Biovesicle surface biomarker detection device

Abstract

The present invention addresses the problem of providing a system, device, and method for quantitatively analyzing, from among test targets resulting from a liquid biopsy, an expression amount of target marker molecules such as a protein present on a surface, by detecting and measuring biological vesicles, having a size of 20 microns or less and including cells, exosomes, and extracellular vesicles, from among body fluid-derived samples.　By capturing, on a support, a body fluid-derived sample containing the plurality of target marker molecules, and by detecting and measuring the plurality of target marker molecules via a distinguishing substance such as an antibody, it became clear that it is possible to provide a method for qualitatively and quantitatively analyzing the expression of the plurality of target marker molecules, and a device and system therefor.

Description

[Technical Field]

[0001] The present invention provides a method, device, and system for detecting and measuring biological vesicles in a body fluid-derived sample, and quantitatively analyzing the expression level of target marker molecules, such as proteins, present on the surface of the vesicles. [Background technology]

[0002] As we enter an aging society, people are increasingly concerned about health and longevity, and the development of diagnostic methods and devices for non-invasive, minimally invasive disease detection is anticipated. In this context, liquid biopsy, a diagnostic method using body fluids, has attracted attention as a less invasive diagnostic technique compared to conventional cytology and tissue diagnosis, particularly in the field of cancer. Currently, the main detection targets for liquid biopsy are cells and nucleic acids circulating in body fluids, including circulating tumor cells (CTCs), circulating DNA (cfDNA), circulating tumor DNA (ctDNA), exosomes, and microRNAs (miRNAs).

[0003] Liquid biopsy testing is a less invasive method compared to conventional biopsies that involve tissue sampling, and has various clinical significance, including disease screening tests, follow-up observation and prognosis estimation, and supporting doctors in deciding on treatment plans. Because it allows for painless testing, it is expected to increase the rate of cancer screening, leading to early cancer detection, improved cure rates, and an extension of healthy life expectancy.

[0004] The aim is to analyze the time-series data obtained using these minimally invasive methods using artificial intelligence (AI) and other methods, and use it for ultra-early detection of disease, prognosis monitoring, and even understanding health conditions.

[0005] Among the test targets for liquid biopsy, methods for detecting cells and biologically derived vesicles (exosomes or extracellular vesicles) include ELISA, FACS, nanoparticle tracking analysis, and NTA. However, optical detection methods are the mainstream, and these detection methods have problems such as complicated sample preparation and difficulty in miniaturizing the measurement device. Compared to traditional biopsies that require tissue collection, liquid biopsy is a less invasive method and has various clinical significance, such as disease screening tests, follow-up observation and prognosis estimation, and supporting physicians' treatment plan decisions.

[0006] Previously, various inventions have been known, including a biosensor device and a biosensor that immobilizes viruses or bacteria using oligosaccharides bound to a support, allowing for selective and highly sensitive detection of viruses or bacteria (Patent Document 1), and a method that applies the principle of sandwich ELISA to immobilize a target protein on a sensor using an antibody bound to a support, and then detects the presence of this target protein by pH changes (Non-Patent Document 1).However, conventional methods have not been able to simply and directly detect target marker molecules, such as biomarkers on the surface of biological vesicles contained in body fluid-derived samples, without pretreatment. [Prior art documents] [Patent documents]

[0007] [Patent Document 1] WO2017 / 126617 [Non-patent literature]

[0008] [Non-Patent Document 1] Biosensors and Bioelectronics 117 (2018) 175-182 Summary of the Invention [Problem to be solved by the invention]

[0009] The present invention aims to provide a method, device, and system for detecting and measuring biological vesicles of 20 microns or less in size, including cells, exosomes, and extracellular vesicles, from bodily fluid-derived samples among the test subjects for liquid biopsy, and for quantitatively analyzing the expression levels of target marker molecules, such as proteins, present on their surfaces. [Means for solving the problem]

[0010] In order to solve the above-mentioned problems, the inventors of the present invention have revealed that it is possible to provide a method for qualitatively and quantitatively analyzing the expression of multiple target marker molecules, as well as a device and system for this purpose, by capturing a body fluid-derived sample containing multiple target marker molecules on a support and detecting and measuring the multiple target marker molecules using an identifying substance such as an antibody.

