Microfluidic sensing chip based on induced electroosmotic flow and preparation method thereof

Abstract

The invention discloses a micro-fluidic sensing chip based on induced electroosmotic flow and a preparation method of the micro-fluidic sensing chip. The micro-fluidic sensing chip based on induced electroosmotic flow comprises a sensing layer and a sample flow cell, the sensing layer comprises a glass substrate and an acceptor layer positioned on one side of the glass substrate; the sensing layer is used for being coupled to the prism through refractive index matching liquid and pressing the sample flow cell to the surface of the acceptor layer through the pressure of the spring to form a sample detection channel; the sample flow cell is located on the side, away from the glass substrate, of the receptor layer and comprises an electrode branch and a sample flow branch at least located on the side, away from the glass substrate, of the electrode layer; when an electric signal is applied to the electrode subsection, induced electroosmotic flow is generated in the sample flow cell to drive an analyte between the electrode subsection and the acceptor layer to move towards the acceptor layer, so that the flow velocity of the analyte near the acceptor layer can be accelerated, the concentration of an analysis sample on one side of the acceptor layer is improved, and the contact probability of the analyte and the acceptor is improved; and the sensitivity of the sensing chip is improved.

Description

Technical Field

[0001] The invention relates to the technical field of cell sensors, and in particular to a microfluidic sensor chip based on inductive electroosmotic flow and a preparation method thereof. Background Art

[0002] As a key technology in modern medicine, cell surface protein analysis technology plays a vital role in the early diagnosis, treatment monitoring and prognosis evaluation of cancer. Compared with traditional analytical diagnostic tools such as immunofluorescence and flow cytometry, label-free detection and analysis technology is revolutionizing the field of biological and chemical analysis because it can directly detect molecules without additional labeling, providing real-time, accurate and cost-effective analytical methods. The importance of this technology is reflected in its ability to provide early diagnosis, avoid complex sample processing, reduce costs and resource consumption, improve detection specificity, reduce invasiveness to patients, and promote the development of precision medicine. In addition, label-free detection technology promotes the realization of personalized treatment plans by discovering new biomarkers in scientific research and clinical applications.

[0003] Microfluidics provides a high-throughput, simple, integrated and customizable, low-cost solution for cell analysis and diagnosis, which has promoted the development of biomedical research and clinical diagnosis. However, at the microscopic scale, the surface effect of the flow cell affects liquid mixing and reaction dynamics, and the laminar flow characteristics of the analyte in the channel are mainly based on molecular diffusion, which limits the improvement of diagnostic efficiency.

[0004] In order to enhance the movement of molecules within microanalytes, the main method currently is to apply an external physical field and drive the analyte molecules by applying a pressure difference through means such as sound, light, heat, electricity, and magnetism. However, existing biosensor chips have the problem of low sensitivity caused by poor contact between the analysis sample and the detection substrate. Summary of the invention

[0005] The embodiment of the present invention provides a microfluidic sensor chip based on induced electroosmotic flow and a preparation method thereof, so as to accelerate the flow rate of the analyte to the receptor layer, increase the concentration of the analysis sample near the receptor layer, enhance the contact probability between the analyte and the receptor, and improve the sensitivity of the sensor chip.

[0006] In a first aspect, an embodiment of the present invention provides a microfluidic sensor chip based on inductive electroosmotic flow, comprising: a sensing layer and a sample flow cell;

[0007] The sensing layer includes a glass substrate and a receptor layer located on one side of the glass substrate; the sensing layer is used to couple to a prism through a refractive index matching liquid, and the sample flow cell is pressed to the surface of the receptor layer by the pressure of a spring to form a sample detection channel;

[0008] The sample flow cell is located on the side of the receptor layer away from the glass substrate, and includes an electrode portion and a first sample flow portion located on the side of the electrode portion away from the glass substrate; when an electrical signal is applied to the electrode portion, an induced electroosmotic flow is generated in the sample flow cell to drive the analyte between the electrode portion and the receptor layer to move toward the receptor layer.

[0009] Optionally, the sample circulation pool further includes a second sample circulation section;

[0010] The second sample flow section is located on a side of the electrode section close to the glass substrate.

[0011] Optionally, the thickness D1 of the second sample flow portion satisfies: 2 μm≤D1≤70 μm.

[0012] Optionally, the electrode division includes a first electrode division and a second electrode division;

[0013] The first electrode division includes a first main body and a first electrode branch that are connected to each other, and the second electrode division includes a second main body and a second electrode branch that are connected to each other; the first electrode branches and the second electrode branches are alternately arranged in sequence along a first direction and both extend along a second direction; the first direction intersects with the second direction.

[0014] Optionally, along the first direction, the width of the first electrode branch is equal to the width of the second electrode branch;

[0015] And the width H1 of the first electrode branch satisfies: 0.5 mm≤H1≤2 mm.

[0016] Optionally, the sample circulation cell includes an inlet and an outlet;

[0017] The analyte flows into the sample inlet and flows out of the sample outlet; the flow direction of the analyte is parallel to the first direction.

[0018] Optionally, along the first direction, a gap is provided between the first electrode branch and the second electrode branch that are adjacent to each other, and along the first direction, a width H2 of the gap satisfies: 0.25 mm ≤ H2 ≤ 0.5 mm;

[0019] The thickness D2 of the electrode portion satisfies: 0.1 mm≤D2≤4 mm.

[0020] Optionally, the glass substrate comprises borosilicate crown glass;

[0021] The receptor layer includes a gold film and an antibody modified by functionalization and located on a side of the gold film away from the glass substrate;

[0022] The thickness D3 of the gold film satisfies: 45nm≤D3≤50nm; the thickness D4 of the borosilicate crown glass satisfies: 0.5mm≤D4≤1nm.

[0023] Optionally, the microfluidic sensor chip based on inductive electroosmotic flow further comprises: a cadmium layer;

[0024] The cadmium layer is located between the gold film and the borosilicate crown glass.

[0025] In a second aspect, an embodiment of the present invention further provides a method for preparing a microfluidic sensor chip based on inductive electroosmotic flow, comprising:

[0026] Providing an electrode subdivision and a sensing layer; the sensing layer comprises a glass substrate and a receptor layer located on one side of the glass substrate;

[0027] Prepare a sample flow cell; the sample flow cell comprises the electrode section and a first sample flow section located on a side of the electrode layer away from the glass substrate;

[0028] Laminating the sample flow cell to the receptor layer;

[0029] An electrical signal is applied to the electrode portion to generate an induced electroosmotic flow in the sample flow cell, so as to drive the analyte between the electrode portion and the receptor layer to move toward the receptor layer.

