 3 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0031] Figure 10 It is a picture showing the monitoring probe of Comparative Example 1 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0032] Figure 11 It is a graph showing the response current - test concentration of Comparative Example 1 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0033] Figure 12 It is a picture showing the monitoring probe of Comparative Example 2 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0034] Figure 13 It is a graph showing the response current - test time of Comparative Example 2 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0035] Figure 14 It is a picture showing the monitoring probe of Comparative Example 3 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0036] Figure 15 It is a picture showing another perspective of the monitoring probe of Comparative Example 3 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0037] Figure 16 It is a picture showing the monitoring probes of Example 1 of the blood ketone sensor and Comparative Examples 1 - 2 involved in the present disclosure.

[0038] Figure 17 It is a picture showing the monitoring probes coated with a polymer film layer of Example 1 of the blood ketone sensor and Comparative Examples 1 - 2 involved in the present disclosure.

[0039] Figure 18 It is a graph showing the response current - test concentration of the linear test of Example 1 of the blood ketone sensor involved in the present disclosure.

[0040] Figure 19 It is a graph of response current - test concentration showing the linear test of Example 1 of the blood ketone sensor involved in the present disclosure.

[0041] Figure 20 It is a graph of response current - test concentration showing the stability test of Example 1 of the blood ketone sensor involved in the present disclosure.

[0042] Figure 21 It is a graph of response current - test concentration showing the stability test of Comparative Example 1 of the blood ketone sensor involved in the present disclosure.

[0043] Figure 22 It is a graph of response current - test concentration showing the stability test of Comparative Example 2 of the blood ketone sensor involved in the present disclosure.

[0044] Figure 23 It is a graph of response current - test concentration showing the linear test of Comparative Example 3 of the blood ketone sensor involved in the present disclosure.

[0045] Figure 24 It is a picture of the working electrode of Comparative Example 4 of the blood ketone sensor involved in the present disclosure.

[0046] Figure 25 It is a picture of the monitoring probe of the sensor of Comparative Example 4 of the blood ketone sensor involved in the present disclosure.

[0047] Figure 26 It is a graph of response current - test time showing the response of the sensor of Comparative Example 4 of the blood ketone sensor involved in the present disclosure.

[0048] Figure 27 It is a picture of the working electrode of the sensor of Comparative Example 5 of the blood ketone sensor involved in the present disclosure.

[0049] Figure 28 It is a graph of response current - test concentration showing the linear test of Comparative Example 5 of the blood ketone sensor involved in the present disclosure.

[0050] Explanation of reference numerals:

[0051] 10... Monitoring probe, 11... Substrate, 12... First working electrode, 121... First analyte enzyme layer, 13... Second working electrode, 131... Second analyte enzyme layer, 14... Semi - permeable membrane, 141... First membrane layer, 142... Transition layer, 143... Second membrane layer, 15... Reference electrode, 16... Counter electrode Detailed description of the invention

[0052] Hereinafter, with reference to the accompanying drawings, preferred embodiments of the present disclosure will be described in detail. In the following description, the same reference numerals are given to the same components, and repeated descriptions are omitted. In addition, the drawings are only schematic diagrams, and the proportional relationship of the sizes between components or the shapes of components may be different from the actual ones.

[0053] It should be noted that the terms "comprising" and "having" in the present disclosure and any variations thereof, for example, a process, method, system, product or device including or having a series of steps or units, do not necessarily limit to those steps or units clearly listed, but may include or have other steps or units not clearly listed or inherent to these processes, methods, products or devices.

[0054] 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, and they only serve as a reading prompt. Such subheadings should neither be understood as used to divide the content of the article, nor should the content under the subheadings be limited only within the scope of the subheadings.

[0055] On the one hand, the present disclosure relates to a multi - analyte monitoring sensor, on the second hand, to a semi - permeable membrane of a multi - analyte monitoring sensor, and on the third hand, to a blood ketone sensor.

[0056] In the present disclosure, the multi - analyte monitoring sensor may also be referred to as a "multi - analyte sensor", a "continuous glucose and blood ketone monitoring sensor", a "glucose and blood ketone sensor", a "glucose + blood ketone sensor", a "physiological parameter sensor" or a "sensor", a "sensing probe", a "sensing head", an "implantation probe", etc.

[0057] The sensor involved in the present disclosure can be used to monitor the 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 body.

[0058] In the present disclosure, the sensor can also be used to detect analytes in the host body. The analytes can be chemical substances in body fluids. For example, the analytes 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 bodies, lactate, oxygen, peroxide, prostate - specific antigen, prothrombin, RNA, thyroid - stimulating hormone or troponin. In addition, the analytes can also be drugs in body fluids. For example, the analytes can be digitoxin, digoxin, theophylline, warfarin or antibiotics (such as gentamicin, vancomycin, etc.).

[0059] Hereinafter, with reference to the accompanying drawings, the multi - analyte monitoring sensor related to the first aspect of the present disclosure will be described.

[0060] Figure 1 It is a schematic diagram showing a multi - analyte monitoring sensor involved in the examples of the present disclosure.

[0061] In this embodiment, the sensor 10 can be used to obtain 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 through the monitoring device 1 applied to itself. In some examples, the sensor 10 can be assembled on the body surface parts such as the host's arm, back, abdomen, waist, leg, etc. where the sensor 10 can be partially implanted to monitor physiological parameters. In some examples, the sensor 10 can be used to implant under the host's skin to obtain sensing signals (such as current signals). In some examples, when the sensor 10 is applied to the host, it can contact the tissue fluid or blood in the host's body to measure the analyte levels in the tissue fluid or blood.

[0062] In some examples, the sensor 10 includes a substrate 11 and working electrodes disposed on the substrate. In some examples, the number of working electrodes can be multiple, and each working electrode can be used to detect different analytes. Thus, the sensor 10 can monitor multiple analytes.

[0063] In some examples, the working electrodes can include a first working electrode 12 and a second working electrode 13. The first working electrode 12 can be used to detect a first analyte, and the second working electrode 13 can be used to detect a second analyte. Thus, the sensor 10 can monitor two analytes.

[0064] In some examples, the first analyte can be any one of glucose, urea, uric acid, ketone bodies, creatinine, ethanol, and lactic acid. In some examples, the second analyte can be any one of glucose, urea, uric acid, ketone bodies, creatinine, ethanol, and lactic acid. Among them, ketone bodies (which can be abbreviated as ketones) are the collective names of acetoacetic acid, β - hydroxybutyric acid, and acetone, which are intermediate products of liver fatty acid oxidation and decomposition. 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's body can be monitored to obtain parameters related to the ketone body level.

[0065] In some examples, the first analyte can be glucose. Thus, the glucose concentration in the body can be monitored. In some examples, the second analyte can be ketone bodies. Thus, the content of ketone bodies in the body can be monitored. It should be noted that, as described above, ketone bodies are the collective names of acetoacetic acid, β - hydroxybutyric acid, and acetone, which are intermediate products of liver fatty acid oxidation and decomposition. Therefore, the second analyte can be any one of acetoacetic acid, β - hydroxybutyric acid, and acetone. Preferably, the second analyte can be β - hydroxybutyric acid.