[0011] More specifically, the present application provides the following aspects to solve the above-mentioned problems: [1] A step of capturing a body fluid-derived sample containing a target marker molecule on a support; binding the captured sample with an identification substance that specifically recognizes the target marker molecule in the sample; detecting or measuring a pH change or an optical change caused by the identifying substance bound to the target marker molecule; A method for detecting or measuring a target marker molecule in a body fluid-derived sample based on a pH change or an optical change, comprising: [2] The method according to [1], wherein a single body fluid-derived sample contains multiple target marker molecules. [3] The method according to [2], in which the step of binding one type of identifying substance to a single body fluid-derived sample and the step of detecting or measuring the resulting pH change or optical change are repeated multiple times. [4] The method according to [3], wherein the enzyme bound to the plurality of recognition substances used in the plurality of steps is one type. [5] The method described in [2], in which a plurality of identifying substances are bound to a single body fluid-derived sample, and then the steps of detecting or measuring the pH change or optical change caused by each of the identifying substances are repeated multiple times. [6] The method according to [5], wherein the enzymes bound to the recognition substances are different enzymes for each of a plurality of different recognition substances. [7] The method according to any one of [1] to [6], wherein the identifying substance is an antibody that specifically binds to the target marker molecule. [8] The method according to any one of [1] to [7], wherein the body fluid-derived sample is a cell or a biological vesicle. [9] The method according to [8], wherein the biologically derived vesicles are exosomes or extracellular vesicles.

[10] The method according to any one of [1] to [9], wherein the pH change or optical change is caused by a substrate-enzyme reaction.

[11] Supports for capturing samples derived from body fluids; A semiconductor sensor that detects or measures a pH change or an optical change caused by an identification substance that specifically recognizes a target marker molecule in a body fluid-derived sample; A device for detecting or measuring a target marker molecule in a sample derived from a body fluid, comprising:

[12] The device according to

[11] , wherein the semiconductor sensor is present in the same structure as the support.

[13] The device according to

[11] or

[12] , wherein the semiconductor sensor is a pH sensor selected from an ISFET, an extended gate FET, and a floating gate FET, or an optical sensor selected from a photodiode, a phototransistor, a CCD, and a CMOS image sensor.

[14] The device according to any one of

[11] to

[13] , wherein the support is selected from the group consisting of a film, glass, elemental silicon, and GaAs.

[15] The device according to any one of

[11] to

[14] , wherein the surface of the support is modified to capture a sample derived from a body fluid.

[16] The device according to

[15] , wherein the modification is carried out by applying a surface modifier selected from an antibody, a cationic polymer, an extracellular matrix protein, a lipid and a heteromolecule having a phosphate group, a phosphonate group, or a silanol group, or a sulfur-containing self-assembled monolayer.

[17] The device according to any one of

[11] to

[16] , wherein the semiconductor sensor electrically detects or measures a pH change or an optical change caused by a substrate-enzyme reaction.

[18] A method for modifying a support for capturing a sample in a body fluid, by applying to the surface of the support a surface modifier selected from an antibody, a cationic polymer, an extracellular matrix protein, a lipid and a heteromolecule having a phosphate group, a phosphonate group, or a silanol group, and a sulfur-containing self-assembled monolayer.

[19] The method described in

[18] for selectively capturing cells or biologically derived vesicles in body fluids.

[20] The method according to

[19] , wherein the biologically derived vesicles are exosomes or extracellular vesicles.

[21] A substance for capturing a body fluid-derived sample on a support, comprising: (1) a portion that binds to the body fluid-derived sample; and (2) a portion that binds to a support material.

[22] The substance according to

[21] , wherein the body fluid-derived sample is a cell or a biologically derived vesicle.

[23] (1) The substance according to

[22] , wherein the site that binds to a sample derived from a body fluid is a hydrocarbon chain or a long-chain fatty acid that binds to the lipid membrane of a cell or biological vesicle.

[24] The substance according to

[23] , which is a hydrocarbon chain or a long-chain fatty acid characterized by having 4 or more carbon atoms.

[25] (2) The substance according to any one of

[21] to

[24] , wherein the site that binds to the support material is composed of a silane compound, a phosphate group or a phosphonate group, a thiol group, or a disulfide group.

[26] The substance is a compound of the following: [ka]

[21] The substance according to

[21] ,

[27] Capturing a single body fluid sample in each of multiple sensing regions of a semiconductor sensor in the device; binding a target marker molecule in a sample to an identification substance that specifically recognizes the target marker molecule; introducing a substrate to induce a pH or optical change near the sensing region; Digitally counting the number of samples derived from body fluids that express a high amount of the target marker molecule based on pH changes or optical changes; Count the number of semiconductor sensors that capture the number of samples derived from body fluids with high expression levels of target marker molecules; A method for analyzing the expression level of a target marker molecule by [Effects of the Invention]

[0012] In the detection and measurement method of the present invention, a body fluid-derived sample carrying multiple target marker molecules on its surface is immobilized on a support, resulting in the capture of multiple target marker molecules on a single support. Each target marker molecule individually undergoes a pH or optical change, and these changes are detected or measured to detect or measure the target marker molecules in the body fluid-derived sample. Furthermore, by capturing each biological vesicle on a single sensor in a one-to-one relationship and integrating these sensors densely on a substrate, the number of biological vesicles expressing high levels of target marker molecules can be counted, allowing the ratio of biological vesicles expressing high levels of target marker molecules to those expressing normal levels to be determined with high accuracy. Using this method or device, the quantitative accuracy and efficiency of liquid biopsies using body fluid-derived samples can be significantly improved. [Brief explanation of the drawings]

