Preparation method and application of glucose sensor based on glucose dehydrogenase

Abstract

The present invention relates to the field of sensor technology, and discloses a preparation method and application of a glucose sensor constructed based on glucose dehydrogenase. This preparation method comprises: S1, taking MCNT-NH2 and placing it in ethanol for ultrasonic dispersion to obtain a dispersion of MCNT-NH2; then coating the MCNT-NH2 dispersion on a substrate to obtain an MCNT-NH2-modified substrate electrode; S2, adding PEG9, TH, PEI and FAD-GDH solution to water, mixing evenly, and heating in a water bath to obtain a hydrogel sensing film; S3, coating the hydrogel sensing film prepared by S2 on the surface of the substrate electrode prepared in S1, and allowing it to stand at room temperature. The present invention realizes the construction of a three-dimensional network of hydrogels by cross-linking redox mediators TH, glucose dehydrogenase and polymers, thereby increasing the loading amount of protease catalysts and mediators on the substrate surface; at the same time, the use of a mediator with a low formula potential improves the sensitivity, anti-interference and accuracy of the glucose sensor for glucose detection.

Description

Technical Field

[0001] The present invention relates to the field of sensor technology, in particular to a preparation method and application of a glucose sensor constructed based on glucose dehydrogenase. Background Art

[0002] Glucose is a common, naturally occurring monosaccharide and an important intermediate metabolite in the human body, participating in key metabolic pathways. Furthermore, glucose is the body's most basic energy source. Glucose dispersed in the blood can be used to provide energy or converted into liver glycogen, muscle glycogen, and other substances for storage. In the fasting state, normal blood glucose concentrations generally remain between 4 and 6 mM. Exceeding or falling below this range can lead to diabetes or hypoglycemia. Therefore, controlling glucose intake and monitoring blood glucose levels in the human body are crucial. Accurate, rapid, simple, and real-time glucose measurement is crucial for both controlling glucose intake and monitoring blood glucose levels in patients with abnormal glucose metabolism. Furthermore, monitoring blood glucose levels can provide a better understanding of a person's health, thereby improving patient health through early diagnosis and treatment. Commonly used disposable test strip blood glucose monitors on the market typically use glucose oxidase or glucose dehydrogenase as enzyme catalysts for blood glucose monitoring. These devices offer advantages such as low blood sample volume, high sensitivity, and high accuracy. However, for patients who require dynamic monitoring and recording of blood glucose levels, disposable blood glucose testing devices have limitations. Therefore, there is an urgent need to develop a simple, economical and effective method that can continuously, directly, quickly and accurately analyze the concentration of glucose in complex samples (such as blood or beverages).

[0003] Electrochemical sensors play an important role in various fields due to their advantages, such as rapid response, good detection selectivity, high sensitivity, convenient signal acquisition, and ease of design based on detection objectives. Their working principle is to use a working electrode as a conversion element. The entire detection process is to identify the analyte through the effective and stable electrochemical signal transmitted after the stabilized biological or chemical material modified on the electrode reacts with the analyte, establish a linear relationship between the analyte parameters and the electrochemical signal, and thus use the electrochemical signal to measure the analyte parameters in practical applications. In particular, bioelectrochemical sensors based on biological enzymes have unique advantages in analyzing complex samples because they combine the advantages of enzyme catalysis, such as high efficiency, selectivity, and specificity. However, because the active site of the enzyme is generally buried deep within its protein shell, the electron transfer kinetics between it and the electrode substrate surface are relatively slow or difficult to achieve. In addition, increasing the enzyme loading on the electrode substrate surface is also crucial to improving the sensitivity of the bioelectrochemical sensor. At the same time, when measuring the glucose concentration in complex samples (such as blood or beverages), the bioelectrochemical sensor should be anti-interference to prevent interference from substances such as ascorbic acid and dopamine that are easily oxidized at low potentials, resulting in inaccurate measurement results.

[0004] Therefore, there is an urgent need to develop a simple and economical method for making a glucose sensor that can achieve sensitive and rapid determination of glucose, has high selectivity for glucose, has good anti-interference ability, and can be used for continuous monitoring of glucose in complex samples. Summary of the Invention

[0005] The technical problem addressed by the present invention is to overcome the shortcomings of the existing technology and provide a method for preparing and applying a reagent-free glucose sensor based on glucose dehydrogenase, enabling continuous and highly selective detection of glucose concentration. The present invention utilizes flavin adenine dinucleotide-dependent glucose dehydrogenase (FAD-GDH) to perform an enzymatic reaction and employs a low-onset-potential redox mediator, thionine (TH), to mediate electron transfer between the electrode substrate and the active center of the glucose dehydrogenase. Simultaneously, a method is developed to immobilize the glucose dehydrogenase and the redox mediator TH on the electrode surface, increasing their loading on the electrode substrate. This allows for efficient catalysis of glucose oxidation and generation of a current signal, rapidly responding to changes in glucose concentration, and enabling continuous, rapid, sensitive, and highly selective detection of glucose concentration over a period of time. The developed electrochemical sensor operates at low overpotentials, preventing interference from substances such as ascorbic acid and dopamine in complex samples. Furthermore, since both the mediator and glucose dehydrogenase are immobilized on the substrate surface, no exogenous reagents need to be added during the detection process, and the mediator and glucose dehydrogenase will not leak into the sample being tested. Therefore, the electrochemical sensor developed in this project, after further miniaturization, also has the potential to serve as an implantable sensor for continuous blood glucose monitoring in living organisms.