[0066] Hereinafter, taking the first analyte as glucose and the second analyte as ketone as an example, the sensor 10 according to the present disclosure will be described in detail.

[0067] In some examples, the sensor 10 may include an enzyme sensing layer disposed on the 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 the enzyme layers of two analytes on the two working electrodes of the sensor 10, the concentration levels of two different analytes in the interstitial fluid can be detected respectively.

[0068] 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 play a role in restricting the permeation of the two analytes, while being beneficial to improving the stability of the sensor. Compared with the solution of separately disposing semi-permeable membranes 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 be permeable to the first analyte and the second analyte simultaneously, thereby being able to simplify 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 be permeable to glucose and ketone simultaneously. 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 polymer film layer 14 can play a role in restricting the permeation of the two analytes, while being beneficial to improving the stability of the sensor; in addition, since the volume of the biosensor 10 is small, the precision requirement for coating the film on the surface of the sensor 10 is relatively high, and usually the dip-coating method is used to cover the surface of the sensor 10 with a film solution to form a film layer covering the sensor 10. Compared with the solution of separately disposing semi-permeable membranes permeable to 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 be permeable to glucose and ketone simultaneously, thereby being able to simplify the manufacturing process of the sensor 10 while maintaining the high sensitivity and accuracy of the sensor 10. Thus, a sensor 10 with a simple manufacturing process and capable of continuously monitoring glucose and ketone simultaneously can be provided.

[0069] In some examples, the permeability coefficient of the polymer film layer 14 for glucose can be from 1 to 1000. Preferably, the permeability coefficient of the polymer film layer 14 for glucose can be from 300 to 800. For example, the permeability coefficient of the polymer film layer 14 for glucose can be 300, 400, 500, 600, 700, or 800. In some examples, the permeability coefficient of the polymer film layer 14 for ketone can be from 1 to 1000. Preferably, the permeability coefficient of the polymer film layer 14 for ketone can be from 200 to 600. For example, the permeability coefficient of the polymer film layer 14 for ketone can be 200, 300, 400, 500, or 600. Herein, the permeability coefficient refers to the concentration difference inside and outside 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 that of glucose inside the polymer film layer.

[0070] In some examples, the thickness ratio of the polymer film layer 14 can be adjusted to enable the polymer film layer 14 to have different permeability coefficients. In this case, by adjusting the permeability coefficient of the polymer film layer 14, the amount of the analyte passing through can be controlled, making the current generated by the working electrode more accurate, and thus enabling the sensor 10 to be more sensitive.

[0071] Figure 2A is a front schematic view of the multi - analyte monitoring sensor involved in the examples of the present disclosure. Figure 2B is a back schematic view of the multi - analyte monitoring sensor involved in the examples of the present disclosure.

[0072] 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. The sensor 10 can be partially implanted under the host's skin to obtain a sensing signal indicating the analyte level, and then transmit the sensing signal to the electronic system; the electronic system can process and / or re - transmit the sensing signal. In this case, the concentration of the analyte in the body can be accurately obtained.

[0073] In some examples, the sensor 10 may further 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 can also form a set of three - electrode circuits with the reference electrode 15 and the counter electrode 16.

[0074] In some examples, the four-electrode dual-channel detection sensor may include a first working electrode 12, a second working electrode 13, a reference electrode 15, and a counter electrode 16. 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, thereby reducing the length of the substrate, and further reducing the implantation depth of the sensor 10. Since the greater the implantation depth, the more obvious the pain generated, thus, the pain caused by the implantation of the sensor 10 can be reduced, and the wearing comfort can be improved.

[0075] In some examples, the first working electrode 12 and the second working electrode 13 may be respectively located on both sides of the substrate 11, and the reference electrode 15 and the counter electrode 16 are respectively located on both sides of the substrate 11. In this case, arranging the four electrodes in pairs on both sides of the substrate can improve the utilization rate of the substrate. At the same time, the voltage drop of the sensor electrodes is optimized, the error can be reduced, and the interference of the electric field on the sensor system can be further reduced. In some examples, since the wearing comfort is related to the implantation depth, the longer the substrate, the greater the implantation depth, and the pain during wearing may also increase. Therefore, arranging the four electrodes in pairs on both sides of the substrate can improve the utilization rate of the substrate. By arranging four electrodes within the limited substrate space, the length of the substrate can be reduced, thereby improving the wearing comfort.

[0076] In some examples, the reference electrode 15 may 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 through the potential difference formed by the reference electrode 15 and the working electrode. Thus, 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 concentration level information of the analyte in the tissue fluid or blood. In some examples, the number of reference electrodes 15 can be one or more, such as 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 as to obtain an accurate current signal, and further improve the sensitivity of the sensor.

[0077] In some examples, the sensor 10 can be divided into an implantation part and a connection part. The implantation part can be placed subcutaneously in a host and connected to an electronic system through the connection part. In this case, by placing the implantation part subcutaneously in the host and electrically connecting the implantation part to the electronic system through the connection part, thereby, it is possible to facilitate the transmission of the sensed signal obtained by the implantation part to the electronic system. In some examples, the implantation part can be rigid. Thereby, it can help to be placed subcutaneously in the host. In other examples, the implantation part can also be flexible. In this case, the foreign body sensation of the host can be reduced.

[0078] In some examples, as described above, the sensor 10 can include a first analyte enzyme layer 121. In this case, the first analyte can react in the first analyte enzyme layer 121 to generate an electric current, and a concentration signal of the first analyte can be obtained.

[0079] In some examples, the first analyte enzyme layer 121 can include a glucose enzyme. Thereby, glucose can react in the first analyte enzyme layer to generate a concentration signal of glucose. In some examples, in the first analyte enzyme layer 121, the content (mass fraction) of the glucose enzyme can be 40% to 70%. For example, the content (mass fraction) of the glucose enzyme 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%. Thereby, it can be beneficial for glucose to react.

[0080] In some examples, the first analyte enzyme layer 121 may include an electron mediator. In some examples, the electron mediator may be a redox polymer. In some examples, the redox polymer may be a metal redox polymer. In some examples, the metal redox polymer may be selected from at least one of poly(vinylferrocene), quaternized poly(4-vinylpyridine) ferricyanide, quaternized poly(1-vinylimidazole) ferricyanide, quaternized poly(4-vinylpyridine) ferrocyanide, quaternized poly(1-vinylimidazole) 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 redox reactions through covalent bonds, coordination bonds or ionic bonds.

[0081] In some examples, in the first analyte enzyme layer 121, the content (mass fraction) of the electron mediator may be 10% to 40%. For example, the content (mass fraction) of the electron mediator may 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 reaction of the first analyte in the first analyte enzyme layer can be promoted.