[0013] [Figure 1] FIG. 1 shows an example of the configuration of a device for detecting and measuring a target protein contained in a body fluid-derived sample according to the present invention. [Figure 2]FIG. 2 is a diagram showing the outline of the structure of the ion-sensitive field effect transistor (ISFET) used in the experiment. [Figure 3] FIG. 3 shows chemical modification and capture of biological vesicles by molecules with oleyl groups. [Figure 4] FIG. 4 shows the NMR spectrum of the synthesized oleyl group-containing molecule. [Figure 5] FIG. 5 shows the results of mass spectrometry of the synthesized oleyl group-containing molecules. [Figure 6] FIG. 6 shows the sensing area and pH response characteristics before and after capturing BT474 cells. [Figure 7] 7 shows visualization of HER2 expression on the surface of BT474 cells. The bar in the figure indicates 50 μm. [Figure 8] FIG. 8 shows the configuration of a biological vesicle surface biomarker detection and measurement device using an ISFET and urease. [Figure 9] Figure 9 shows the detection and measurement of biological vesicle surface markers based on pH changes. [Figure 10] FIG. 10 shows the response of an ISFET without cell trapping to urea. [Figure 11] FIG. 11 shows the difference between the response of the biological vesicle surface biomarker detection and measurement device obtained in FIG. 9 and the response of the ISFET not capturing cells obtained in FIG. [Figure 12] FIG. 12 shows the configuration of a device in which cells as biological vesicles are bound to an ISFET and biomarkers on the surface of the biological vesicles are detected and measured using glucose oxidase. [Figure 13] FIG. 13 shows that when a biomarker on the surface of a biological vesicle is detected and measured using glucose oxidase, the difference in output potential is a signal that reflects the expression level of the biomarker (EGFR). [Figure 14]FIG. 14 is a diagram showing that the magnitude of the potential change reflects the magnitude of the expression level of the biomarker (EGFR) on the cell membrane surface. [Figure 15] FIG. 15 shows the results of measuring the fluorescence intensity of a fluorescently labeled anti-human EGFR antibody bound to a biomarker (EGFR) on the surface of each cell by FACS. [Figure 16] FIG. 16 shows the results of examining the correlation between the expression level of a biomarker (EGFR) on the cell membrane surface measured using ISFET and the result measured using FACS. [Figure 17] FIG. 17 shows a schematic diagram of anti-CD31 antibody-containing magnetic beads, anti-CD105 antibody-containing primary antibody-secondary antibody-enzyme complex, and the complexation of these with MV. [Figure 18] FIG. 18 shows the relationship between the change in output potential of the ISFET and the glucose concentration when MVs are captured with antibody-containing magnetic beads and a magnet. [Figure 19] FIG. 19 is a schematic diagram showing how molecules having an oleyl group are immobilized on the gate surface of an ISFET and MVs are captured in the vicinity of the gate of the ISFET. [Figure 20] FIG. 20 shows the amount of protein expression on biological vesicles captured by oleyl groups, measured from the relationship between the change in output potential of the ISFET and the glucose concentration. DETAILED DESCRIPTION OF THE INVENTION

[0014] <Liquid biopsy testing> The present invention aims to significantly expand the scope of test targets for liquid biopsy. Until now, test targets for liquid biopsy have been limited to specific soluble proteins, and it has not been possible to simultaneously immobilize multiple types of proteins on a support, making it impossible to detect and measure multiple types of proteins from a single sample.

[0015] On the other hand, the detection and measurement method of the present invention has the advantage that a body fluid-derived sample carries multiple target marker molecules on its surface, and the body fluid-derived sample is immobilized on a support, thereby capturing multiple target marker molecules on a single support. By individually inducing pH or optical changes for each of the multiple target marker molecules captured on the support and detecting or measuring these changes simultaneously or sequentially, multiple target marker molecules carried in a single body fluid-derived sample can be simultaneously or sequentially detected or measured.

[0016] Furthermore, by capturing each biological vesicle on a single sensor and integrating these sensors densely on a substrate, it is possible to count the number of biological vesicles that express a high amount of target marker molecules, and to determine with high precision the ratio of biological vesicles that express a high amount of target marker molecules to those that express a normal amount. The use of this method or device can significantly improve the quantitative accuracy and efficiency of liquid biopsies using bodily fluid-derived samples.

[0017] <Method for detecting or measuring target marker molecules in a body fluid-derived sample> In one embodiment of the present invention, an example of the configuration of a system for detecting and measuring target marker molecules present on the surface of a body fluid-derived sample is shown in Figure 1. In one embodiment of the present invention, this detection and measurement system is used to capture a body fluid-derived sample containing a target marker molecule (e.g., a biomarker protein) on a support including a semiconductor sensor (sensing region), and a recognition substance that specifically recognizes the target marker molecule present on the surface of the body fluid-derived sample is reacted to specifically bind to the target marker molecule. In Figure 1, an antibody is illustrated as an example of the recognition substance. An enzyme is directly or indirectly bound to this recognition substance, and the target marker molecule in the body fluid-derived sample can be detected or measured by detecting or measuring the pH change or optical change caused by the recognition substance-enzyme complex.