[0006] The technical solution adopted by the present invention to solve its technical problem is:

[0007] A method for preparing a reagent-free glucose sensor based on glucose dehydrogenase comprises the following steps:

[0008] Step S1: preparing a base electrode:

[0009] MCNT-NH2 (amino carbon nanotubes) were ultrasonically dispersed in ethanol (EtOH) to obtain a MCNT-NH2 dispersion. The MCNT-NH2 dispersion was then coated on a substrate to prepare an MCNT-NH2-modified substrate electrode.

[0010] Step S2: preparing a hydrogel sensing membrane:

[0011] PEG9 (bis(2,5-dioxopyrrolidin-1-yl)-4,7,10,13,16,19,22,25,28-nonaoxatriacontane dioate), TH (thionine), PEI (polyethyleneimine), and FAD-GDH (flavin adenine dinucleotide-dependent glucose dehydrogenase) were added to water and mixed uniformly to prepare a mixed solution. The mixed solution was then heated in a water bath to prepare a hydrogel sensing membrane FAD-GDH / PEI / TH.

[0012] Step S3: preparing a glucose sensor:

[0013] The hydrogel sensing film prepared in step S2 is coated on the upper surface of the base electrode prepared in step S1, and the finished glucose sensor is obtained by standing at room temperature.

[0014] Furthermore, in step S1, the concentration of MCNT-NH2 in the MCNT-NH2 dispersion is 2 mg / mL.

[0015] Furthermore, in step S1, each cm 2 The substrate was coated with 200–300 μL of MCNT-NH2 dispersion.

[0016] Furthermore, in step S1, the substrate is selected from one of a glassy carbon (GC) electrode, a gold (Au) electrode and a carbon cloth (CC) substrate, and finally the sensors FAD-GDH / PEI / TH / MCNT-NH2 / GCE, FAD-GDH / PEI / TH / MCNT-NH2 / Au and FAD-GDH / PEI / TH / MCNT-NH2 / CC are prepared.

[0017] Furthermore, in step S2, the water bath temperature is 35° C., and the water bath time is 30 to 60 minutes.

[0018] Furthermore, in the mixed solution of step S2, the concentration of FAD-GDH is 1.75-6.5 mg / mL, the concentration of PEG9 is 2.2 mg / mL, the concentration of TH is 2 mg / mL, and the concentration of PEI is 0.3-1 mg / mL. The FAD-GDH solution is prepared by adding FAD-GDH to a phosphate buffer solution (PBS) with a pH of 7.0 and mixing thoroughly.

[0019] Furthermore, in step S3, each cm 2 The base electrode is coated with 40-200 μL of hydrogel sensing film.

[0020] The solvent-free glucose sensor prepared in this invention uses flavin adenine dinucleotide-dependent glucose dehydrogenase (FAD-GDH) as an enzyme catalyst, capable of highly selectively oxidizing glucose to gluconolactone. TH, which contains amino groups and has a low formula potential, is selected as a mediator to mediate electron transfer between the electrode and the enzyme's active center. PEG9 is used as a crosslinker, and PEI is used as a polymer backbone. A three-dimensional network hydrogel membrane is constructed by crosslinking the components. Aminated carbon nanotubes (MCNT-NH2) are used as conductive scaffolds and crosslinking sites. This allows for the immobilization of more enzymes and mediators on the substrate surface than with planar electrodes, increasing the modified electrode's response current per unit glucose concentration and, therefore, the sensitivity of the sensor membrane. All components are placed in a 35°C water bath to construct the redox hydrogel sensor membrane. The reaction process between the crosslinker and the amino groups is as follows:

[0021] .

[0022] The working principle of the glucose sensor prepared by the present invention is as follows Figure 1 As shown, the electron transfer process is as follows:

[0023]

[0024] Glucose is first oxidized by the oxidized glucose dehydrogenase GDH (FAD) to gluconolactone. Simultaneously, GDH (FAD) is reduced to GDH (FADH2), representing reaction 1. Next, the oxidized mediator (TH(ox)) present in the system oxidizes the reduced GDH (FADH2) back to GDH (FAD), returning it to its oxidized state and allowing it to continue reacting with glucose. Simultaneously, the TH molecule itself is reduced to TH(red), representing reaction 2. Third, the reduced TH(red) is oxidized by the applied potential at the electrode, returning it to its oxidized state, TH(ox), representing reaction 3. The oxidation current of TH(red) during reaction 3 is measured through the electrode substrate to indirectly determine the corresponding glucose content.