[0082] In some examples, the first analyte enzyme layer 121 may include a crosslinking agent. In some examples, the crosslinking agent may be PEGDGE. In some examples, the crosslinking agent may be PEGDGE400. In some examples, in the first analyte enzyme layer 121, the content (mass fraction) of the crosslinking agent may be 0% to 20%. For example, the content (mass fraction) of the crosslinking agent 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, long-term monitoring by the sensor can be facilitated.

[0083] In some examples, the area of the first analyte enzyme layer 121 can be larger than the area of the second analyte enzyme layer 131. Since the amounts of glucose and ketone that can permeate through the polymer film 14 are different, and usually the amount of glucose is less than that of ketone, 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 minimizing 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.

[0084] In some examples, the area ratio of the first analyte enzyme layer 121 to the second analyte enzyme layer 131 can be greater than 1:1. Thus, it is beneficial 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 can be below 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 can be from 1:1 to 1.5:1. Thus, the reaction sensitivity of the first analyte can be improved, and then the concentration level of the analyte in the interstitial 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 can 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 interstitial fluid can be more accurately reflected.

[0085] In some examples, the first analyte enzyme layer 121 and / or the second analyte enzyme layer 131 can be linear. Thus, the utilization rate of the substrate can be increased, and at the same time, the sensitivity of the sensor can be improved. Limited by the existing processing level, since the sensor has a small volume, when coating an extremely small analyte enzyme layer 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 coated on the working electrode by drop coating, and the dot-like solutions dropped on the working electrode can diffuse and connect to form a whole linear solution. In this case, compared with the multi-point discrete arrangement scheme, a larger area of the enzyme layer can be arranged on the working electrode with a limited area, thereby improving the sensitivity of the sensor 10. In some examples, when drop coating, the adjacent two drops of solution can intersect, so that a linear analyte enzyme layer can be formed on the working electrode in a way that multiple points are connected into a line. In this case, the processing difficulty of the sensor 10 can be reduced by connecting multiple points into a line, so that a linear analyte enzyme layer can be formed, and then the utilization rate of the substrate can be increased.

[0086] In the present disclosure, the statement that "the first analyte enzyme layer 121 and / or the second analyte enzyme layer 131 is linear" may at least include three embodiments: the first analyte enzyme layer 121 is linear, the second analyte enzyme layer 131 is linear, and both 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 illustration. In this case, by adjusting the area of the analyte enzyme layer, it is beneficial to improve the sensitivity and accuracy of the sensor 10 when monitoring two analytes.

[0087] Figure 3 is a schematic diagram showing the polymer film layer 14 involved in the examples of the present disclosure. In Figure 3 the schematic illustration of the analyte enzyme layer is omitted.

[0088] In some examples, the first working electrode 12 may have a base layer 140. Thereby, the stability of the sensor 10 can be improved. In some examples, the base layer 140 has electrical conductivity. Thereby, the current signal generated by the analyte reaction can be conducted. In some examples, the base layer 140 may be made of at least one material selected from gold, glassy carbon, graphite, silver, silver chloride, palladium, titanium, and iridium. In this case, the base layer 140 can have good electrical conductivity and can inhibit the occurrence of electrochemical reactions on the base layer 140, thereby improving the stability of the base layer 140.

[0089] In some examples, the polymer film layer 14 may cover the base layer 140 of the working electrode (see Figure 3 ). Thereby, the film manufacturing process of the sensor 10 can be simplified, and at the same time, the biocompatibility of the sensor 10 can be improved, which is beneficial to the long-term use of the sensor 10.

[0090] In some examples, the polymer film layer 14 may include a vinylpyridine polymer. Thereby, it is beneficial to form a polymer film layer 14 that has a restricted permeation effect on glucose and ketones.

[0091] In some examples, the mass fraction of the vinylpyridine polymer in the polymer film layer 14 is 80% to 100%. Thereby, it is beneficial to form the polymer film layer 14 that has a restricted permeation effect on glucose and ketones and is biocompatible. Through the polymer film, the amounts of glucose and ketones can be controlled within the linear range of the sensor, improving the sensitivity and accuracy of the sensor. In some examples, the mass fraction of the vinylpyridine polymer 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%. Thereby, it is beneficial to form a stable polymer film layer 14, and at the same time, it is beneficial to enhance the restricted permeation effect of the polymer film layer 14 on glucose and ketones.

[0092] In some examples, the polymer film layer 14 may further include a crosslinking agent and a regulator. The crosslinking agent can be polyethylene glycol diglycidyl ether. The regulator can be a polydimethylsiloxane regulator. Thereby, the stability of the polymer film layer 14 can be enhanced, and further, the performance of the sensor 10 can be improved.

[0093] In some examples, the polymer film layer 14 may include a first film layer 141. The first film layer 141 can be disposed on the enzyme sensing layer. Thereby, the permeation of the analyte in the tissue fluid can be restricted, and at the same time, other interfering substances can be blocked from permeating, improving the reaction efficiency of the enzyme sensing layer, and further, the sensitivity and stability of the sensor can be enhanced.

[0094] In some examples, the polymer film layer 14 may include a second film layer 143. The second film layer 143 can have biocompatibility. Thereby, the occurrence of inflammatory phenomena caused by implanting the sensor 10 into the body can be reduced. In some examples, the second film layer 143 can be disposed on the first film layer 141. In this case, a polymer film layer 14 with restricted permeation and biocompatibility effects can be formed, and further, it is beneficial to enhance the sensitivity and stability of the sensor 10.

[0095] In some examples, the polymer film layer 14 may further include a transition layer 142. The transition layer 142 can be disposed between the first film layer 141 and the second film layer 143. Thereby, a stable polymer film layer 14 can be formed. That is to say, in some examples, the polymer film layer 14 may include a first film layer 141 disposed on the enzyme sensing layer in sequence, 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. Thereby, the bonding stability of the polymer film layer 14 can be improved, which is beneficial to enhancing the stability of the sensor 10.

[0096] 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. Thereby, it is beneficial to the first film layer. In some examples, the water-swellable homopolymer may be one selected from polystyrene, polyurethane, ethoxyethyl polyacrylate, ethoxypropyl polyacrylate, poly(2-vinylpyridine), poly(4-vinylpyridine), hydroxyethyl polymethacrylate, and hydroxyethyl polyacrylate. Thereby, a film layer with diffusion control performance can be formed, such that there is a fixed concentration ratio of the analyte on both sides of the polymer film layer, thereby being able to expand the linear response range of the biosensor.

[0097] In some examples, the second film layer 143 is formed of 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 one selected from polyvinylpyrrolidone, polyvinyl alcohol, chitosan, carboxymethyl chitosan, chitosan salts, alginic acid, alginates, hyaluronic acid, hyaluronates, cellulose ethers, cellulose esters, polyvinylpyrrolidone, polyacrylamide, polyacrylic acid, polyvinyl alcohol, sodium polystyrene sulfonate, polyethylene glycol, and polyethylene glycol-polypropylene glycol copolymers. Thereby, the biocompatibility of the biosensor can be improved.