[0018] In the present invention, detection refers to qualitatively detecting the presence or absence of a target marker molecule, and measurement refers to quantifying the amount of the target marker molecule present. In the method of the present invention, the pH change or optical change resulting from the recognition substance occurs according to the amount of the target marker molecule, and therefore detection and measurement can be performed according to the sensitivity of the semiconductor sensor.

[0019] The method of the present invention not only enables quantitative measurement of the expression level of target marker molecules on the surface of biological vesicles, but also allows for high-density integration on a substrate of a predetermined size due to the use of semiconductor technology for fabrication. Therefore, by capturing each biological vesicle on a single sensor and setting a predetermined threshold for pH change or optical change, the number of biological vesicles whose expression level of target marker molecules is equal to or greater than a predetermined value can be counted. It is known that actual body fluids contain a variety of biological vesicles, including both those with high and low expression levels of target marker molecules. Using the method of the present invention, the number of biological vesicles with high and low expression levels of target marker molecules can be individually counted and their ratios calculated. The proportion of biological vesicles with high expression levels of target marker molecules among the biological vesicles contained in a body fluid is an important indicator for evaluating the progression of disease and the effectiveness of treatment, and may also serve as an indicator for ultra-early diagnosis of disease.

[0020] The object of detection or measurement in the method and device of the present invention is a target marker molecule present on the surface of a body fluid-derived sample captured on the surface of a support composed of a semiconductor sensor (sensing region). Multiple target marker molecules may be present on the surface of a single body fluid-derived sample, and the target marker molecule to be detected or measured may be one or more of these. When detecting or measuring only a single specific target marker molecule among the multiple target marker molecules present on the surface of a single body fluid-derived sample, an identification substance capable of specifically binding to that target marker molecule may be used. On the other hand, when detecting or measuring multiple target marker molecules among the multiple target marker molecules present on the surface of a single body fluid-derived sample, multiple identification substances capable of specifically binding to each of the target marker molecules may be used.

[0021] In the present invention, an identifying substance can be bound to the surface of a body fluid-derived sample depending on the amount of target marker molecules present on the surface of the body fluid-derived sample. That is, a larger amount of identifying substance can be bound to a body fluid-derived sample surface where the target marker molecules are highly expressed, compared to a body fluid-derived sample where the expression level is low. As a result, a pH change or optical change proportional to the amount of target marker molecules present on the surface of the body fluid-derived sample occurs before and after adding a substance for measuring the amount of identifying substance.

[0022] The sample derived from a body fluid may be any sample derived from a body fluid that can be collected from a living body, such as blood, urine, cerebrospinal fluid, saliva, etc. In the present invention, blood is preferred as the body fluid.

[0023] In the present invention, it is preferable to use cells or biologically derived vesicles as the biologically derived sample, since these samples can contain multiple types of target marker molecules. Here, the biologically derived vesicles refer to exosomes and extracellular vesicles.

[0024] Proteins, glycans, peptides, and other molecules present on the surface of bodily fluid samples can be used as target marker molecules to be detected and measured. These target marker molecules can include marker molecules already known to be associated with diseases or pathological conditions, or marker molecules that change in conjunction with the effects of treatment as a result of disease treatment. For example, proteins known as biomarker proteins for disease or prognosis, such as HER2, EpiCAM, CD9, CD63, CD147, EPS8, cytokeratin, and EGFR, can be selected as target marker molecules present on the surface of bodily fluid samples.

[0025] In the present invention, the most commonly usable identification substance for detecting or measuring the above-mentioned target marker molecule is an antibody when the target marker molecule is a target protein, or an antibody or lectin when the target marker molecule is a sugar chain.

[0026] When an antibody is used as an identifying substance for a target marker molecule, the target marker molecule can be specifically detected or measured by using a primary antibody that specifically binds to the target marker molecule and a secondary antibody that specifically binds to the primary antibody and is conjugated with an enzyme. The primary antibody used in this step specifically binds to a specific target marker molecule, but it must be detected or measured separately from other primary antibodies in subsequent steps. Therefore, each primary antibody used must be derived from a different animal species (e.g., mouse, rat, rabbit, guinea pig, goat, etc.).

[0027] For these primary antibodies, it is necessary to select a secondary antibody according to the animal species of origin and the type of antibody (IgG, IgM, etc.) For example, if the primary antibody is mouse-derived IgG, an anti-mouse IgG antibody should be used as the secondary antibody, and if the primary antibody is goat-derived IgG, an anti-goat IgG antibody should be used as the secondary antibody. This technique is well known in the art.