[0025] The beneficial effects of the present invention are: the present invention has a reasonable design, an economical and simple preparation method, and has the following advantages:

[0026] (1) The present invention introduces PEI as a polymer skeleton and combines it with the crosslinker PEG9 through covalent interaction, which can block the π-π stacking of TH with the surface of carbon cloth and carbon nanotubes (MCNT-NH2) to a certain extent, making it have a certain fluidity in the hydrogel, promoting the contact between the mediator and the enzyme active center, and increasing the rate of mediated electron transfer. At the same time, the rich amino groups of PEI are used to react with the crosslinker PEG9 to fix FAD-GDH and TH to form a three-dimensional hydrogel. This strategy avoids the complex synthesis steps of redox polymers. In addition, the introduction of carbon nanotubes (MCNT-NH2) can improve the hydrophilicity of the carbon cloth substrate and, as a crosslinking site and conductive support, enable the hydrogel sensor membrane to be fixed on the surface of the carbon cloth electrode with a high loading amount, thereby improving the stability and sensitivity of the electrode; at an applied potential of 0.2 V vs. Ag / AgCl / 3 M KCl, when the glucose concentration reaches 60 mM, the catalytic current can reach a maximum of 1.4 mA cm -2 ,like Figure 7 As shown in (a);

[0027] (2) The redox mediator TH selected in the present invention has a low onset potential and high electron transfer kinetics. When measuring glucose in complex samples, it can avoid the influence of factors such as ascorbic acid (AA), dopamine (DOP), and uric acid (UA), thereby greatly improving the anti-interference and accuracy of glucose detection.

[0028] (3) The glucose sensor of the present invention introduces hydrogel technology and constructs a hydrogel sensing membrane by cross-linking polymers and components, thereby achieving the immobilization of protease catalysts and mediators in the hydrogel on the substrate surface. No exogenous reagents are required during the detection process, which can achieve reagent-free detection and avoid the leakage of mediators and glucose dehydrogenase into the sample to be tested, which would contaminate the detection system. BRIEF DESCRIPTION OF THE DRAWINGS

[0029] In order to more clearly illustrate the specific embodiments of the present invention or the technical solutions in the prior art, the following briefly introduces the drawings required for use in the specific embodiments or the description of the prior art. Obviously, the drawings described below are some embodiments of the present invention. For ordinary technicians in this field, other drawings can be obtained based on these drawings without paying any creative work.

[0030] Figure 1 This is a diagram of the electron transfer path of glucose oxidation catalyzed by the glucose sensor prepared by the present invention;

[0031] Figure 2 Schematic diagram of a glucose sensor prepared in Example 1 using a hydrogel membrane based on glucose dehydrogenase;

[0032] Figure 3 Figures 1 and 13 are the electrochemical performance diagrams of Example 1 and Comparative Example 13; (a) and (c) are cyclic voltammograms of the glassy carbon electrode in Ar-saturated 0.2 M PBS (pH 7.4) containing 0.25 mM TH / TB, with a scan rate of 10 mV / s; (b) and (d) are cyclic voltammograms of FAD-GDH / PEI / TH / MCNT-NH2 / CC (b) and FAD-GDH / PEI / TB / MCNT-NH2 / CC (d) in Ar-saturated 0.2 M PBS (pH 7.4) with (red line) and without (black line) 30 mM glucose, with a scan rate of 5 mV / s;

[0033] Figure 4 1 is a summary graph of the catalytic oxidation current of glucose by FAD-GDH / PEI / TH / MCNT-NH2 / CC prepared according to the PEI amount and enzyme content in Examples 1, 9 to 12, wherein (a) is a graph showing different PEI mass-to-volume ratios (Examples 1, 9 and 10); (b) is a graph showing the catalytic current response of different FAD-GDH mass-to-volume ratios (Examples 1, 11 and 12) in Ar-saturated 0.2 M PBS (pH 7.4) containing 30 mM glucose;

[0034] Figure 5 The FAD-GDH / PEI / TH hydrogel membranes prepared in Examples 1 to 6 with different modification solution volumes were modified on MCNT-NH2 / CC (0.0625 cm 2 ) Glucose oxidation current response diagram of FAD-GDH / PEI / TH / MCNT-NH2 / CC prepared on the surface of the membrane in 0.1 M PBS (pH 7.4) saturated with Ar containing 30 mM glucose (applied voltage: 0.2 V vs. Ag / AgCl / 3 M KCl);