[0030] The technical solution provided by the embodiment of the present invention is to set an electrode division. When an electrical signal is applied to the electrode division, an AC electric field will be generated in the sample circulation pool. At the place where the sample circulation pool contacts the analyte, since the surface of the sample circulation pool carries an induced charge, particles of opposite charge will be absorbed in the analyte to form a double electric layer. Under the action of the electric field, the analyte in the double electric layer and the charged cells on the membrane surface will be moved by the Coulomb force of the electric field. Under this irregular electric field, the vortex motion of the analyte is formed as the intensity of the AC electric field changes, which can break the traditional laminar diffusion restriction. The vortex motion of the analyte generated by this electric field drive accelerates the transport of the analyte to the receptor layer within a fixed time, and the concentration of the analysis sample near the receptor layer is increased by translating, rotating, etc. the analysis sample to the receptor layer, and the contact probability of the analyzed sample with the receptor layer receptor is increased, thereby improving the sensitivity of the sensor chip. BRIEF DESCRIPTION OF THE DRAWINGS

[0031] Figure 1 A schematic diagram of the structure of a microfluidic sensor chip based on inductive electroosmotic flow provided by an embodiment of the present invention;

[0032] Figure 2 A modeling diagram of a channel in a simulation experiment provided by an embodiment of the present invention;

[0033] Figure 3 A schematic diagram of the effect of different 0KV-3KV voltage changes on velocity distribution in numerical simulation provided by an embodiment of the present invention;

[0034] Figure 4 A scatter plot of the ratio of the velocity of the analyte to the initial velocity of the numerical simulation provided by an embodiment of the present invention as a function of voltage;

[0035] Figure 5 A schematic diagram of the effect of different 0KV-3KV voltage changes on the concentration distribution of the glass substrate position provided by the numerical simulation of the embodiment of the present invention;

[0036] Figure 6 for Figure 5 The corresponding numerical simulation is a curve diagram of the relationship between the concentration integral value and the voltage after integrating the concentration curve;

[0037] Figure 7 A schematic diagram of the effect of different changes in the frequency of electrical signals on the concentration distribution at the glass substrate position provided by the numerical simulation of the embodiment of the present invention;

[0038] Figure 8 A schematic diagram of the concentration distribution on the side of the flow cell channel within 1 second of the inductive electroosmotic flow chip numerically simulated under a 3KV voltage and a 10Hz frequency electrical signal provided in an embodiment of the present invention;

[0039] Fig. 9 Schematic diagram of concentration distribution of the control group without electrical signal loading;

[0040] Fig.10 A schematic diagram of the velocity distribution on the side of the flow cell channel of the sensor chip provided in an embodiment of the present invention under a 3KV voltage and a 10Hz frequency electrical signal;

[0041] Fig.11 It is a schematic diagram of the velocity distribution of the control group without electrical signal loading;

[0042] Fig.12 An optical path and overall architecture diagram of a SPRi system provided in an embodiment of the present invention;

[0043] Fig.13 A schematic diagram of system sensitivity detection when the sensor chip is not powered on in the SPRi system provided by an embodiment of the present invention;

[0044] Fig.14 A schematic diagram of system sensitivity detection when a sensor chip is powered on in a SPRi system provided by an embodiment of the present invention;

[0045] Fig.15 A schematic top view of an electrode division provided by an embodiment of the present invention;

[0046] Fig.16 A schematic diagram of numerical simulation of the effect of different electrode branch widths on flow velocity provided by an embodiment of the present invention;

[0047] Fig.17 A schematic diagram of a numerical simulation of the effect of different ratios of first electrode branch widths to second electrode branch widths on flow velocity provided by an embodiment of the present invention;

[0048] Fig.18 A schematic diagram of a numerical simulation of the effect of different gap widths on flow velocity provided by an embodiment of the present invention;

[0049] Fig.19 A schematic diagram of fluorescence imaging of fluorescent microspheres on a sensing substrate layer in a sensor chip without applying an electrical signal provided by an embodiment of the present invention;

[0050] Fig. 20 A schematic diagram of fluorescence imaging of fluorescent microspheres on a sensing substrate layer in a sensor chip with an electrical signal applied provided by an embodiment of the present invention;

[0051] Fig.21 A schematic diagram of the trajectory and velocity analysis of fluorescent microspheres in a sensor chip without applying an electrical signal provided by an embodiment of the present invention;

[0052] Fig. 22 A schematic diagram of the trajectory and velocity analysis of fluorescent microspheres in a sensor chip to which an electrical signal is applied provided by an embodiment of the present invention;

[0053] Fig.23 A schematic diagram of the detection limit analysis of the sensor chip system provided by an embodiment of the present invention;

[0054] Fig.24 A schematic flow chart of a method for preparing a microfluidic sensor chip based on induced electroosmotic flow provided in an embodiment of the present invention. DETAILED DESCRIPTION

[0055] The present invention will be further described in detail below in conjunction with the accompanying drawings and embodiments. It is to be understood that the specific embodiments described herein are only used to explain the present invention, rather than to limit the present invention. It should also be noted that, for ease of description, only parts related to the present invention, rather than all structures, are shown in the accompanying drawings.

[0056] Figure 1 A schematic diagram of the structure of a microfluidic sensor chip based on inductive electroosmotic flow provided by an embodiment of the present invention is shown in FIG. Figure 1As shown, the microfluidic sensor chip based on induced electroosmotic flow includes: a sensing layer 10 and a sample flow cell 20; the sensing layer 10 includes a glass substrate 101 and a receptor layer 102 located on one side of the glass substrate 101; the sensing layer 10 is used to couple to a prism through a refractive index matching liquid, and the sample flow cell 20 is pressed to the surface of the receptor layer 102 by the pressure of a spring to form a sample detection channel 30; the sample flow cell 20 is located on the side of the receptor layer 102 away from the glass substrate 101, and the sample flow cell 20 includes an electrode section 201 and a first sample flow section 2022 located on the side of the electrode section 201 away from the glass substrate 101; when an electrical signal is applied to the electrode section 201, an induced electroosmotic flow is generated in the sample flow cell 20 to drive the analyte 40 between the electrode section 201 and the receptor layer 102 to move toward the receptor layer 102.

[0057] Specifically, electroosmotic flow technology refers to the movement of analytes caused by the fixed charges on the inner wall of the microchannel and the double electric layer formed. Under the action of an external electric field, the ions in the double electric layer move, driving the surrounding analytes to move together. This flow can be used to drive analytes in microfluidic chips, including nucleic acids, proteins, cells, etc. Exemplarily, the analyte can be understood as an analytical sample.