[0098] In some examples, the transition layer 142 is formed of a third type of polymer. Among them, the third type of polymer is a copolymer formed from 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. Thereby, the adhesion stability and mechanical strength of the polymer film layer 14 can be improved, and further the stability of the sensor 10 can be improved. Thereby, a transition layer with enhanced adhesion to the first film layer 141 and the second film layer 143 can be formed, thereby being beneficial to the formation of the polymer film layer 14, and further being able to help expand the linear response range of the sensor 10 and improve the biocompatibility of the sensor 10.

[0099] In some examples, the first film layer 141 may include a vinylpyridine polymer. Thereby, a film layer with a restricted permeation effect on two analytes can be formed.

[0100] In some examples, the transition layer 142 may include a copolymer of vinylpyridine and styrene. Thereby, it is beneficial to form a uniform and stable polymer film layer 14, and further improve the stability of the polymer film layer.

[0101] In some examples, the second film layer 143 may include a copolymer of vinylpyridine and styrene. Thereby, it is beneficial to improve the film-forming performance of the polymer film layer 14, and further improve the sensitivity of the sensor 10.

[0102] In some examples, the second analyte enzyme layer 131 may include hydroxybutyrate dehydrogenase. In this case, hydroxybutyric acid in the interstitial fluid can react under the action of hydroxybutyrate dehydrogenase, so that the concentration level of hydroxybutyric acid can be obtained through the signal generated by the reaction, and then the concentration of ketone bodies in the interstitial fluid can be mapped with the concentration level of hydroxybutyric acid. In some examples, in the second analyte enzyme layer 131, the content of hydroxybutyrate dehydrogenase may be 10% to 20%. For example, the content of hydroxybutyrate dehydrogenase may be 10%, 11%, 12%, 13%, 14%, 15%, 16%, 17%, 18%, 19%, or 20%. Thereby, the redox reaction of ketone in the second analyte enzyme layer 131 can be promoted.

[0103] In some examples, the second analyte enzyme layer 131 may include diaphorase. In this case, the electrons generated by the reaction of β-BHB can be transferred through diaphorase. In some examples, in the second analyte enzyme layer 131, the content of diaphorase may be 5% to 20%. For example, the content of diaphorase may be 5%, 6%, 7%, 8%, 9%, 10%, 11%, 12%, 13%, 14%, 15%, 16%, 17%, 18%, 19%, or 20%. Thereby, the reaction of the second analyte enzyme layer 131 can be promoted, and at the same time, the electron transfer efficiency in the working electrode can be improved.

[0104] In some examples, the second analyte enzyme layer 131 may include coenzyme. Thereby, the oxidation of hydroxybutyrate dehydrogenase can be promoted, so that the electrons generated by the reaction of β-BHB can be transferred. In some examples, in the second analyte enzyme layer 131, the content of coenzyme may be 10% to 30%. For example, the content of coenzyme may be 10%, 12%, 14%, 15%, 16%, 18%, 19%, 20%, 21%, 22%, 25%, 26%, 27%, 28%, 29%, or 30%. Thereby, the redox reaction between hydroxybutyrate dehydrogenase and β-BHB can be promoted.

[0105] In some examples, the second analyte enzyme layer 131 may include an electron mediator. In some examples, the electron mediator may be a redox polymer. In some examples, the redox polymer may be a metal redox polymer. In some examples, the metal redox polymer may be selected from at least one of poly(vinylferrocene), quaternized poly(4-vinylpyridine) ferricyanide, quaternized poly(1-vinylimidazole) ferricyanide, quaternized poly(4-vinylpyridine) ferrocyanide, quaternized poly(1-vinylimidazole) 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 redox reactions through covalent bonds, coordination bonds or ionic bonds, 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 may be 10% to 30%. For example, the content of the electron mediator may 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 at the same time, by setting a macromolecular electron mediator in the enzyme layer, hydroxybutyrate dehydrogenase can be covalently bound to the electron mediator, thereby further improving the stability.

[0106] 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.

[0107] The second aspect of the present disclosure relates to a semi-permeable membrane of a multi-analyte monitoring sensor. The semi-permeable membrane of a multi-analyte monitoring sensor related to the second aspect of the present disclosure is the same as the polymer membrane layer 14 in the sensor 10 related to the first aspect of the present disclosure. For the structure of the polymer membrane layer 14, the settings of its various components and parameters, and the preparation method, reference can be made to the above description, and details will not be repeated here.

[0108] In the second aspect of the present disclosure, a semi-permeable membrane with high stability and capable of restricting the permeation of glucose and ketones simultaneously can be provided. Covering the electrodes of the multi-analyte sensor 10 with this semi-permeable membrane can help improve the sensitivity of the sensor 10.

[0109] The third aspect of the present disclosure relates to a blood ketone sensor. The blood ketone sensor related to the third aspect of the present disclosure is basically the same as the sensor 10 related to the first aspect of the present disclosure, except that the first working electrode and the first analyte enzyme layer provided on the first working electrode in the sensor 10 related to the first aspect are removed, and the second working electrode and the second analyte enzyme layer for analyzing ketone (i.e., the enzyme layer for detecting ketone) are retained, and other parts are the same as the sensor 10 related to the first aspect, which will not be elaborated herein. In some examples, in the blood ketone sensor, the second working electrode 13 and the reference electrode 15 can be provided on the same side of the substrate 11, and the counter electrode 16 can be provided on the other side of the substrate 11. Thus, the utilization rate of the substrate can be improved. Through the blood ketone sensor related to the third aspect of the present disclosure, the ketone body level in the host body can be measured.

[0110] Hereinafter, the multi-analyte monitoring sensor and the blood ketone sensor provided by the present disclosure will be described in detail with reference to examples and comparative examples, but they should not be construed as limiting the protection scope of the present disclosure.

[0111] [Blood Glucose + Blood Ketone Sensor]

[0112] Example 1

[0113] First, prepare a monitoring probe with front and rear dual working electrodes, namely the first working electrode and the second working electrode.

[0114] Secondly, prepare the reagent components of the ketone body sensitive layer (osmium coordination metal polymer, β-hydroxybutyric acid dehydrogenase, diaphorase, NAD + , stabilizer, crosslinking agent) according to Table 1, and dissolve each substance in the buffer solvent of HEPES buffer solution according to the concentration parameters in the table to obtain the β-hydroxybutyric acid sensitive layer solution.

[0115] Subsequently, prepare a glucose sensitive layer solution according to the concentration of osmium coordination metal polymer being 10 mg / mL, glucose oxidase (GOD) being 12.5 mg / mL, and crosslinking agent PEGDGE being 2.5 mg / mL.

[0116] Deposit the glucose sensitive layer solution on the first working electrode, and deposit the β-hydroxybutyric acid sensitive layer solution on the second working electrode, both forming a linear pattern, where the glucose sensing layer area is about 0.27 mm 2 , and the ketone body sensing layer area is about 0.27 mm 2 ; that is, the area ratio of the blood ketone to the blood glucose sensing layer is 1:1; perform curing to obtain an electrode with a glucose and ketone body dual-component responsive sensing layer.