[0028] In the present invention, the enzyme conjugated to the identification substance for detecting or measuring the target marker molecule can be an enzyme that, when a substrate for the enzyme is introduced, generates or consumes protons (H+) through the substrate-enzyme reaction, causing a change in the pH of the solution, or an enzyme that produces an optical change as a result of the substrate-enzyme reaction. In this method, the target marker molecule can be detected or measured by measuring the pH change or optical change after introducing the substrate, using the pH or light intensity before introducing the substrate as the reference.

[0029] When the enzyme selected is one that produces or consumes protons as a result of its reaction with a substrate, the amount of protons changes near the target marker molecule depending on the expression level of the target marker molecule in the body fluid-derived sample when the substrate is introduced and the enzyme reaction occurs, and the pH changes accordingly. More specifically, when the substrate is introduced and the enzyme reaction occurs, the pH changes depending on the expression level of the target marker molecule on the surface of the body fluid-derived sample, and this pH change is converted into an electrical signal by a semiconductor sensor such as a pH sensor and detected and measured. By detecting and measuring this electrical signal, the expression of the target marker molecule can be analyzed qualitatively and quantitatively.

[0030] Examples of enzymes that can be used to generate or consume protons (H+) as a result of reaction with a substrate include urease, glucose oxidase, penicillinase, glucose dehydrogenase, acetylcholinesterase, and creatinine deiminase. By reacting the above enzymes with their respective substrates, such as urea, glucose, penicillin, glucose, acetylcholine, and creatinine, protons are generated or consumed, changing the pH near the reaction site.

[0031] A pH sensor detects or measures the pH difference between a reference electrode and a measurement electrode and converts it into an electrical signal. Several types of sensors have been developed, including glass electrodes, metal oxide electrodes, fiber optic pH sensors, ion-sensitive field-effect transistors (ISFETs; see Figure 2), extended-gate FETs, and floating-gate FETs. When using an ISFET as a pH sensor, the gate (i.e., sensing area) of the ISFET (see Figure 2) can be optimized to fit the size of the body fluid sample, such as a cell or biological vesicle, to be captured. This allows for the capture of a single body fluid sample on the gate (sensing area) of a single transistor. In this state, the expression of a target protein on the surface of a single biological vesicle can be detected or measured by detecting or measuring the pH change generated by introducing a substrate.

[0032] Alternatively, if the enzyme selected produces an optical change (e.g., luminescence) as a result of its reaction with a substrate, the amount of product that produces the optical change varies depending on the expression level of the target marker molecule on the surface of the body fluid-derived sample, resulting in an optical change near the body fluid-derived sample. This optical change is converted into an electrical signal by a semiconductor sensor such as an optical sensor and then detected and measured. By detecting and measuring this electrical signal, the expression of the target marker molecule can be analyzed qualitatively and quantitatively.

[0033] Examples of enzymes that produce an optical change as a result of reaction with a substrate include alkaline phosphatase (ALP) and horseradish peroxidase (HRP). In the case of ALP, the substrate for the enzyme may be a dioxetane-based chemiluminescent substrate, such as 3-(2'-spiroadamantane)4-methoxy-4-(3'-phosphoryloxy)phenyl-1,2-dioxetane (AMPPD), CSPD (registered trademark), or CDP-Star. TMIn the case of HRP, by reacting with TMB (tetramethylbenzidine) or a luminol-based chemiluminescent substrate, the reaction product can cause an optical change in the vicinity of the reaction site.

[0034] An optical sensor is a sensor that senses light and converts it into an electrical signal. Several types have been developed, including photodiodes, phototransistors, CCDs, and CMOS image sensors, and any of these may be used in the present invention. For example, a phototransistor can capture a single body fluid sample on the gate (sensing area) of a single transistor by optimizing the sensing area to match the size of the body fluid sample, such as a cell or biological vesicle, to be captured. In this state, the expression of a target protein on the surface of a single biological vesicle can be detected or measured by detecting or measuring the optical change generated by introducing a substrate.

[0035] When detecting or measuring multiple target marker molecules, a body fluid-derived sample can be captured in the sensing region of a semiconductor sensor such as a pH sensor or phototransistor, and then the steps of binding an identifying substance and detecting or measuring pH changes or optical changes can be repeated multiple times for a single body fluid-derived sample, or multiple identifying substances can be bound to a single body fluid-derived sample, and then the steps of measuring pH changes or optical changes can be repeated multiple times.

[0036] Among these, when the step of binding an identifying substance and the step of detecting or measuring a pH change or an optical change are repeated multiple times, the enzyme bound to the multiple identifying substances (e.g., antibodies) used in the multiple steps may be of one type.