[0058] Specifically, electrical signals of different strengths and frequencies may be applied to both ends of the electrode section 201 to drive the generation of induced electroosmotic flow. Figure 2 A modeling diagram of a channel in a simulation experiment provided by an embodiment of the present invention, Figure 3 A schematic diagram of the effect of different 0KV-3KV voltage changes on velocity distribution in numerical simulation provided by an embodiment of the present invention, Figure 4 The scatter plot of the ratio of the analyte velocity to the initial velocity as a function of voltage provided by the numerical simulation of the embodiment of the present invention is as follows: Figure 2-Figure 4 As shown in FIG. 1 , the width of the first electrode branch, the gap width between the first electrode branch and the second electrode branch, and the non-width of the second electrode branch are 2 mm, 0.5 mm, and 2 mm, respectively, to simulate the effect of the voltage amplitude of the applied electric signal on the analyte flow velocity at the sensing layer position. The analyte velocity increases with the increase of the voltage, and as shown in FIG. Figure 4 As shown, the flow rate changes linearly with the voltage.

[0059] Figure 5 A schematic diagram of the effect of different 0KV-3KV voltage changes on the concentration distribution of the glass substrate position provided by the numerical simulation of the embodiment of the present invention, Figure 6 for Figure 5 The corresponding numerical simulation is a graph of the relationship between the concentration integral value and the voltage after integrating the concentration curve. Figure 7 The numerical simulation provided by the embodiment of the present invention shows the influence of different changes in the frequency of electrical signals on the concentration distribution of the glass substrate. Figure 5-Figure 7As shown, by simulating the influence of the voltage parameters and frequency parameters of the electrical signal on the analyte concentration distribution at the sensing layer position, when the voltage of the applied electrical signal is ≥1 kV and the frequency of the electrical signal is ≥20 Hz, the particle concentration distribution in the channel tends to be stable and does not change much.

[0060] Specifically, Figure 8 A schematic diagram of the concentration distribution on the side of the flow cell channel within 1 second of the inductive electroosmotic flow chip numerically simulated under a 3KV voltage and a 10Hz frequency electrical signal provided in an embodiment of the present invention, Fig. 9 is a schematic diagram of the concentration distribution of the control group without electrical signal loading, such as Figure 8 and Fig. 9 As shown, the darker the red position, the higher the concentration. Figure 8 As shown, by applying an electrical signal, a vortex is generated in the flow cell, so that the concentration of the analyzed sample changes with the vortex, so that the sample concentration is high near the sensing layer, while Fig. 9 The sample concentration is low near the electrode and near the gold film.

[0061] Fig.10 This is a schematic diagram of the velocity distribution on the side of the flow cell channel of the sensor chip provided by the embodiment of the present invention under a 3KV voltage and a 10Hz frequency electrical signal. Fig.11 is a schematic diagram of the velocity distribution of the control group without electrical signal loading, such as Fig.10 and Fig.11 As shown, under the action of the electrical signal, the flow rate of the analyte to the glass substrate can be accelerated, thereby improving the sensitivity of the sensor chip.

[0062] Exemplarily, the electrical signal may have a voltage of 1000V-2000V, a current of 30μA-100μA, and a frequency of 20Hz-120Hz.

[0063] Specifically, the analyte 40 may be located in the sample detection channel 30, and the analyte may be understood as a sample to be detected. The analyte 40 may flow into the sample inlet through the first stainless steel needle 501, and flow out of the sample outlet through the second stainless steel needle.

[0064] Specifically, the sample flow cell 20 can be pressed to the surface of the receptor layer 102 by the pressure of the spring to form a sample detection channel 30, thereby realizing the preparation of a sensor chip, which can be applied to a surface plasmon resonance immunosensor (SPRi) system to realize label-free immunoassay. Fig.12 The optical path and overall architecture diagram of a SPRi system provided by an embodiment of the present invention, refer to Figure 1 and Fig.12The SPRi system includes a halogen lamp and a collimating lens L1, a narrowband filter F1, a polarizer P1, an SPRi cell sensor, an analyzer P2, a zoom imaging lens L2, and a complementary metal oxide semiconductor, which are arranged in sequence according to the propagation order of light. This intensity-type SPRi sensing system uses a halogen lamp as a light source. The white light generated by the halogen lamp is introduced into the optical path system of the SPRi through a multimode optical fiber, and then passes through the collimating lens L1 and the narrowband filter F1 to become monochromatic parallel light, and then passes through the polarizer P1 to become P polarized light (since S polarized light does not produce SPR effect, it is filtered out as stray light). The P polarized light is incident on the SPRi cell sensor at a certain angle to produce an SPR effect, and finally the reflected light is received by the complementary metal oxide semiconductor (CMOS) through the analyzer P2 and the zoom imaging lens L2 for imaging.

[0065] Specifically, since the electrode division 201 is arranged in the sample circulation pool 20, conductive clips are inserted at both ends of the electrode division 201, and a sample circulation pool with an induced electroosmotic flow function can be completed. In this way, when an electric signal is applied to both ends of the electrode division 201, an induced electroosmotic flow is generated in the sample circulation pool 20, that is, an alternating electric field is generated in the sample circulation pool 20. At the place where the sample circulation pool and the analyte are in contact, since the surface of the sample circulation pool 20 carries an induced charge, particles with opposite charges are absorbed in the analyte 40 to form a double electric layer. Under the action of the electric field, the analyte in the double electric layer and the charged cells on the membrane surface will be moved by the Coulomb force of the electric field. Under this irregular electric field, the vortex motion of the analyte is formed as the intensity of the alternating electric field changes, which can break the traditional laminar diffusion limitation. The vortex motion of the analyte generated under the drive of this electric field accelerates the transport of the analyte to the receptor layer within a fixed time. By making the analyte sample move in translation, rotation and other movements toward the receptor layer, the concentration of the analyte near the receptor layer is increased, and the contact probability between the analyte and the receptor layer receptor is increased, thereby improving the sensitivity of the sensor chip.

[0066] The microfluidic sensor chip based on induced electroosmotic flow provided by the embodiment of the present invention is provided with electrode divisions. When an electrical signal is applied to the electrode divisions, an alternating electric field is generated in the sample flow cell. At the place where the sample flow cell contacts the analyte, since the surface of the sample flow cell carries an induced charge, particles with opposite charges are absorbed in the analyte to form a double electric layer and generate electroosmotic flow. The driving effect of the electroosmotic flow can further increase the concentration of the analyzed sample near the receptor layer, increase the contact probability between the analyzed sample and the receptor layer receptor, and further improve the sensitivity of the sensor chip.

[0067] Optional, continue to refer to Figure 1The sample circulation pool 20 also includes a second sample circulation section 2021 ; the second sample circulation section 2021 is located on a side of the electrode section 201 close to the glass substrate 101 .