[0117] Table 1. Reagent Components of the Ketone Body Sensitive Layer

[0118] Substance Concentration (mg / mL) Metal polymer 10 β-Hydroxybutyric dehydrogenase 6 Diaphorase 6 NAD+ 6 Human serum albumin 6 PEGDGE 5

[0119] Then, ethanol was prepared as a solvent, and a diffusion-limiting membrane solution was obtained by configuring polyvinylpyridine and PEGDGE according to Table 2. The diffusion-limiting membrane solution was coated onto the electrode deposited with the sensitive layer reagent to fabricate a two-component monitoring working electrode with a membrane layer of about 25 μm.

[0120] Table 2. Composition of the diffusion-limiting membrane solution

[0121] Substance Concentration (mg / mL) Polyvinylpyridine 100 Polydimethylsiloxane 4 PEGDGE 35

[0122] An 80% ethanol aqueous solution was prepared, and a compatibility membrane solution was obtained by configuring poly(4-vinylpyridine-co-polystyrene) and PEGDGE at corresponding concentrations according to Table 3; the compatibility membrane solution was coated onto the diffusion-limiting membrane of the working electrode to fabricate a working electrode with a compatibility membrane layer of about 15 μm.

[0123] Table 3. Composition of the compatibility membrane solution

[0124]

[0125] An 80% ethanol aqueous solution was prepared, and a biocompatible membrane solution was obtained by configuring (poly(4-vinylpyridine-g-polyethylene glycol)-co-polystyrene) and PEGDGE at corresponding concentrations according to Table 4; the biocompatible membrane solution was coated onto the compatibility membrane of the working electrode to fabricate a working electrode with a biocompatible membrane layer of about 15 μm, resulting in a working electrode with a total membrane thickness of about 56 μm.

[0126] Table 4. Composition of the biocompatible membrane solution

[0127]

[0128] The sensor of Example 1 was immersed in a standard PBS buffer solution (pH 7.4) at 37°C; a working voltage was applied until a constant background current was reached, and the response current at corresponding concentrations of glucose / β-hydroxybutyrate sodium was measured. The concentrations were as follows: 2.2 mM / 1 mM, 5 mM / 2 mM, 10 mM / 3 mM, 15 mM / 4 mM, 20 mM / 6 mM, 25 mM / 8 mM, with the former being the glucose concentration and the latter being the β-hydroxybutyrate sodium concentration.

[0129] Figure 4 It shows the response current - test concentration curve of Example 1 of the blood glucose + blood ketone sensor involved in the present disclosure. As Figure 4 shown, the two-component sensor of Example 1 showed good linearity in the response to glucose and β-hydroxybutyrate sodium. Specifically, it can be seen that the linear curve R of the response current of its sensor to glucose and β-hydroxybutyrate sodium 2All are greater than 0.996, indicating a good linear relationship. The sensitivity of the glucose working electrode is 0.79 nA / mM, and the sensitivity of the ketone body working electrode is 2.35 nA / mM.

[0130] Example 2

[0131] First, prepare a monitoring probe with front and rear dual working electrodes, namely the first working electrode and the second working electrode respectively.

[0132] Secondly, prepare the β-hydroxybutyric acid sensitive layer solution and the glucose sensitive layer solution with the same formulation as in Example 1, where the metal polymer is a ruthenium coordination metal polymer. Deposit the glucose sensitive layer solution on the first working electrode and the β-hydroxybutyric acid sensitive layer solution on the second working electrode, both forming a linear pattern. Among them, the area of the glucose sensing layer is about 0.3 mm 2 , and the area of the ketone body sensing layer is about 0.27 mm 2 ; Carry out curing to obtain an electrode with a glucose and ketone body dual-component responsive sensing layer.

[0133] Coat the diffusion-limiting membrane, compatibility membrane and biocompatibility membrane in the same manner as in Example 1 to obtain the sensor of Example 2 with a blood glucose to blood ketone sensing layer area ratio of 1.11:1.

[0134] Figure 5 is a picture showing the monitoring probe of Example 2 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0135] Immerse the sensor of Example 2 in a standard PBS buffer solution (pH 7.4) at 37°C; apply a working voltage until a constant background current is reached, and measure the response current at corresponding concentrations of glucose / sodium β-hydroxybutyrate. The respective concentrations are the following values: 2.2 mM / 1 mM, 5 mM / 2 mM, 10 mM / 3 mM, 15 mM / 4 mM, 20 mM / 6 mM, 25 mM / 8 mM, with the former being the glucose concentration and the latter being the sodium β-hydroxybutyrate concentration.

[0136] Figure 6 is a graph showing the response current - test time curve of Example 2 of the blood glucose + blood ketone sensor involved in the present disclosure. Figure 7 is a graph showing the response current - test concentration curve of Example 2 of the blood glucose + blood ketone sensor involved in the present disclosure. As Figure 6 , Figure 7 shown, the dual-component sensor of Example 2 shows good linearity in the response to glucose and sodium β-hydroxybutyrate. Specifically, as Figure 7 shown, the response current linear curves R of its sensor to glucose and sodium β-hydroxybutyrate 2All 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.

[0137] Immerse the sensor of Example 2 in a standard PBS buffer solution (pH 7.4) at 37 °C for 336 hours (14 days). Figure 8 It shows the glucose / β-hydroxybutyrate response current-time curve of Example 2 of the blood glucose + blood ketone sensor involved in the present disclosure. As Figure 8 shown, during the test time from day 0 to day 14, both current signals of the sensor remained in a stable state.

[0138] Example 3

[0139] Prepare a monitoring probe with dual working electrodes in the same manner as in Example 1. Deposit the glucose-sensitive layer solution on the first working electrode and deposit the β-hydroxybutyric acid-sensitive layer solution on the second working electrode, both forming a linear pattern, where the area of the glucose sensing layer is about 0.39 mm 2 , and the area of the ketone body sensing layer is about 0.26 mm 2 ; that is, the area ratio of blood ketone to blood glucose sensing layer is 1:1.5; perform curing to obtain an electrode with a dual-component response sensing layer containing glucose and ketone body.

[0140] Next, coat the diffusion-limiting membrane, compatibility membrane, and biocompatible membrane in the same manner as in Example 1 to obtain the sensor of Example 3 with an area ratio of blood ketone to blood glucose sensing layer of 1:1.5.

[0141] Immerse the sensor of Example 3 in a standard PBS buffer solution (pH 7.4) at 37 °C; apply a working voltage until a constant background current is reached, and measure the response current at corresponding concentrations of glucose / β-hydroxybutyrate. Each concentration is the following value, 2.2 mM / 1 mM, 5 mM / 2 mM, 10 mM / 3 mM, 15 mM / 4 mM, 20 mM / 6 mM, 25 mM / 8 mM, with the former being the glucose concentration and the latter being the β-hydroxybutyrate concentration.