[0037] For example, when a primary antibody and a secondary antibody-enzyme complex are used as the identification substance, and the step of binding the identification substance and the step of detecting or measuring a pH change or an optical change are repeated multiple times for a single body fluid-derived sample to detect or measure multiple target marker molecules, the following steps are performed for a single body fluid-derived sample captured on a support: performing a step of binding a primary antibody against a first target marker molecule and a step of measuring a pH change or an optical change; performing a step of binding a primary antibody against a second target marker molecule and a step of measuring a pH change or an optical change; This process is repeated multiple times with a third target marker molecule, a fourth target marker molecule, and so on. It may also be carried out by

[0038] Alternatively, in a case where a plurality of identifying substances are bound to a single body fluid-derived sample and then the process of measuring pH changes or optical changes is repeated multiple times, the enzymes bound to the identifying substances are different for each of the different identifying substances.

[0039] For example, when a primary antibody and a secondary antibody-enzyme complex are used as identification substances to detect or measure multiple target marker molecules, if the above multiple types of identification substances are to be specifically and simultaneously bound to multiple target marker molecules on the surface of a single body fluid-derived sample, after binding multiple primary antibodies, measuring a pH change or an optical change for the first target marker molecule; measuring a pH change or an optical change for the second target marker molecule; Repeating multiple times with a third target marker molecule, a fourth target marker molecule, and so on. It may also be carried out by

[0040] In either method, the expression of a first target marker molecule that specifically binds to the first recognition substance can be qualitatively and quantitatively examined by introducing a substrate for a first enzyme bound to a first recognition substance, causing an enzymatic reaction, and detecting or measuring a pH change or optical change. Furthermore, the expression of a second target marker molecule that specifically binds to a second primary antibody can be qualitatively and quantitatively examined by introducing a substrate for a second enzyme bound directly or indirectly to a second recognition substance, causing an enzymatic reaction, and detecting or measuring a pH change or optical change. In this way, by sequentially introducing enzyme substrates bound to primary antibodies into the reaction field, causing an enzymatic reaction, and measuring the pH change, multiple target marker molecules present on the surface of a single body fluid-derived sample can be sequentially analyzed.

[0041] <Device for detecting or measuring target marker molecules in a body fluid-derived sample> In another aspect, the present invention provides a detection and measurement device for realizing a detection and measurement system for target marker molecules present on the surface of a body fluid-derived sample, as shown in Figure 1. This device is provided as comprising a support for capturing the body fluid-derived sample, and a semiconductor sensor for measuring a pH change or optical change that occurs as a result of an identification substance that specifically recognizes the target marker molecule in the body fluid-derived sample.

[0042] In this device, an enzyme is directly or indirectly bound to a target marker molecule present on the surface of a body fluid-derived sample captured on the support described above, and the pH or optical change that occurs when a substrate corresponding to the enzyme is added is converted into an electrical signal, which can be detected and measured, thereby detecting or measuring the target marker molecule in the body fluid-derived sample.

[0043] The semiconductor sensors held in this device can be pH sensors such as glass electrodes, metal oxide electrodes, optical fiber pH sensors, ion-sensitive field effect transistors (ISFETs, see Figure 2), extended gate FETs, and floating gate FETs for measuring pH changes, or optical sensors such as photodiodes, phototransistors, CCDs, and CMOS image sensors for measuring optical changes.

[0044] In the system of the present invention, the pH change or optical change caused by the target marker molecules present on the surface of a body fluid sample is weak. Therefore, in order to increase the efficiency of detecting and measuring the pH change or optical change, it is desirable that the support that captures the body fluid sample and the semiconductor sensor that detects and measures the pH change or optical change are as close as possible to each other. In the closest possible proximity, the surface of the semiconductor sensor can be used as the semiconductor sensor. In this case, the support and the semiconductor sensor are present in the same structure. In other words, when the support and the semiconductor sensor are present in the same structure, the support is the same as the semiconductor sensor, and the support is selected from a membrane, glass, ISFET, extended-gate FET, floating-gate FET, and phototransistor.

[0045] The support may be made of any material that can capture a body fluid-derived sample, and metal oxides such as silicon oxide (glass), iridium oxide, or tantalum oxide, or nitrides such as silicon nitride can be used.

[0046] <Sensor integration> The semiconductor sensors used in the present invention can be integrated by miniaturizing each unit of the semiconductor sensor. pH sensors and optical sensors, which are examples of the semiconductor sensors used in the present invention, have both been realized as small sensors on the order of millimeters or micrometers, and these small semiconductor sensors can be functionally arranged on a support and integrated for measurement.

[0047] For example, the ion-sensitive field effect transistor (ISFET) exemplified as a pH sensor is a micrometer-order sensor fabricated using semiconductor microfabrication technology, allowing for high-density integration and arraying. The configuration of the present invention can be constructed on individual transistors of a high-density arrayed ISFET. Therefore, a device including a highly integrated semiconductor sensor can be provided for detecting or measuring target marker molecules in body fluid samples.