[0068] Specifically, the sample circulation cell can be a polydimethylsiloxane sample circulation cell. The first sample circulation subsection 2022 is located above the electrode subsection 201, that is, the second sample circulation subsection 2021 is located between the analyte 40 and the electrode subsection 201, so that the second sample circulation subsection 2021 can be used to insulate and prevent the analyte 40 from directly contacting the electrode subsection 201. As a comparative example, in the prior art, in order to achieve the accelerated motion of the electrically driven analyte, a low electrical signal is usually loaded into the sensing layer 10, so that the analyte is directly in direct contact with the sensing layer. During the experiment, the detection signal will be interfered by the bubbles generated by the electrolysis of water, and the negative impact of the electrically induced Joule heat on the protein cannot be controlled. In the embodiment of the present invention, the electrode portion 201 that generates an electric field is placed above the sample detection channel 30, and a second sample flow portion 2021 is provided between the electrode portion 201 and the sample detection channel 30. In this way, the second sample flow portion 2021 can act as an insulator to prevent the electrode portion 201 from directly contacting the protein in the analyte 40, that is, the electrode portion 201 and the analyte 40 are non-contact. This can reduce sample contamination and avoid the risk of protein inactivation caused by electrochemical reactions and thermal effects, thereby providing an efficient and reliable technical platform for biological label-free real-time analysis and solving the problems of protein denaturation caused by electrical manipulation and low efficiency of microfluidic biosensor systems.

[0069] Specifically, the first sample circulation section 2022 can be understood as a sample circulation section located above the electrode section 201. It should be noted that when preparing the sample circulation pool 20, polydimethylsiloxane glue doped with polytetrafluoroethylene of different concentrations is first prepared and poured into the circulation pool mold twice, and heated and cured at 150°C for 20 minutes and 200°C for 30 minutes respectively. Among them, 10%-50% of the total volume of glue is injected for the first time, and the designed electrode section 201 is placed on the glue surface. Taking the electrode section 201 as an example of being orthogonal to the direction of the circulation pool channel, the remaining volume of glue is injected. After curing and demolding, a sample circulation pool with an electrode section 201 embedded above the channel insulation layer is obtained, and conductive clips are inserted on both sides of the electrode section 201. Holes are punched at both ends of the channel of the initial sample circulation pool, and the sample detection channel and the microtube are connected by a stainless steel needle tube to obtain a sample circulation pool.

[0070] Specifically, in the process of using the sensor chip, the peak-to-peak value of the electric signal applied at both ends of the electrode division 201 is about 1 kV, the current is about 60 μA, the frequency is 80 Hz, and at least 3×10 4The cell suspension of 1.5 cells / mL was injected at a flow rate of 10 μL / min and an injection time of at least 10 min to study the specific capture of cells, calculate the change in SPR signal intensity, and study the detection limit and sensitivity of the system.

[0071] Fig.13 A schematic diagram of system sensitivity detection when the sensor chip is not powered on in the SPRi system provided by an embodiment of the present invention, Fig.14 A schematic diagram of system sensitivity detection when the sensor chip is powered on in the SPRi system provided by an embodiment of the present invention, such as Fig.13 and Fig.14 As shown, the sensor chip provided by the embodiment of the present invention is directly injected into the living cell suspension in the label-free SPRi cell immune sensing experiment. Fig.13 and Fig.14 The dashed lines in the figure represent the fitting of the three control group curves. The slopes of the dashed lines can represent the reaction speed. Fig.13 and Fig.14 It can be found that when powered on, the sensor chip provided by the embodiment of the present invention has a fast response speed. 6 cell / mL, it can distinguish human epidermal growth factor receptor 2 (HER2) positive cell lines and HER2 negative cell lines without labeling, quickly and accurately within 45 seconds. After three repeated experiments, the detection limit of the sensor chip of the present invention in detecting HER2 positive cells is 3×10 4 cell / mL, while the detection limit of traditional methods is 10 5 For the experimental results of traditional fluorescent-labeled flow cytometry and unprocessed label-free analysis sensor chips, the present invention can reduce the detection limit by one order of magnitude.

[0072] In clinical applications, overexpression of human epidermal growth factor receptor 2 is closely related to the aggressiveness, poor prognosis and sensitivity to certain therapeutic drugs of breast cancer. Therefore, its detection is of great significance for guiding clinical treatment decisions. HER2 was selected as a key molecular marker in the diagnosis and treatment of breast cancer as the target of research and detection. In the application results, the sensitivity and detection limit of the immunohistochemistry (IHC) detection method commonly used in clinical practice and the publicly published label-free SPRi immunosensor were compared horizontally. The present invention was used for SPRi cell sensors. The results of the embodiments of the present invention have the advantages of fast detection speed, label-free, high throughput, simple operation, etc. in the application. The results of repeated experiments further confirmed these advantages and showed a high degree of consistency and repeatability, which is of great significance for clinical diagnosis and biomedical research.

[0073] Optional, continue to refer to Figure 1, the thickness D1 of the second sample circulation section 2021 satisfies: 2μm≤D1≤70μm.

[0074] Exemplarily, when D1 is less than 2 μm, it indicates that the thickness of the second sample flow section 2021 is relatively thin, so that the distance between the electrode section 201 and the analyte 40 is relatively close, and the electrically induced Joule heat has a greater impact on the protein, which in turn affects the detection accuracy of the sensor chip. Exemplarily, when D1 is greater than 70 μm, it indicates that the distance between the electrode section 201 and the analyte 40 is relatively far, and when a voltage is applied to both ends of the electrode section 201, it is not conducive to the analyte moving toward the glass substrate under the action of the AC electric field, and thus the flow rate of the analyte to the glass substrate cannot be accelerated, and the concentration of the analyte near the glass substrate cannot be guaranteed, which in turn affects the detection accuracy of the sensor chip. The technical solution provided in the embodiment of the present invention sets 2μm≤D1≤70μm, so that the distance between the electrode portion 201 and the analyte 40 is moderate. On the one hand, the insulation effect between the electrode portion 201 and the analyte 40 can be ensured through the second sample flow portion 2021, thereby preventing the electrode portion from directly contacting the analyte and causing protein failure. On the other hand, the driving ability for the analyte to flow to the receptor layer can be improved, thereby improving the sensitivity of the sensor chip.

[0075] Optional, Fig.15 A schematic top view of an electrode division provided in an embodiment of the present invention, such as Figure 4 As shown, the electrode division 201 includes a first electrode division 2011 and a second electrode division 2012; the first electrode division 2011 includes a first main body 20111 and a first electrode branch 20112 that are connected, and the second electrode division 2012 includes a second main body 20121 and a second electrode branch 20122 that are connected; the first electrode branch 20112 and the second electrode branch 20122 are alternately arranged in sequence along a first direction (the X direction shown in the figure) and both extend along a second direction (the Y direction shown in the figure); the first direction X intersects with the second direction Y.