[0035] Figure 611 and 12; (a) is a cyclic voltammogram of FAD-GDH / PEI / TH / MCNT-NH2 / CC in Ar-saturated 0.2 M PBS (pH 7.4) at different scan rates, and the scan rate from inside to outside is: 2~200 mV / s; (b) is the dependence of the peak current of the anode (red square) and cathode (blue square) obtained from (a) on the scan rate, both of which are linearly related to the scan rate; Example 1 (c), Comparative Example 11 (d) and Comparative Example 12 (e) in Ar-saturated 0.2 M PBS (pH 7.4) without (black line) glucose, with 5 mM (red line) glucose, with 10 mM (blue line) glucose, with 30 mM (green line) glucose, and with 30 mM glucose and free thionine (purple), the scan rate is 5 mV / s;

[0036] Figure 7 Figure 1 shows the response of the catalytic current density of the glucose dehydrogenase-based glucose sensor prepared using FAD-GDH / PEI / TH / MCNT-NH2 / CC in Example 1 (a) and FAD-GDH / TH / MCNT-NH2 / CC in Comparative Example 12 (b) as a function of glucose concentration (0-150 mM). The Michaelis-Menten equation was fitted to obtain the Michaelis-Menten kinetic parameters. The inset shows the linear response curves of the catalytic current to glucose concentration for the corresponding groups. All electrodes were evaluated in Ar-saturated 0.2 M PBS (pH 7.4). Chronoamperometry was performed at a potential of 0.2 V (vs. Ag / AgCl / 3 M KCl). Error bars represent standard deviations (n=3).

[0037] Figure 8 Cyclic voltammograms of (a) FAD-GDH / PEI / TH / MCNT-NH2 / GCE of Example 7 and (b) FAD-GDH / PEI / TH / MCNT-NH2 / Au of Example 8 in Ar-saturated 0.2 M PBS (pH 7.4) without (black line), with 10 mM (blue line) glucose, and with 30 mM (red line) glucose;

[0038] Figure 9 Figures 1 and 2 show the results of the anti-interference experimental test of the continuous sensing performance of the hydrogel membrane in Example 1; (a) shows the current response after adding interfering substances (Ep: epinephrine; DA: dopamine; UA: uric acid; APAP: acetaminophen; AA: ascorbic acid, as indicated by the arrows) when continuously monitoring glucose using the chronoamperometry method; and (b) shows the standard addition method analysis diagram obtained by fitting the actual sample using the chronoamperometry method (applied voltage: -0.1 V vs. Ag / AgCl / 3 M KCl). DETAILED DESCRIPTION

[0039] It should be noted that the following detailed descriptions are illustrative and intended to provide further explanation of the present application. Unless otherwise specified, all technical and scientific terms used herein have the same meaning as commonly understood by those skilled in the art to which the present application belongs.

[0040] It should be noted that the terms used herein are only for describing specific embodiments and are not intended to limit the exemplary embodiments according to the present application. As used herein, unless the context clearly indicates otherwise, the singular also includes the plural. In addition, it should be understood that when the terms "comprise" and / or "include" are used in this specification, they indicate the presence of features, steps, operations, devices, components and / or combinations thereof.

[0041] The following will clearly and completely describe the technical solutions of the present invention in conjunction with the embodiments. Obviously, the embodiments described are only some embodiments of the present invention, not all embodiments. Based on the embodiments of the present invention, all other embodiments obtained by ordinary technicians in this field without making creative efforts are within the scope of protection of the present invention.

[0042] In the hydrogel membrane modification solution FAD-GDH / PEI / TH / PEG9 system, PEG9 contains an NHS ester group that can covalently bind to FAD-GDH, TH, PEI, and MCNT-NH2, which contain primary amine groups. By cross-linking these components, a three-dimensional network hydrogel membrane is formed. The innovative introduction of PEI and MCNT-NH2 in this invention further increases the thickness of the hydrogel membrane on the electrode substrate, significantly increasing the loading of FAD-GDH and TH on the substrate surface. Furthermore, due to the use of long-chain crosslinkers, the cross-linked TH can oscillate within a certain range within the hydrogel, mediating rapid electron transfer between the electrode surface and the enzyme's active center, FAD, thereby constructing a sensitive glucose sensor.

[0043] Example 1

[0044] A method for preparing a reagent-free glucose sensor based on glucose dehydrogenase is as follows:

[0045] (1) Take 4 mg of amino carbon nanotubes in 2 mL of ethanol and disperse them evenly under ultrasonication for 3 h to obtain MCNT-NH2 dispersion. Take 12~20 μL of the dispersion and apply it to a carbon cloth substrate (0.0625 cm 2 ), a carbon cloth electrode was prepared; after the carbon cloth substrate was modified with amino carbon nanotubes, the hydrophobicity of the carbon cloth surface could be reduced, and the amino-containing carbon nanotubes could also cross-link with the cross-linking agent in the modified hydrogel to achieve a stronger fixation effect;

[0046] (2) Preparation of hydrogel sensing membrane: 0.4 μL PEG9 (100 mg / mL in H2O), 7.2 μL TH (5 mg / mL in H2O), 0.8 μL PEI (10 mg / mL in H2O) and 3 μL FAD-GDH (20 mg / mL in pH 7.0 PBS (phosphate buffer solution)) were mixed evenly in 6.6 μL water to prepare a mixed solution. The final concentrations of the mixed solution were: FAD-GDH was 3.3 mg / mL, PEG9 was 2.2 mg / mL, TH was 2 mg / mL, and PEI was 1 mg / mL. The mixed solution was then placed in a 35 °C aqueous solution and heated in a water bath for 60 min. The water bath heating treatment was beneficial to the reaction of the NHS ester group of PEG9 with the primary amine, which promoted the cross-linking of the components and formed a three-dimensional network hydrogel. While enhancing the electron transfer efficiency, it also increased the loading amount of FAD-GDH and TH on the substrate surface, thereby improving the efficiency of the catalytic glucose oxidation reaction.