[0142] Figure 9 It shows the response current-test concentration curve of Example 3 of the blood glucose + blood ketone sensor involved in the present disclosure. As Figure 9 shown, the dual-component sensor of Example 3 shows good linearity in the response to glucose and β-hydroxybutyrate. Specifically, the linear curves R of the response current of its sensor to glucose and β-hydroxybutyrate 2 are both greater than 0.997, indicating a good linear relationship. The sensitivity of the glucose working electrode is 1.22 nA / mM, and the sensitivity of the ketone body working electrode is 2.52 nA / mM.

[0143] Comparative Example 1

[0144] First, prepare a monitoring probe with front and rear dual working electrodes, namely the first working electrode and the second working electrode respectively.

[0145] Secondly, prepare the β-hydroxybutyric acid sensitive layer solution and the glucose sensitive layer solution with the same formulation as in Example 1. Deposit the glucose sensitive layer solution on the first working electrode and deposit the β-hydroxybutyric acid sensitive layer solution on the second working electrode, both forming a linear pattern. Among them, the area of the glucose sensing layer is about 0.3 mm 2 , and the area of the ketone body sensing layer is about 0.45 mm 2 , that is, the area ratio of the glucose enzyme layer to the ketone body enzyme layer is 1:1.5; carry out curing to obtain a dual-component response sensing layer electrode containing glucose and ketone bodies.

[0146] Then coat the above-obtained dual-component response electrode according to the same three film layer formulations as in Example 1, and finally obtain a dual-component sensor with a film layer thickness of about 57 microns.

[0147] Figure 10 It is a picture showing the monitoring probe of Comparative Example 1 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0148] Immerse the sensor of Comparative Example 1 in a standard PBS buffer solution (pH 7.4) at 37°C; apply a working voltage to the relative reference electrode until a constant background current is reached, and measure the response current at the corresponding concentrations of glucose / sodium β-hydroxybutyrate. Each concentration is the following value, 2.2 mM / 1 mM, 5 mM / 2 mM, 10 mM / 3 mM, 15 mM / 4 mM, 20 mM / 6 mM, 25 mM / 8 mM, with the former being the glucose concentration and the latter being the sodium β-hydroxybutyrate concentration;

[0149] Figure 11 It is a graph showing the response current - test concentration curve of Comparative Example 1 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0150] The current response of the dual-component sensor in Comparative Example 1 to the concentration of sodium β-hydroxybutyrate shows good linearity, while the current response to the concentration of glucose is poor. Specifically, as Figure 11 shown, the linear curves R of the response current of its sensor to glucose and sodium β-hydroxybutyrate 2 are 0.877 and 0.995 respectively, which indicates that the glucose sensor does not have linearity in the range of 0 - 25 mM. This may be because when the high-concentration response occurs, the reaction amount on the ketone body side is relatively large, and the product will affect the response of the working electrode on the glucose side, resulting in the non-linearity problem of glucose at high concentrations.

[0151] Comparative Example 2

[0152] First, prepare a monitoring probe with front and rear dual working electrodes, namely the first working electrode and the second working electrode respectively.

[0153] Secondly, prepare the β-hydroxybutyric acid sensitive layer solution and the glucose sensitive layer solution with the same formulation as in Example 1. Deposit the glucose sensitive layer solution on the first working electrode and deposit the β-hydroxybutyric acid sensitive layer solution on the second working electrode, both forming a linear pattern, where the area of the glucose sensing layer is about 0.3 mm 2 , and the area of the ketone body sensing layer is about 0.27 mm 2 ; Carry out curing to obtain an electrode with a glucose and ketone body dual-component responsive sensing layer.

[0154] Then, prepare an 80% ethanol aqueous solution as the solvent, and configure the corresponding concentration of alcohol solutions of polyvinylpyridine-b-polystyrene and PEGDGE according to Table 5 to obtain a membrane solution. Coat the membrane solution onto the electrode deposited with the sensitive layer reagent to make a dual-component monitoring working electrode with a membrane layer of about 60 microns.

[0155] Table 5. Composition of the diffusion-limiting membrane solution

[0156] Substance Concentration (mg / mL) Polyvinylpyridine 100 Polydimethylsiloxane 4 PEGDGE 35

[0157] Figure 12 is a picture showing the monitoring probe of Comparative Example 2 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0158] Immerse the sensor of Comparative Example 2 in a standard PBS buffer solution (pH 7.4) at 37°C; apply a working voltage until a constant background current is reached, and measure the response current at the corresponding concentrations of glucose / sodium β-hydroxybutyrate. The respective concentrations are the following values: 2.2 mM / 1 mM, 5 mM / 2 mM, 10 mM / 3 mM, 15 mM / 4 mM, 20 mM / 6 mM, 25 mM / 8 mM, with the former being the glucose concentration and the latter being the sodium β-hydroxybutyrate concentration.

[0159] Figure 13 is a graph showing the response current - test time curve of Comparative Example 2 of the blood glucose + blood ketone sensor involved in the present disclosure. As Figure 13 shown, the dual-component sensor of Comparative Example 2 shows good linearity in the response to glucose and sodium β-hydroxybutyrate. Specifically, the linear curves of the response current of its sensor to glucose and sodium β-hydroxybutyrate are R 2All are greater than 0.99, indicating a good linear relationship; among them, the sensitivity of the glucose working electrode is 0.30 nA / mM, and the sensitivity of the ketone body working electrode is 0.73 nA / mM. However, compared with Example 2, the sensitivity of the two electrodes in Comparative Example 2 is lower, indicating that adding a compatibility film layer and a biocompatible film layer can significantly improve the sensitivity of the electrode.

[0160] Comparative Example 3

[0161] First, prepare a monitoring probe with front and rear dual working electrodes, namely the first working electrode and the second working electrode respectively.

[0162] Secondly, prepare the β-hydroxybutyric acid sensitive layer solution and the glucose sensitive layer solution with the same formulation as in Example 1. Deposit the glucose sensitive layer solution on the first working electrode and deposit the β-hydroxybutyric acid sensitive layer solution on the second working electrode, both forming a linear pattern, where the area of the glucose sensing layer is about 0.3 mm 2 , and the area of the ketone body sensing layer is about 0.27 mm 2 ; Cure to obtain an electrode with a glucose and ketone body dual-component responsive sensing layer.

[0163] Then, prepare ethanol as a solvent, and prepare an alcohol solution with corresponding concentrations of polyvinylpyridine and PEGDGE according to Table 6 to obtain a diffusion-limiting membrane solution. Coat the diffusion-limiting membrane solution onto the electrode deposited with the sensitive layer reagent to make a dual-component monitoring working electrode with a membrane layer of about 35 microns.

[0164] Table 6 Composition of the diffusion-limiting membrane solution

[0165] Substance Concentration (mg / mL) Polyvinylpyridine 100 Polydimethylsiloxane 4 PEGDGE 35

[0166] Finally, prepare an 80% ethanol aqueous solution, and prepare a solution with corresponding concentrations of (poly(4-vinylpyridine-g-polyethylene glycol))-copoly-polystyrene and PEGDGE according to Table 7 to obtain a biocompatible membrane solution; coat the biocompatible membrane solution onto the diffusion-limiting membrane of the working electrode to make a working electrode with a biocompatible membrane layer of about 25 microns and a total membrane thickness of about 60 microns.