[0076] Specifically, the first electrode subdivision 2011 and the second electrode subdivision 2012 can be connected to the positive and negative poles of the voltage respectively, so that the electrical signal can be applied to the electrode subdivision 201 to generate electroosmosis and drive the flow rate of the analyte to the glass substrate 101. Specifically, the first main body 20111 and the second main body 20121 are arranged along the second direction Y, and the first electrode branch 20112 is connected to the first main body 20111, and the second electrode branch 20122 is connected to the second main body 20121, so that the first electrode subdivision 2011 and the second electrode subdivision 2012 both form a shape similar to "comb teeth". The first electrode branch 20112 and the second electrode branch 20122 are alternately arranged in sequence along the first direction X and both extend along the second direction Y, that is, the second electrode branch 20122 is located between two adjacent first electrode branches 20112, so that the electric field strength can be ensured, thereby accelerating the flow rate of the analyte to the receptor layer, increasing the concentration of the analyzed sample in the receptor layer, increasing the contact probability between the analyte and the receptor, and improving the sensitivity of the sensor chip.

[0077] Exemplarily, the material of the electrode portion 201 may be indium tin oxide.

[0078] Optional, continue to refer to Fig.15 , along the first direction X, the width of the first electrode branch 20112 is equal to the width of the second electrode branch 20122; and the width H1 of the first electrode branch 20112 satisfies: 0.5mm≤H1≤2mm.

[0079] Specifically, along the first direction X, the width of the first electrode branch 20112 is equal to the width of the second electrode branch 20122. In this way, when a voltage is applied to both ends of the first electrode branch 20112 and the second electrode branch 20122, the electric field generated by the electrode portion 201 can be ensured to be uniform, thereby ensuring that the speed at which the analytes on one side of the first electrode branch 20112 and the second electrode branch 20122 move toward the receptor layer 102 is consistent, thereby ensuring the sensitivity of the sensor chip and preventing the different widths of the first electrode branch 20112 and the second electrode branch 20122 from causing a speed difference in the process of the analytes on one side of the first electrode branch 20112 and the second electrode branch 20122 moving toward the receptor layer 102, thereby affecting the sensitivity of the sensor chip.

[0080] Exemplarily, when H1 is less than 0.5 mm, it indicates that the width of the first electrode branch 20112 and the second electrode branch 20122 is relatively small, which is not conducive to the preparation of the electrode branch. When H1 is greater than 2 mm, it indicates that the width of the first electrode branch 20112 and the second electrode branch 20122 is relatively large, which is not conducive to matching the size of the electrode branch with the sample flow cell, thereby affecting the function of the sensor chip. The embodiment of the present invention sets 0.5 mm ≤ H1 ≤ 2 mm, indicating that the width of the first electrode branch and the second electrode branch is moderate, which can ensure the simplicity of the preparation process of the electrode branch on the one hand, and the performance of the sensor chip on the other hand.

[0081] Specifically, in the process of designing the electrode division, the width of the first electrode branch was set to 0.1 mm, 0.25 mm, 0.5 mm, 1.00 mm, 1.50 mm and 2.00 mm respectively for multiple control tests to determine the effect of the electrode branch width on the flow rate. Fig.16 A schematic diagram of numerical simulation of the effect of different electrode branch widths on flow velocity provided by an embodiment of the present invention, such as Fig.16 As shown, when the width of the first electrode portion is 0.5 mm ≤ H1 ≤ 2 mm, the flow rate of the analyte to the glass substrate can be accelerated, the concentration of the analysis sample on one side of the glass substrate can be increased, and the sensitivity of the sensor chip can be improved. Fig.17 A schematic diagram of a numerical simulation of the influence of different ratios of first electrode branch widths to second electrode branch widths on flow velocity provided by an embodiment of the present invention, such as Fig.17 As shown, the ratio of the first electrode branch width to the second electrode branch width is 1:1, 2:1, 3:1, 4:1, 3:2 and 3:5, which can accelerate the flow rate of the analyte to the glass substrate, increase the concentration of the analysis sample on one side of the glass substrate, and improve the sensitivity of the sensor chip.

[0082] Optional, continue to refer to Figure 1 The sample flow cell 20 includes an inlet and an outlet; the analyte 40 flows in at the inlet 100 and flows out at the outlet 200; the flow direction of the analyte 40 is parallel to the first direction X.

[0083] Specifically, the analyte 40 flows into the sample inlet through the first stainless steel needle 501 and flows out of the sample outlet through the second stainless steel needle 502. The flow direction of the analyte 40 can be understood as the direction from the sample inlet to the sample outlet. Specifically, the flow direction of the analyte 40 is parallel to the first direction X, that is, the flow direction of the analyte 40 is parallel to the arrangement direction of the first electrode branch and the second electrode branch.

[0084] Exemplarily, the flow direction of the analyte 40 is parallel to the extension direction of the first electrode branch and the second electrode branch, so that on the one hand, diversified settings of the sensor chip can be achieved, and on the other hand, the performance of the sensor chip can be guaranteed.

[0085] Exemplarily, the sample flow cell may include at least 4 channels. Each channel includes an inlet and an outlet. The material of the sample flow cell is silicone rubber doped with materials of different concentrations; the doped material is polytetrafluoroethylene powder with particle sizes of nanometers and 10 micrometers, and the doping concentrations are 0%, 1%, 5%, 10%, 15%, and 20%, respectively.

[0086] Optional, continue to refer to Figure 1 and Fig.15 Along the first direction X, there is a gap 2013 between the adjacent first electrode branch 20112 and the second electrode branch 20122, and along the first direction X, the width H2 of the gap 2013 satisfies: 0.25mm≤H2≤0.5mm; the thickness D2 of the electrode portion 201 satisfies: 0.1mm≤D2≤4mm.

[0087] Specifically, in the process of designing the electrode division, the width of the gap 2013 between the first electrode branch 20112 and the second electrode branch 20122 was set to 0.25mm, 0.50mm, 1.00mm, 1.50mm, 2.00mm and 2.50mm respectively for multiple control experiments to determine the effect of the gap width on the flow rate. Fig.18 A schematic diagram of a numerical simulation of the effect of different gap widths on flow velocity provided by an embodiment of the present invention, such as Fig.18 As shown, when the width of the gap 2013 is set to 0.25 mm ≤ H2 ≤ 0.5 mm, the flow rate of the analyte to the glass substrate can be accelerated, the concentration of the analysis sample on one side of the glass substrate can be increased, and the sensitivity of the sensor chip can be improved.

[0088] For details, please refer to Figure 1 The thickness D2 of the electrode portion 201 satisfies: 0.1 mm ≤ D2 ≤ 4 mm. This can ensure that the preparation process of the electrode portion 201 is simple on the one hand, and the performance of the sensor chip on the other hand.