[0048] The size (area) of the sensing region of the semiconductor sensor is preferably in the range of 100 nm to 50 μm in terms of the long side, and can be optimized depending on the measurement target. For example, in the case of detecting cells, the long side is preferably 5 to 50 μm, and the size is preferably adjusted to an optimal size depending on the type of cell. In addition, in the case of detecting exosomes or extracellular vesicles, the long side of the sensing region is preferably in the range of 100 to 500 nm.

[0049] In this way, by capturing a single body fluid sample on each of the densely integrated individual semiconductor sensors (pH sensors or light sensors), introducing a substrate, and measuring the pH or light intensity, it is possible to digitally count the number of body fluid samples that express a high amount of target marker molecules, and quantitatively analyze the expression levels of the target marker molecules. Specifically: capturing a single body fluid-derived sample on each of a plurality of sensing regions of a semiconductor sensor in the device; binding a target marker molecule in a sample to an identification substance that specifically recognizes the target marker molecule; introducing a substrate to induce a pH or optical change near the sensing region; Digitally count the number of samples derived from body fluids that express a high amount of target marker molecules based on pH changes or optical changes; Count the number of semiconductor sensors that capture the number of samples derived from body fluids with high expression levels of target marker molecules; This can provide a method for analyzing the expression level of a target marker molecule.

[0050] As described above, by capturing each biological vesicle on a single sensor and setting a predetermined threshold for pH change or optical change, it is possible to count the number of biological vesicles whose expression level of a target marker molecule is equal to or greater than a predetermined value. It is known that actual body fluids contain a variety of biological vesicles, and that biological vesicles with high and low expression levels of target marker molecules are mixed. Using the method of the present invention, the number of biological vesicles with high and low expression levels of target marker molecules can be counted individually and their ratios calculated. The proportion of biological vesicles with high expression levels of target marker molecules among the biological vesicles contained in body fluids is an important indicator for evaluating the progression of disease and the effectiveness of treatment, and is also thought to be an indicator for ultra-early diagnosis of disease.

[0051] <Modification of the support surface> The support surface may be modified in various ways to capture biological fluid samples such as cells and biological vesicles. For example, modifications can be made by coating the support surface with antibodies, cationic polymers such as poly-L-lysine, extracellular matrix proteins such as collagen, fibronectin, and vitronectin, surface modifiers selected from heteromolecules containing lipids such as oleyl groups and phosphate, phosphonate, or silanol groups, or sulfur-containing self-assembled monolayers. To introduce these modifications to the support surface, the support surface may be pretreated with a silane coupling agent and then various modifications may be introduced.

[0052] As mentioned above, one example of a surface modifier for capturing a body fluid-derived sample on a support is a heteromolecule containing a lipid such as an oleyl group and a phosphate group, phosphonate group, or silanol group. This specific configuration can be used to capture a body fluid-derived sample on a support, consisting of (1) a site that binds to a body fluid-derived sample, such as a biological vesicle, and (2) a site that binds to the support material (Figure 3). In the above heteromolecule, the site that binds to the body fluid-derived sample (1) binds to the lipid membrane of cells or biological vesicles, and is composed of a hydrocarbon chain or a long-chain fatty acid. The hydrocarbon chain or long-chain fatty acid used in this embodiment preferably has four or more carbon atoms, and for example, a lipid-based oleyl group can be used. Meanwhile, the site that binds to the support material (2) can bind to the surface of a support made of silicon oxide (glass), silicon nitride, or a metal oxide such as tantalum oxide. This site that binds to the support material (2) consists of an organosilane compound, a phosphate group, a phosphonate group, a thiol group, or a disulfide group. By forming a molecule having the above structures (1) and (2) on a support, the oleyl group fuses with the lipid membrane of the body fluid-derived sample (e.g., biological vesicles), allowing the body fluid-derived sample to be captured on the support. Furthermore, by forming this molecule on the sensing region of a semiconductor sensor, the body fluid-derived sample can be captured on the sensing region. An example of a heteromolecule in the present invention is a molecule having the following formula:

[0053] [ka]

[0054] can be used.

[0055] Below, we will show an example of detection and measurement using an enzyme that generates protons (H+) and changes the pH depending on the expression level of target marker molecules on the surface of a body fluid-derived sample, and a pH sensor as the semiconductor sensor. The principle is the same when using an enzyme that generates an optical change, but an optical sensor is used as the semiconductor sensor, and detection is possible in the same way. [Example]

[0056] Example 1: Surface modification of the sensing region In this example, the configuration of a device in which a pH sensor is used as the semiconductor sensor was examined.

[0057] In the experiment, the pH sensor used was the ISFET (ISFET.COM) shown in Figure 2. This pH sensor uses tantalum pentoxide (Ta2O5) with a thickness of 40 nm as the pH-sensitive membrane material, with a gate length of 10 μm and a gate width of 300 μm, which forms the sensing region for hydrogen ions (protons).