[0167] Table 7 Composition of the biocompatible membrane solution

[0168]

[0169] Figure 14 is a picture showing the monitoring probe of Comparative Example 3 of the blood glucose + blood ketone sensor involved in the present disclosure. Figure 15 is a picture showing another perspective of the monitoring probe of Comparative Example 3 of the blood glucose + blood ketone sensor involved in the present disclosure.

[0170] As Figure 14 ,Figure 15 As shown, it can be seen that in the absence of a compatibility film layer, the morphology of the semi-permeable membrane of the fabricated sensor shrinks, and the thickness is uneven up and down, which is not conducive to the preparation of sensors with high consistency.

[0171] [Blood ketone sensor]

[0172] Example 1, Comparative Examples 1-2

[0173] First, prepare a monitoring probe with a working electrode.

[0174] Secondly, prepare the sensitive layer reagent raw materials for each blood ketone sensor example and comparative example according to Table 8 (ruthenium coordination metal polymer, β-hydroxybutyrate dehydrogenase, myocardial xanthine oxidase, 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 oscillate to completely dissolve it to obtain the β-hydroxybutyrate sensitive layer solution for each example.

[0175] Deposit the β-hydroxybutyrate sensitive layer solution on the working electrode to form a linear pattern with a width of about 160 microns and a length of about 2.1 mm; cure it to obtain an enzyme sensing layer.

[0176] Table 8. Examples of blood ketone sensors

[0177]

[0178] Figure 16 It is a picture showing the monitoring probes of the sensors of Example 1 and Comparative Examples 1-2 of the blood ketone sensor involved in the present disclosure.

[0179] Then, prepare pure ethanol as a solvent, and configure an alcohol solution with corresponding concentrations of polyvinylpyridine, polydimethylsiloxane, and PEGDGE according to Table 9 to obtain a diffusion-limiting membrane solution. Coat the diffusion-limiting membrane solution on the working electrode on which the β-hydroxybutyrate sensitive layer solution has been deposited to make a working electrode with a diffusion-limiting membrane layer of about 20 microns.

[0180] Table 9. Composition of the diffusion-limiting membrane solution

[0181] Substance Concentration (mg / mL) Polyvinylpyridine 92 Polydimethylsiloxane 3 PEGDGE 30

[0182] Prepare an 80% ethanol aqueous solution, and configure a solution with corresponding concentrations of poly(4-vinylpyridine)-co-poly(styrene) and PEGDGE according to Table 10 to obtain a compatibility membrane solution; coat the compatibility membrane solution on the working electrode coated with the diffusion-limiting layer membrane to make a working electrode with a compatibility membrane layer of about 15 microns.

[0183] Table 10. Composition of the compatibility membrane solution

[0184]

[0185] Prepare an 80% aqueous ethanol solution, and prepare corresponding concentration solutions of (poly(4-vinylpyridine)-g-poly(ethylene glycol))-copoly-polystyrene and PEGDGE according to Table 11 to obtain a biocompatible membrane solution; coat the working electrode coated with the compatibility membrane layer with the biocompatible membrane solution to make a working electrode with a biocompatible membrane layer of about 15 microns and a total membrane thickness of 50 microns.

[0186] Table 11. Composition of the biocompatible membrane solution

[0187]

[0188] Figure 17 It is a picture showing the monitoring probes coated with the polymer membrane layer of Example 1 and Comparative Examples 1-2 of the blood ketone sensor involved in the present disclosure.

[0189] Finally, perform linear and stability tests on the ketone biosensors obtained in Example 1 and Comparative Examples 1-2:

[0190] (1) Linear test:

[0191] Immerse the ketone biosensor in a standard PBS buffer solution (pH 7.4) at 37°C; apply a working voltage until the ketone biosensor reaches a constant background current, and add sodium β-hydroxybutyrate with concentrations of 0 mM, 1 mM, 2 mM, 3 mM, 4 mM, 6 mM, and 8 mM to the solution to measure the linearity of the reaction. After each addition of sodium β-hydroxybutyrate, allow the solution to equilibrate for 10 minutes, and continuously stir the solution during the measurement to make its concentration uniform.

[0192] Figure 18 It is a response current - test concentration graph showing the linear test of Example 1 of the blood ketone sensor involved in the present disclosure. Figure 19 It is a response current - test concentration curve graph showing the linear test of Example 1 of the blood ketone sensor involved in the present disclosure.

[0193] As Figure 18 、 19 shown, the ketone biosensor of the blood ketone sensor in Example 1 shows good linearity in the 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.

[0194] (2) Stability test:

[0195] At 37°C, immerse the ketone biosensor in 3 mM sodium β-hydroxybutyrate (using PBS buffer solution with pH 7.4 as the solvent) and continuously measure for 14 days.

[0196] Figure 20 It is a response current - test concentration curve graph showing the stability test of Example 1 of the blood ketone sensor involved in the present disclosure. Figure 21 It is a response current - test concentration curve graph showing the stability test of Comparative Example 1 of the blood ketone sensor involved in the present disclosure. Figure 22 It is a response current - test concentration curve graph showing the stability test of Comparative Example 2 of the blood ketone sensor involved in the present disclosure.

[0197] As Figures 20 - 22 shown, Comparative Example 1 could not maintain 14 - day stability at 37°C, and its performance began to decline on the 5th day, while both Example 1 and Comparative Example 2 could maintain 14 - day stability because a stabilizer was added to their enzyme sensing layers, achieving longer - term stability of the ketone sensor. The response value of Example 1 is higher than that of Comparative Example 2, indicating that using HEPES buffer as the solvent for the sensitive layer reagent is superior to using PBS buffer as the solvent for the sensitive layer reagent.

[0198] Comparative Example 3

[0199] A monitoring probe with a working electrode was prepared in the same manner as in Example 1 of blood ketone, and an enzyme sensing layer was set on the working electrode to obtain a sensor without a semi - permeable membrane.

[0200] The sensor of Comparative Example 3 was exposed to a solution containing 1 mM sodium β - hydroxybutyrate standard PBS buffer (pH 7.4) at 37°C for 1 hour.

[0201] Figure 23 It is a response current - test concentration graph showing the linear test of Comparative Example 3 of the blood ketone sensor involved in the present disclosure.

[0202] As Figure 23 shown, within the 1 - hour test time, the current signal generated by the sensor gradually decreased from 240 nA to below 155 nA, showing poor stability. The test results indicate that the sensor without a semi - permeable membrane has poor stability when testing the concentration of sodium β - hydroxybutyrate solution.

[0203] Comparative Example 4

[0204] A monitoring probe with a working electrode was prepared in the same manner as in Comparative Example 4 of blood ketone, and an enzyme sensing layer was set on the working electrode to obtain a sensor without a semi - permeable membrane.