[0089] Optional, continue to refer to Figure 1 The glass substrate 101 includes borosilicate crown glass; the receptor layer 102 includes a gold film and a functionally modified antibody located on a side of the gold film away from the glass substrate 101; the thickness D3 of the gold film satisfies: 45nm≤D3≤50nm; the thickness D4 of the borosilicate crown glass satisfies: 0.5mm≤D4≤1nm.

[0090] Specifically, a layer of gold film is deposited on a borosilicate crown glass having a size of 1 mm*18 mm*18 mm by magnetron sputtering. Exemplarily, the thickness of the gold film may be 47 nm.

[0091] For example, when the thickness D3 of the gold film is greater than 50 nm, it indicates that the thickness of the gold film is too large, so the difference in light reflection and refraction is relatively large. When the thickness D3 of the gold film is less than 45 nm, it indicates that the thickness of the gold film is too small, so it is difficult to plate the gold film on the borosilicate crown glass by magnetron sputtering, the process is difficult, and the thickness of the gold film is difficult to control. The embodiment of the present invention sets 45 nm ≤ D3 ≤ 50 nm, so that on the one hand, it matches the preparation process of the gold film and reduces the process difficulty, and on the other hand, it can ensure the reflection and refraction efficiency of the receptor layer 102 to light.

[0092] Specifically, the thickness D4 of the borosilicate crown glass satisfies: 0.5 mm ≤ D4 ≤ 1 nm. Thus, the thickness of the borosilicate crown glass is moderate and can ensure the stability of the substrate.

[0093] Optional, continue to refer to Figure 1 The microfluidic sensor chip based on induced electroosmotic flow also includes: a cadmium layer 60; the cadmium layer 60 is located between the gold film and the borosilicate crown glass.

[0094] Specifically, since the gold film has poor adhesion to the borosilicate crown glass and is easy to fall off, a cadmium layer is plated between the gold film and the glass substrate to increase the stability of the gold film.

[0095] Specifically, functional modification is performed on the gold film to form a receptor layer.

[0096] Exemplarily, the thickness of the cadmium layer 60 may be 2 nm.

[0097] Fig.19 A schematic diagram of fluorescence imaging of fluorescent microspheres on a sensing substrate layer in a sensor chip without applying an electrical signal provided by an embodiment of the present invention, Fig. 20 A schematic diagram of fluorescence imaging of fluorescent microspheres on a sensing substrate layer in a sensor chip with an electrical signal applied provided by an embodiment of the present invention, Fig.21 A schematic diagram of the trajectory and velocity analysis of fluorescent microspheres in a sensor chip without applying an electrical signal provided by an embodiment of the present invention, Fig. 22 The schematic diagram of the trajectory and velocity analysis of the fluorescent microspheres in the sensor chip to which the electrical signal is applied according to the embodiment of the present invention is as follows: Figure 19-22As shown in the figure, in the same time, i.e., 50s, the analysis sample without the application of the electric signal only moved 85μm, but after the electric signal was applied to the electrode portion, the analysis sample could move 240μm in the same time. It can be seen that the statistical analysis of the movement trajectory data of the fluorescent microspheres can be used to conclude that the application of the inductive electroosmotic chip in the present invention accelerates the movement speed of the fluorescent microspheres. Under the action of the electric signal, the flow speed of the analyte to the receptor layer can be accelerated, thereby improving the sensitivity of the sensor chip.

[0098] In the experiment to verify the effect of the inductive electroosmotic flow driving effect of the chip of the present invention on the distribution of the fluorescent flow rate of the microspheres, fluorescent microspheres with a diameter of 5 μm were selected as the experimental objects. A fluorescent microsphere solution with a concentration of 50 μg / mL was injected into the circulation pool at a flow rate of 15 μL / min. An electrical signal with a peak-to-peak value of 1 kV was applied to both ends of the chip, the current was controlled at 30-60 μA, and the frequency was adjusted between 20-80 Hz. A charge-coupled device camera was used to collect 50 s of fluorescent signals at a speed of 50 frames / s. The collected signals were analyzed and processed by data analysis software to obtain the motion trajectory of the fluorescent microspheres near the antibody sensing layer, and the speed of the particles under inductive electroosmotic flow treatment and without treatment was analyzed and compared.

[0099] Continue to refer Figure 19-22 In the experiment to verify the effect of the induced electroosmotic driving effect of the chip of the present invention on the fluorescence intensity distribution of the microspheres, fluorescent microspheres with a diameter of 5 μm were selected as the experimental objects. A fluorescent microsphere solution with a concentration of 50 μg / mL was injected into the flow cell at a flow rate of 15 μL / min. An electrical signal with a peak-to-peak value of 1 kV was applied to both ends of the chip, the current was controlled at 30 μA-60 μA, and the frequency was adjusted between 20 Hz-80 Hz. The Z-stack module function of the confocal microscope was used to scan the 1000 nm high flow cell from bottom to top, and a picture was collected at a height interval of 10 nm. A total of 100 pictures were collected, and the sum of the fluorescence intensity in each picture was calculated. The horizontal axis was the Z-axis position, and the vertical axis was the fluorescence intensity to draw the fluorescent microsphere concentration distribution curve, and the sum of the fluorescence intensity of the particles under induced electroosmotic flow treatment and without treatment was analyzed and compared. The results show that under the same experimental conditions, according to the statistical results, the application of the present invention makes the fluorescence intensity of the microspheres in the channel 1.5 times that of the untreated control group, indicating that the application of the present invention successfully drives more fluorescent microspheres, which are pre-concentrated in the sample channel, increasing the contact probability of antigen and antibody. This experimental result also consistently verifies the previous simulation results.

[0100] Fig.23 A schematic diagram of the detection limit analysis of the sensor chip system provided in an embodiment of the present invention is shown in FIG. Fig.23 As shown, the cell suspension was serially diluted to prepare 3×10 5 cells / mL and 3×10 4cells / mL of BT-474 (HER2 positive) cell suspension, and 5×10 5 cells / mL of MCF-10A (HER2 negative expression) cell suspension. The above three cell suspensions were injected into the circulation pool at the same time, with a flow rate of 10μL / min. The injection time was 10 minutes without treatment. The buffer was changed to PBS for 10 minutes to remove the cells that were not captured. Then, an electric signal with a peak-to-peak value of about 1kV and a current of about 60μA was applied to both ends of the chip treated with induced electroosmotic flow. The three cell suspensions were repeatedly injected under the conditions of 80Hz for 10 minutes. According to the changes in the SPR signal, the changes in signal intensity under different concentrations, different types and different conditions were calculated, and the data were statistically analyzed to obtain the detection limit of the system. The experimental results are as follows Fig.23 As shown, it can be distinguished from 3×10 4 The sensing signal of BT-474 (HER2 positive) cells with a concentration of 5×10 5 cells / mL of MCF-10A (HER2 negative expression) cells. The application of the present invention can make the lower limit of cell concentration detection reach 3×10 4 cell / mL, which is an order of magnitude lower than the detection limit of existing immunosensors.