[0047] (3) Take 5 μL of the prepared hydrogel sensor film and apply it to the carbon cloth electrode (0.0625 cm 2 ) and allowed to stand at room temperature for 24 h to solidify. The entire system was FAD-GDH / PEI / TH / PEG9, and the finished glucose sensor was obtained.

[0048] Figure 2 This is a schematic diagram of the reagent-free glucose sensor based on glucose dehydrogenase prepared in Example 1. As can be seen from the figure, the glucose sensor comprises a three-dimensional hydrogel structure formed by coupling PEI and TH with the crosslinker PEG9. This 3D structure enables the immobilization of larger amounts of the enzyme and mediator TH. The presence of MCNT-NH2 in the redox hydrogel acts as a support for the hydrogel membrane, improving membrane stability.

[0049] Example 2 to Example 6

[0050] The difference from Example 1 is that in step (3), the amount of hydrogel sensing film applied is different from that of Example 1, which is 5 μL per electrode (0.0625 cm 2 ), the coating amount of the hydrogel sensing membrane of each electrode in Examples 2 to 6 is 2.5 μL, 3.5 μL, 6 μL, 10 μL and 12.5 μL, respectively.

[0051] Example 7

[0052] The difference from Example 1 is that a glassy carbon electrode is used instead of a carbon cloth electrode.

[0053] Example 8

[0054] The difference from Example 1 is that gold electrodes are used instead of carbon cloth electrodes.

[0055] Example 9 and Example 10

[0056] The difference from Example 1 is that in step (3), the final concentration of PEI in the mixed solution is different. Unlike the PEI concentration of 1 mg / mL in Example 1, the PEI concentrations in Examples 9 and 10 are 0.3 mg / mL and 0.6 mg / mL, respectively.

[0057] Example 11 and Example 12

[0058] The difference from Example 1 is that in step (3), the final concentration of FAD-GDH in the mixed solution is different. Unlike the FAD-GDH concentration of 3.3 mg / mL in Example 1, the FAD-GDH concentrations in Examples 11 and 12 are 1.75 mg / mL and 6.5 mg / mL, respectively.

[0059] Comparative Example 1 to Comparative Example 10

[0060] In the above comparative examples, different base electrodes, different mediators, different mediator fixing methods, different FAD-GDH fixing methods and different cross-linking agents were replaced to prepare different glucose sensors.

[0061] Comparative Example 11

[0062] The difference from Example 1 is that in this comparative example, in step (2), the mixed solution is not heated in a water bath at 35°C.

[0063] Comparative Example 12

[0064] The difference from Example 1 is that in this comparative example, PEI is not used as the polymer skeleton to construct the hydrogel structure.

[0065] Comparative Example 13

[0066] The difference from Example 7 is that in this comparative example, toluidine blue TB is used as a medium, and finally a glucose enzyme biosensor FAD-GDH / TB / MCNT-NH2 / CC in which electron transfer is mediated by TB is obtained.

[0067] Then, the glucose sensors prepared in Examples 1 to 12 and Comparative Examples 1 to 13 were tested by measuring the catalytic current density in a phosphate buffer solution containing 30 mM glucose at an applied potential of 0.2 V. The test results are shown in Table 1.

[0068] Table 1 Test results of Examples 1 to 12 and Comparative Examples 1 to 13

[0069]

[0070]

[0071] As shown in Table 1, in Comparative Examples 1 to 6, the catalytic currents of 1,2-naphthoquinone (1,2-NQ), 9,10-phenanthrenequinone (9,10-PQ), and thionine (TH) as mediators were compared in different quinone media. Although the catalytic current of the glucose sensor prepared with 1,2-NQ was higher, because its molecular structure does not contain amino groups, it cannot be well immobilized on the substrate surface through cross-linking. During the measurement process, 1,2-NQ will leak into the electrolyte, resulting in poor stability of the prepared glucose sensor and contamination of the test sample.