[0205] Prepare an 80% ethanol aqueous solution, and configure solutions with corresponding concentrations of poly(4-vinylpyridine)-co-poly(styrene) and PEGDGE according to the following table to obtain the inner layer membrane solution; coat the working electrode with the enzyme sensing layer with the inner layer membrane solution to make a working electrode containing a poly(4-vinylpyridine)-b-poly(styrene) membrane layer of about 35 microns.

[0206] Table 12 Composition of the Inner Layer Membrane Solution

[0207]

[0208]

[0209] Figure 24 It is a picture showing the working electrode of Comparative Example 4 of the blood ketone sensor involved in the present disclosure.

[0210] Prepare an 80% ethanol aqueous solution, and configure solutions with corresponding concentrations of (poly(4-vinylpyridine)-g-poly(ethylene glycol))-co-poly(styrene) and PEGDGE according to Table 13 to obtain the outer layer membrane solution; coat the working electrode coated with the inner layer membrane with the outer layer membrane solution to make a working electrode with an outer layer membrane of about 15 microns and a total membrane thickness of 50 microns.

[0211] Table 13 Composition of the Outer Layer Membrane Solution

[0212]

[0213] Figure 25 It is a picture showing the monitoring probe of the sensor of Comparative Example 4 of the blood ketone sensor involved in the present disclosure.

[0214] Finally, perform a stability test on the obtained ketone sensor; that is, at 37 °C, immerse the ketone biosensor of Comparative Example 4 in 3 mM sodium β-hydroxybutyrate (PBS buffer solution pH 7.4 as the solvent); continuously measure for 3 days.

[0215] Figure 26 It is a graph showing the response current - test time curve of the sensor of Comparative Example 4 of the blood ketone sensor involved in the present disclosure.

[0216] As Figure 26 shown, the initial response current of the sensor of Comparative Example 4 is 15.6 nA, and the response current of the sensor decreases to 11.2 nA after three days, with a total decrease of 28.2%, and an average decrease of 9.4% per day, which indicates that the sensor coated only with poly(4-vinylpyridine)-b-poly(styrene) and (poly(4-vinylpyridine)-g-poly(ethylene glycol))-co-poly(styrene) membranes has poor stability.

[0217] The main reason is that, compared with Example 1 of blood ketone, the film layer combination has a relatively high transmittance to sodium β-hydroxybutyrate, indicating that the diffusion limitation of this film is not strong enough to reduce the transmittance of sodium β-hydroxybutyrate and restrict the escape of small molecules from the enzyme layer, resulting in poor stability.

[0218] Comparative Example 5

[0219] First, prepare a monitoring probe with a working electrode.

[0220] Prepare the corresponding reagents according to the same enzyme layer formulation as in Example 1 of blood ketone. Use an ultra-micro pipetting robot to drop the β-hydroxybutyric acid-sensitive layer solution onto the working electrode to form 6 independent dot patterns with a diameter of 150 microns; perform curing to obtain an enzyme sensor with 6 dot patterns.

[0221] Figure 27 It is a picture showing the working electrode of the sensor of Comparative Example 5 of the blood ketone sensor involved in the present disclosure.

[0222] Then, coat the above 6-dot pattern enzyme sensor with a film layer according to the same film layer coating scheme as in Example 1 of blood ketone to obtain the ketone body sensor of Comparative Example 5.

[0223] Finally, perform a linearity test on the obtained ketone body sensor of Comparative Example 5: immerse the ketone body biosensor in a standard PBS buffer solution (pH 7.4) at 37°C; apply a working voltage until the ketone body biosensor reaches a constant background current, and add sodium β-hydroxybutyrate with concentrations of 1 mM, 2 mM, 3 mM, 4 mM, 5 mM, 6 mM, 7 mM, and 8 mM to the solution to measure the linearity of the reaction. After each addition of sodium β-hydroxybutyrate, continuously stir the solution during the measurement to make its concentration uniform.

[0224] Figure 28 It is a graph showing the response current - test concentration of the linearity test of the blood ketone sensor of Comparative Example 5 involved in the present disclosure.

[0225] As Figure 28 shown, it can be seen that it shows good linearity in the concentration range of 0 - 8 mM for sodium β-hydroxybutyrate, but its response sensitivity is relatively low, that is, the slope of the linear fitting curve is 0.677 nA / mM, while the sensitivity of the sensor in Example 1 is 2.08 nA / mM. Comparative Example 5 is only 32.5% of Example 1. This is because the area of the enzyme layer in Comparative Example 5 is smaller, resulting in a lower reaction area with the sodium β-hydroxybutyrate passing through the film layer. This indicates that increasing the area of the enzyme layer can effectively increase the sensor sensitivity.

[0226] The foregoing describes various examples of the present disclosure in the specific implementation manners. Although these descriptions directly describe the above examples, it should be understood that those skilled in the art can conceive of modifications and / or variations to the specific examples shown and described herein. Any such modifications or variations that fall within the scope of this specification are also intended to be included therein. Unless otherwise specified, the intention of the inventor is that the terms in the specification and claims be given the ordinary and customary meaning of those of ordinary skill in the art.

Claims

1. A multi - analyte sensor, characterized in that, it includes a working electrode, an enzyme sensing layer disposed on the working electrode, and a polymer film layer covering the enzyme sensing layer, the working electrode includes a first working electrode for detecting a first analyte and a second working electrode for detecting a second analyte, the first analyte is glucose, and the second analyte is β - hydroxybutyric acid; the enzyme sensing layer includes 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, and the area of the first analyte enzyme layer is greater than or equal to the area of the second analyte enzyme layer; the polymer film layer is configured to be permeable to both the first analyte and the second analyte, and has different permeability coefficients for the first analyte and the second analyte.

2. The multi - analyte sensor according to claim 1, characterized in that, the first analyte enzyme layer includes glucose enzyme, an electron mediator, and a cross - linker, the mass fraction of the glucose enzyme is 40% to 70%, the mass fraction of the electron mediator is 10% to 40%, and the mass fraction of the cross - linker is 0% to 20%.

3. The multi - analyte sensor according to claim 1, characterized in that, the second analyte enzyme layer includes hydroxybutyrate dehydrogenase, riboflavin - 5'-phosphate, coenzyme, and an electron mediator, in the second analyte enzyme layer, the mass fraction of hydroxybutyrate dehydrogenase is 10% to 20%, the mass fraction of riboflavin - 5'-phosphate 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%.

4. The multi - analyte sensor according to claim 1, characterized in that, it further includes a substrate, and the first working electrode and the second working electrode are disposed on the substrate and are respectively located on both sides of the substrate.

5. The multi - analyte sensor according to claim 1, characterized in that, the polymer film layer includes a vinylpyridine polymer and a cross - linker, and the mass fraction of the vinylpyridine polymer in the polymer film layer is 80% to 100%.

6. The multi - analyte sensor according to claim 1 or 5, characterized in that, the polymer film layer includes a regulator, and the regulator is polydimethylsiloxane.

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

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

1.

9. The multi - analyte sensor according to claim 1, characterized in that, 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.

10. The multi - analyte sensor according to claim 9, characterized in that, 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, which is a copolymer formed from 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.

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