[0101] In summary, the microfluidic sensor chip based on induced electroosmotic flow provided by the embodiment of the present invention, by setting the electrode division, when the electrical signal is applied to the electrode division, an AC electric field will be generated in the sample circulation pool, where the sample circulation pool and the analyte are in contact, because the surface of the sample circulation pool carries an induced charge, so that particles of opposite charge are absorbed in the analyte to form a double electric layer. Under the action of the electric field, the analyte in the double electric layer and the charged cells on the membrane surface will be moved by the Coulomb force of the electric field. Under this irregular electric field, the vortex motion of the analyte is formed as the intensity of the AC electric field changes, which can break the traditional laminar diffusion restriction. The vortex motion of the analyte generated under this electric field drive accelerates the transport of the analyte to the receptor layer within a fixed time, and the concentration of the analysis sample near the receptor layer is increased by analysing the sample to do translation, rotation and other movements to the receptor layer, and the contact probability of the analyzed sample with the receptor layer receptor is increased, thereby improving the sensitivity of the sensor chip. In addition, unlike the defects of protein inactivation or electrode degradation caused by direct contact or direct loading of electrodes into the sensing layer in the past, the design of the electrode division in the embodiment of the present invention avoids the negative impact of electrothermal effect or electrolytic reaction on antigen and antibody, thus ensuring the accuracy of the detection system.

[0102] Based on the same inventive concept, the embodiment of the present invention also provides a method for preparing a microfluidic sensor chip based on inductive electroosmotic flow. Fig.24A schematic diagram of a process for preparing a microfluidic sensor chip based on electroosmotic flow according to an embodiment of the present invention is shown in FIG. Fig.24 As shown, the preparation method comprises:

[0103] S101, providing an electrode portion and a sensing layer; the sensing layer includes a glass substrate and a receptor layer located on one side of the glass substrate.

[0104] For details, please refer to Fig.15 The preparation method of the electrode division 201 may be to first theoretically simulate the effects of electrode width, gap, and electrode ratio on the analyte velocity and analyte substance concentration under the same electric field conditions, and design an indium tin oxide conductive glass etching electrode based on the theoretical simulation results. According to the mold size of the circulation pool, the electrode size provided in the embodiment of the present invention may be 17mm*10mm*0.7mm. The electrode branch width obtained by etching the electrode may be 0.5mm-2mm, and the gap may be 0.25mm-0.5mm. The ratio of the first electrode branch width to the second electrode branch width may be 1:1, 2:1, 3:1, 4:1, 3:2, and 3:5. The theoretical simulation model was established using finite element analysis software, and numerical simulations were performed on the side of the circulation pool containing the chip. The effects of changes in electrode width, electrode gap, and distribution ratio on changes in analyte velocity at the position of the sensing layer in the channel were simulated, such as Figure 16-18 As shown, the simulation results show that when the electrode width is 2mm-2.5mm, the electrode gap is 0.25mm-0.5mm, and the ratio of the first electrode branch to the second electrode branch is 1:1 or 2:1, the analyte velocity at the sensing layer position is the largest.

[0105] Specifically, at room temperature, the gold film is incubated in a 1 mol / L glutathione reduced solution for 6-8 hours to complete glutathione functionalization; a phosphate buffer solution is injected into the substrate that has completed glutathione functionalization at a flow rate of 10 μL / min through a peristaltic pump for 10 minutes, and a phosphate buffer solution with a concentration of 50 μg / mL nanoantibody protein for 30 minutes and 10 minutes, respectively, to complete the fixation of the nanoantibody; finally, a 1% bovine serum albumin solution is passed for 20-30 minutes, and then a phosphate buffer solution is injected for 20 minutes to 30 minutes to complete the blocking of the sensing layer. The receptor layer is prepared through the steps of functionalization-antibody fixation-washing-blocking-re-washing.

[0106] S102, preparing a sample flow cell; the sample flow cell comprises an electrode portion and a first sample flow portion located on a side of the electrode layer away from the glass substrate.

[0107] For details, please refer to Figure 1When preparing the sample circulation cell 20, first prepare polydimethylsiloxane glue doped with polytetrafluoroethylene of different concentrations, pour it into the circulation cell mold twice, and heat and cure it at 150°C for 20 minutes and 200°C for 30 minutes respectively. Among them, 10%-50% of the total volume of glue is injected for the first time, and the designed electrode division 201 is placed on the glue surface. Taking the electrode division 201 as an example of being orthogonal to the direction of the circulation cell channel, the remaining volume of glue is injected. After curing and demolding, a sample circulation cell with an electrode division 201 embedded above the channel insulation layer is obtained, and conductive clips are inserted on both sides of the electrode division 201. Holes are punched at both ends of the channel of the initial sample circulation cell, and the sample detection channel and the microtube are connected by a stainless steel needle tube to obtain a sample circulation cell.

[0108] S103, attaching the sample flow cell to the receptor layer.

[0109] Specifically, the receptor layer is closely attached to a sample flow cell to obtain a microfluidic sensor chip for sensing electroosmotic flow, and its liquid sealing performance is tested.

[0110] Exemplarily, the test method for liquid sealing can be to inject water into the sample detection channel to check whether there is water leakage. If there is no water leakage, it means that the sealing of the sensor chip is good, otherwise the sensor chip needs to be repaired.

[0111] S104, applying an electrical signal to the electrode portion to generate an induced electroosmotic flow in the sample flow cell to drive the analysis sample between the electrode portion and the receptor layer to move toward the receptor layer.

[0112] For details, please refer to Figure 1 , since the electrode division 201 is arranged in the sample circulation pool 20, when a voltage is applied to both ends of the electrode division 201, an induced electroosmotic flow is generated in the sample circulation pool 20, that is, an AC electric field is generated in the sample circulation pool 20. At the place where the sample circulation pool and the analyte are in contact, since the surface of the sample circulation pool 20 carries an induced charge, particles of opposite charge are absorbed in the analyte 40 to form a double electric layer. Under the action of the electric field, the analyte in the double electric layer and the charged cells on the membrane surface are moved by the Coulomb force of the electric field. Under this irregular electric field, the vortex motion of the analyte is formed as the intensity of the AC electric field changes, which can break the traditional laminar diffusion restriction. The vortex motion of the analyte generated under the drive of this electric field accelerates the transport of the analyte to the receptor layer within a fixed time, and the concentration of the analysis sample near the receptor layer is increased by translating the analysis sample to the receptor layer, rotating and other movements, and increasing the contact probability of the analyzed sample with the receptor layer receptor, thereby improving the sensitivity of the sensor chip.