[0072] Figure 3The figures are the electrochemical performance diagrams of Example 1 and Comparative Example 13; wherein, (a) and (c) are cyclic voltammograms of the glassy carbon electrode in Ar-saturated 0.2 M PBS (pH 7.4) containing 0.25 mM TH / TB, with a scan rate of 10 mV / s; (b) and (d) are cyclic voltammograms of FAD-GDH / PEI / TH / MCNT-NH2 / CC (b) and FAD-GDH / PEI / TB / MCNT-NH2 / CC (d) in Ar-saturated 0.2 M PBS (pH 7.4) with the presence (red line) and absence (black line) of 30 mM glucose, with a scan rate of 5 mV / s. (a) and (c) of the figure are cyclic voltammograms of a glassy carbon electrode in 0.2 M PBS (pH 7.4) containing 0.25 mM TH / TB. A pair of electrochemically reversible redox peaks are obtained, corresponding to the mutual conversion between the redox states of TH and TB. From this, the redox potentials of toluidine blue (TB) and thionine (TH) can be obtained as -0.191 V vs. Ag / AgCl / 3 M KCl and -0.147 V vs. Ag / AgCl / 3 M KCl, respectively. The redox potential of thionine is significantly different from that of the FAD-GDH active center (-0.34 V vs. Ag / AgCl / 3 M KCl), thus theoretically providing a greater driving force for electron transfer when used as a mediator. Figures (b) and (d) show that the glucose sensors FAD-GDH / PEI / TH / MCNT-NH2 / CC and FAD-GDH / PEI / TB / MCNT-NH2 / CC constructed with TH and TB as mediators exhibit a pair of distinct redox peaks in glucose-free solutions, indicating that TH or TB is immobilized within the hydrogel membrane. Upon addition of 30 mM glucose, both glucose sensors exhibit catalytic oxidation currents centered around the mediator's redox peak. However, the catalytic current of the TH mediator group is significantly higher than that of the TB group. Therefore, it was finally decided to use TH as a mediator to mediate the transfer of electrons between the FAD-GDH active center and the electrode, and to construct FAD-GDH / PEI / TH / MCNT-NH2 / CC.

[0073] Figure 4Figures 1 and 12 show the effects of PEI content and enzyme content on the glucose catalytic current of the glucose sensors prepared using the FAD-GDH / PEI / TH / MCNT-NH2 / CC system. (a) shows the effect of different PEI mass-to-volume ratios (Examples 1, 9, and 10) on the glucose oxidation current of the glucose sensors prepared; (b) shows the catalytic current response of different FAD-GDH mass-to-volume ratios (Examples 1, 11, and 12) in 0.2 M PBS (pH 7.4) containing 30 mM glucose and saturated with Ar. The figure shows that the glucose sensor with a PEI content of 1 mg / mL and an enzyme (FAD-GDH) concentration of 3.27 mg / mL achieves the highest current density. To achieve optimal sensitivity, both must be optimized: controlling the PEI content to ensure enzyme immobilization, matching the enzyme content to prevent aggregation, and ensuring mass transfer efficiency for optimal catalytic performance. Therefore, Example 1 is the optimal example.

[0074] Figure 5 The FAD-GDH / PEI / TH hydrogel membranes with different modified volumes prepared in Examples 1 to 6 were modified on MCNT-NH2 / CC (0.0625 cm 2 ) in Ar-saturated 0.1 M PBS (pH 7.4) containing 30 mM glucose (applied voltage: 0.2 V vs. Ag / AgCl / 3 M KCl). This figure explores the effect of redox hydrogel film thickness on the catalytic current. With increasing film thickness, the current density first increases and then decreases. The current density reaches its maximum when the FAD-GDH / PEI / TH mixture modification amount is 5 μL / root. Moderately increasing the membrane thickness can load more enzymes and enhance catalytic ability. However, excessively thick membranes increase the diffusion resistance of the substrate glucose in the formed hydrosol and the electron transfer efficiency in the hydrosol, ultimately leading to a decrease in catalytic effect.

[0075] Figure 6The figures are the electrochemical performance diagrams of Example 1, Comparative Example 11 and Comparative Example 12; wherein, (a) is the cyclic voltammogram of FAD-GDH / PEI / TH / MCNT-NH2 / CC in Ar-saturated 0.2 M PBS (pH 7.4) at different scan rates, and the scan rate from inside to outside is: 2~200 mV / s; (b) is the dependence of the peak current of the anode (red square) and cathode (blue square) obtained from (a) on the scan rate; the cyclic voltammograms of Example 1 (c), Comparative Example 11 (d) and Comparative Example 12 (e) in Ar-saturated 0.2 M PBS (pH 7.4) without (black line), with 5 mM (red line) glucose, with 10 mM (blue line) glucose, with 30 mM (green line) glucose, and with 30 mM glucose and free thionine (purple), with a scan rate of 5 mV / s. As can be seen from the figure, for Example 1, FAD-GDH / PEI / TH / MCNT-NH2 / CC shows a pair of redox peaks centered at -0.147 V, corresponding to the mutual conversion of TH (red) and TH (ox), and the peak current increases with the increase of the scan rate ( Figure 6 , a), and the peak current is linearly related to the scan rate ( Figure 6 , b), indicating that the system is surface-controlled, demonstrating the successful construction of a reagent-free redox hydrogel.