[0113] For example, continue to refer to Figure 1 and Fig.15, taking 0% undoped polydimethylsiloxane glue material as an example, according to the size of the mold, the electrode section with a thickness of 0.7mm is processed, the size is 17mm*10mm, the electrode width is 2mm, the electrode gap is 0.5mm, and the electrode ratio is 1:1. When making, add 220μL of polydimethylsiloxane glue to the mold, place the conductive glass electrode directly above the flow cell channel direction (taking orthogonal placement as an example), heat and cure at 200℃ for 20 minutes, inject 780μL of polydimethylsiloxane glue after curing, and continue to cure at 200℃ for 30min. Figure 1 As shown, after curing and demolding, a flow cell with electrodes embedded above the channel insulation layer is obtained, and conductive clips are inserted on both sides of the electrodes. Holes are punched at both ends of the channel of the initial sample flow cell, and the flow cell sample channel and microtube are connected through a stainless steel needle tube to obtain a microfluidic flow cell with inductive electroosmosis function, which is subsequently used in the SPRi cell sensing system to achieve label-free immunoassay.

[0114] The preparation method of the microfluidic sensor chip based on induced electroosmotic flow provided by the embodiment of the present invention is to prepare an electrode division. When an electrical signal is applied to the electrode division, an AC electric field will be generated in the sample circulation pool. In the place where the sample circulation pool contacts the analyte, since the surface of the sample circulation pool carries an induced charge, particles of opposite charge will be absorbed in the analyte to form a double electric layer. Under the action of the electric field, the analyte in the double electric layer and the charged cells on the membrane surface will be moved by the Coulomb force of the electric field. Under this irregular electric field, the vortex motion of the analyte is formed as the intensity of the AC electric field changes, which can break the traditional laminar diffusion restriction. The vortex motion of the analyte generated under this electric field drive accelerates the transport of the analyte to the receptor layer within a fixed time, and the concentration of the analysis sample near the receptor layer is increased by translating the analysis sample to the receptor layer, rotating and other movements, and increasing the contact probability of the analyzed sample with the receptor layer receptor, thereby improving the sensitivity of the sensor chip.

[0115] Note that the above are only preferred embodiments of the present invention and the technical principles used. Those skilled in the art will understand that the present invention is not limited to the specific embodiments described herein, and that various obvious changes, readjustments, combinations and substitutions can be made by those skilled in the art without departing from the scope of protection of the present invention. Therefore, although the present invention has been described in more detail through the above embodiments, the present invention is not limited to the above embodiments, and may include more other equivalent embodiments without departing from the concept of the present invention, and the scope of the present invention is determined by the scope of the appended claims.

Claims

1. A microfluidic sensor chip based on inductive electroosmotic flow, characterized in that: include: Sensing layer and sample flow cell; The sensing layer includes a glass substrate and a receptor layer located on one side of the glass substrate; The sensing layer is used to couple to the prism through a refractive index matching liquid, and the sample flow cell is pressed to the surface of the receptor layer by the pressure of a spring to form a sample detection channel; The sample flow cell is located on the side of the receptor layer away from the glass substrate, and includes an electrode portion and a first sample flow portion located on the side of the electrode portion away from the glass substrate; when an electrical signal is applied to the electrode portion, an induced electroosmotic flow is generated in the sample flow cell to drive the analyte between the electrode portion and the receptor layer to move toward the receptor layer.

2. The microfluidic sensor chip based on inductive electroosmotic flow according to claim 1, characterized in that: The sample circulation cell also includes a second sample circulation subsection; The second sample flow section is located on a side of the electrode section close to the glass substrate. 3 . The microfluidic sensor chip based on induced electroosmotic flow according to claim 2 , wherein the thickness D1 of the second sample flow portion satisfies: 2 μm≤D1≤70 μm.

4. The microfluidic sensor chip based on inductive electroosmotic flow according to claim 1, characterized in that: The electrode subdivision includes a first electrode subdivision and a second electrode subdivision; The first electrode division includes a first main body and a first electrode branch that are connected to each other, and the second electrode division includes a second main body and a second electrode branch that are connected to each other; the first electrode branches and the second electrode branches are alternately arranged in sequence along a first direction and both extend along a second direction; the first direction intersects with the second direction.

5. The microfluidic sensor chip based on inductive electroosmotic flow according to claim 4, characterized in that: Along the first direction, the width of the first electrode branch is equal to the width of the second electrode branch; And the width H1 of the first electrode branch satisfies: 0.5 mm≤H1≤2.5 mm.

6. The microfluidic sensor chip based on inductive electroosmotic flow according to claim 4, characterized in that: The sample circulation pool comprises a sample inlet and a sample outlet; The analyte flows into the sample inlet and flows out of the sample outlet; the flow direction of the analyte is parallel to the first direction.

7. The microfluidic sensor chip based on inductive electroosmotic flow according to claim 4, characterized in that: Along the first direction, a gap is provided between the first electrode branch and the second electrode branch that are adjacent to each other, and along the first direction, a width H2 of the gap satisfies: 0.25 mm ≤ H2 ≤ 0.5 mm; The thickness D2 of the electrode portion satisfies: 0.1 mm≤D2≤4 mm.

8. The microfluidic sensor chip based on inductive electroosmotic flow according to claim 1, characterized in that: The glass substrate comprises borosilicate crown glass; The receptor layer includes a gold film and an antibody modified by functionalization and located on a side of the gold film away from the glass substrate; The thickness D3 of the gold film satisfies: 45nm≤D3≤50nm; the thickness D4 of the borosilicate crown glass satisfies: 0.5mm≤D4≤1nm.

9. The microfluidic sensor chip based on inductive electroosmotic flow according to claim 8, characterized in that: The microfluidic sensor chip based on inductive electroosmotic flow further comprises: a cadmium layer; The cadmium layer is located between the gold film and the borosilicate crown glass.

10. A method for preparing a microfluidic sensor chip based on inductive electroosmotic flow, characterized in that: include: Providing an electrode subdivision and a sensing layer; the sensing layer comprises a glass substrate and a receptor layer located on one side of the glass substrate; Prepare a sample flow cell; the sample flow cell comprises the electrode section and a first sample flow section located on a side of the electrode layer away from the glass substrate; Laminating the sample flow cell to the receptor layer; An electrical signal is applied to the electrode portion to generate an induced electroosmotic flow in the sample flow cell, so as to drive the analyte between the electrode portion and the receptor layer to move toward the receptor layer.

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