[0076] In the PEI system ( Figure 6 , c), the constructed glucose sensor has a higher catalytic current density for the same concentration of glucose, and its catalytic current density changes more obviously with the glucose concentration. In addition, after further adding free redox mediator TH to the electrolyte, the catalytic current remains unchanged, which indicates that the electron transfer rate in the constructed hydrogel membrane is fast enough. For the system without PEI ( Figure 6 ,e), after adding the free redox mediator TH, the catalytic oxidation current is significantly improved, indicating that there are defects in the electron transfer path in the constructed hydrogel membrane. The glucose sensor constructed by this method has low sensitivity to glucose detection. The comparison of the two figures effectively proves that the three-dimensional hydrogel constructed by PEI as a polymer skeleton in this system can load more glucose dehydrogenase and redox mediator TH, and form an effective electron transfer path. In addition, the importance of the water bath step in the preparation of the hydrogel membrane is compared. When there is no water bath heating step ( Figure 6 d). The catalytic current of the constructed glucose sensor is lower than that after the water bath. This is because the water bath heating step facilitates the reaction between the NHS ester group in the crosslinker PEG9 and the primary amine. Furthermore, considering the optimal activity temperature of the enzyme, the water bath was heated at 35°C for 1 hour to promote the coupling reaction between the NHS ester group and the primary amine.

[0077] Figure 7 The graph shows the response of the catalytic current density of the glucose dehydrogenase-based glucose sensors prepared in Example 1 (a) and Comparative Example 12 (b) to glucose concentration (0–150 mM). The Michaelis-Menten kinetic parameters were obtained by fitting the Michaelis-Menten equation. The inset shows the linear response curve of the catalytic current of the glucose sensors to glucose concentration. All electrodes were evaluated in Ar-saturated 0.2 M PBS (pH 7.4). The data shown are from chronoamperometric experiments conducted at a potential of 0.2 V (vs. Ag / AgCl / 3 M KCl). Error bars represent standard deviations (n=3). The graph shows the glucose concentration dependence of the catalytic current density measured at an applied voltage of 0.2 V vs. Ag / AgCl / 3 M KCl for Example 1 compared to Comparative Example 12. At each glucose concentration, the current response of the glucose sensor prepared with the PEI-containing redox hydrogel is greater, demonstrating its higher sensitivity for glucose detection. Moreover, in Example 1 containing PEI, the linear detection range of glucose is wider (0.76~20 mM) than that of Comparative Example 12 without PEI ( Figure 7 , a).

[0078] Figure 8 Figures 7(a) and 8(b) show cyclic voltammograms of FAD-GDH / PEI / TH / MCNT-NH2 / GCE in Example 7 (a) and FAD-GDH / PEI / TH / MCNT-NH2 / Au in Ar-saturated 0.2 M PBS (pH 7.4) containing no glucose (black line), 10 mM glucose (blue line), and 30 mM glucose (red line). As shown in the figures, in an electrolyte containing 30 mM glucose, milliampere-level catalytic currents were still observed in glucose sensors prepared using the glassy carbon electrode (Example 7) and the gold electrode (Example 8) as electrode substrates. This demonstrates that the hydrogel-based glucose sensor developed in this invention is not only highly sensitive but also widely applicable to electrode substrates. Glucose sensors prepared using different electrode substrates all achieved high catalytic current values ​​for glucose, demonstrating the wide applicability of this method.

[0079] Figure 9The graphs show the results of the anti-interference experimental test of the continuous sensing performance of the hydrogel membrane in Example 1; (a) shows the current response after adding interfering substances (Ep: epinephrine; DA: dopamine; UA: uric acid; APAP: acetaminophen; AA: ascorbic acid, as indicated by the arrows) when continuously monitoring glucose using the chronoamperometry method; (b) shows the actual sample concentration measured using the standard addition method (applied voltage: -0.1 V vs. Ag / AgCl / 3 M KCl). As can be seen from the graph, under an applied voltage of -0.1 V (vs. Ag / AgCl / 3 M KCl), the chronoamperometry method is used to continuously monitor the glucose and interfering substances that may exist in the human body in Example 1 ( Figure 9 , a), first, when glucose was added to make the glucose concentration in the test sample reach 5 mM (at 100 s), an obvious stable catalytic current was observed. Then, 0.2 mM epinephrine, 1 mM dopamine, 0.2 mM uric acid, 1 mM acetaminophen and 1 mM ascorbic acid were gradually added (as shown in Figure 2). Figure 9 , as indicated by the arrow in a), no obvious disturbance was observed in the catalytic current of glucose. Subsequently, after all interfering substances were added, 5 mM glucose was added to make the glucose concentration in the test sample reach 10 mM, and a rapid increase in the catalytic current response was further observed. It was proved that the glucose sensor prepared by the redox hydrogel film not only has a high anti-interference ability under continuous detection working conditions, but also maintains a fast and highly sensitive response to glucose. The glucose concentration in the actual sample was detected using the glucose sensor of Example 1. 2 mL of 5% glucose injection was taken and diluted 80 times with 0.2 M PBS (pH 7.4) to obtain the actual sample to be tested. The theoretical value of its glucose concentration is 3.5 mM. By the standard addition method, 5% glucose was added to the sample solution (c x ) were added with glucose standard solution (0.5 M) to make the glucose concentration c x +1 mM, c x +3 mM, c x +6 mM, c x +10 mM, the catalytic current density when different concentrations of glucose were added was measured by chronoamperometry, and the results were Figure 9 In (b), a linear fitting is performed to obtain the linear equation y=28.24x+102.25, and the y value of the fitting curve is 0, which means the actual sample concentration c is obtained. x The sample recovery rate was 103.71%, which proved that the glucose sensor prepared by the redox hydrogel film had high accuracy.

[0080] In summary, the present invention has developed a simple method to construct a reagent-free glucose sensor by cross-linking a redox mediator, glucose dehydrogenase, and a polymer. The hydrogel construction immobilizes the protease catalyst and mediator, while the mediator can oscillate within a certain range, mediating rapid electron transfer. This improves the sensitivity of the glucose sensor while preventing contamination of the detection system by the mediator during the detection process. Furthermore, the three-dimensional network of the hydrogel increases the loading capacity of the protease catalyst and mediator on the substrate surface, improving the glucose sensor's sensitivity to glucose detection. Furthermore, the glucose sensor produced by the present invention utilizes a redox mediator with a low onset potential, enabling glucose measurements at lower applied potentials, resulting in improved interference resistance and accuracy.

[0081] Furthermore, the method of the present invention has broad applicability. Given the importance of immobilizing components in the design of wearable devices and in vivo detection devices, the glucose sensor of the present invention incorporates a hydrogel technology concept to prevent components from becoming free and dispersible in the blood. By constructing a redox-capable hydrogel using a polymer, the glucose sensor achieves immobilization of the protease catalyst and mediator, preventing contamination of the detection system by the mediator's solvation. Therefore, the glucose sensor has great potential for application in wearable devices and in vivo detection devices.

[0082] Finally, it should be noted that the above embodiments are only used to illustrate the technical solutions of the present invention, rather than to limit it. Although the present invention has been described in detail with reference to the above embodiments, those skilled in the art should understand that they can still modify the technical solutions described in the above embodiments, or replace some or all of the technical features therein with equivalents. However, these modifications or replacements do not cause the essence of the corresponding technical solutions to deviate from the scope of the technical solutions of the embodiments of the present invention.

Claims

1. A method for preparing a glucose sensor based on glucose dehydrogenase, characterized in that: The specific steps include: Step S1: preparing a base electrode: MCNT-NH2 is placed in ethanol and ultrasonically dispersed to obtain a MCNT-NH2 dispersion; the MCNT-NH2 dispersion is then coated on a substrate to prepare a MCNT-NH2 modified substrate electrode; Step S2: preparing a hydrogel sensing membrane: Bis(2,5-dioxopyrrolidin-1-yl)-4,7,10,13,16,19,22,25,28-nonaoxatriacontanoate PEG9, thionine TH, polyethyleneimine PEI, and flavin adenine dinucleotide-dependent glucose dehydrogenase FAD-GDH are added to water and mixed uniformly to prepare a mixed solution; the mixed solution is then heated in a water bath to prepare a hydrogel sensing membrane; Step S3: preparing a glucose sensor: The hydrogel sensing film prepared in step S2 is coated on the upper surface of the base electrode prepared in step S1, and the finished glucose sensor is obtained by standing at room temperature; In step S2, the water bath temperature is 35° C., and the water bath time is 30 to 60 min. In the mixed solution of step S2, the concentration of FAD-GDH is 1.75 to 6.5 mg / mL, the concentration of PEG9 is 2.2 mg / mL, the concentration of TH is 2 mg / mL, and the concentration of PEI is 0.3 to 1 mg / mL. In step S3, each cm 2 The base electrode is coated with 40-200 μL of hydrogel sensing film.

2. The method for preparing a glucose sensor based on glucose dehydrogenase according to claim 1, characterized in that: In step S1, the concentration of MCNT-NH2 in the MCNT-NH2 dispersion is 2 mg / mL.

3. The method for preparing a glucose sensor based on glucose dehydrogenase according to claim 1, wherein: In step S1, each cm 2 The substrate was coated with 200–300 μL of MCNT-NH2 dispersion.

4. The method for preparing a glucose sensor based on glucose dehydrogenase according to claim 1, wherein: In the step S1, the substrate is selected from one of a glassy carbon electrode, a gold electrode and a carbon cloth substrate.

5. An application of a glucose sensor based on glucose dehydrogenase prepared by the method according to any one of claims 1 to 4, characterized in that: Application of the glucose sensor in detecting glucose concentration.

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