[0058] A glass tube with an inner diameter of 5 mm was attached to the surface of the ISFET so as to include this gate region, using a thermosetting epoxy resin that was cured at 120°C for 2 hours, and cells were captured inside the tube as a body fluid sample to form a reaction field.

[0059] To capture biological fluid samples in the sensing region (the gate of the ISFET), the Ta2O5 surface was chemically modified. The Ta2O5 surface was first washed with a mixture of 60 μl of NH4OH (10% V / V) and 140 μl of HO2 (20% V / V), then rinsed with water. It was then immersed in a poly-L-lysine solution (0.01% w / v) at room temperature for 10 minutes to adsorb the poly-L-lysine onto the Ta2O5 surface. The surface was then washed with pure water and dried. Because the poly-L-lysine has a positive charge due to the amino groups, while the surface of biological vesicles has a negative charge, biological vesicles can be captured on the poly-L-lysine surface through electrostatic interaction. The device thus prepared was used in the following examples.

[0060] Example 2: Synthesis of oleyl group-containing molecules In this example, an oleyl group-containing molecule was synthesized to be used in an embodiment in which biological vesicles are captured on the surface of a sensor substrate (FIG. 3).

[0061] 20 mg of 10-Carboxydecylphosphonic acid (10CDPA) (Dojindo) was added to 1 mL of tetrahydrofuran (THF), and then 35 mg of (4,6-dimethoxy-1,3,5-triazin-2-yl)-4-methylmorpholinium chloride (DMT-MM) (Wako) and 6 μl of oleylamine (Wako) were added, mixed, and allowed to stand at room temperature for 2 hours to react.

[0062] Excess DMT-MM and excess oleylamine were removed from this mixture, and the resulting solid aggregates were dissolved in 1,4-dioxane solvent and freeze-dried under liquid nitrogen for 3-4 hours to produce powder.

[0063] The NMR spectrum of the oleyl group-containing molecule synthesized by the above process is shown in Figure 4. This confirms that the amide bond reaction has progressed. The mass spectrometry results of the synthesized oleyl group-containing molecule are shown in Figure 5. The chemical formula of the target compound is C 29 H 58 The molecular weight of the compound synthesized was 516.45, while the molecular weight of NO4P was 515.4, and it was confirmed that the target compound represented by the following chemical formula had been synthesized.

[0064] [ka]

[0065] Example 3: Cell culture In this example, cells and their culture were examined as examples of samples derived from body fluids.

[0066] The human breast cancer cell lines BT474 (ATCC No. HTB-20), MCF10A (ATCC No. CRL-10317), and MDA-MB-231 (ATCC No. HTB-26) (American Type Culture Collection (ATCC), Manassas, VA)) were used as body fluid samples.

[0067] The culture medium for MCF10A cells was MEGM TM Mammary Epithelial Cell Growth Medium (Lonza, Walkersville, MD, USA) was used, and the culture medium for other cells was RPMI1640 1X (Thermo Fisher Scientific, MA, USA) containing 10% Fetal Bovine Albumin (FBS), penicillin-Streptomycin 1X (Wako, Japan), and glutamax 1X (Thermo Fisher Scientific, MA, USA).

[0068] All of the above cells were first cultured on cell culture dishes at 37°C in a 5% CO2 atmosphere. 5 BT474 cells at 1000 cells / ml were trypsinized and seeded onto the sensing region of an ISFET (ISFET.COM, Inc.) (discussed in Example 1) chemically modified with poly-L-lysine or oleyl group-containing molecules, and cultured overnight at 37°C in a 5% CO2 atmosphere.

[0069] The pH response characteristics of the ISFET were evaluated with cells trapped in the sensing region of the ISFET. An Ag / AgCl electrode with a salt bridge was placed in the glass tube formed in the sensing region of the ISFET, and the potential response of the ISFET was measured when buffer solutions of pH 4, 7, and 9 were introduced. The results are shown in Figure 6. Even when the Ta2O5 surface of the sensing region was chemically modified with poly-L-lysine or oleyl group-containing molecules and cells were then trapped and cultured on the modified surface, the pH response characteristics of the ISFET remained almost unchanged, and it was found that good pH response characteristics such as sensitivity and stability were maintained.

[0070] Example 4: Confirmation of HER2 expression by confocal fluorescence microscopy In this example, the expression state of HER2, which is used as one of the markers for breast cancer detection on the cell membrane surface, on BT474 cells was confirmed by a conventional method.

[0071] BT474 cells were trypsinized, seeded on glass dishes, and cultured overnight at 37°C in a 5% CO2 atmosphere. A 4% paraformaldehyde solution in phosphate buffer was added dropwise and incubated for 20 minutes, followed by three washes with phosphate buffer. To prevent nonspecific binding, a blocking solution consisting of 10% v / v normal goat serum (Vector Laboratory), 0.2% v / v Tween 20 (Nacalai Tesque; 28353-85), and 0.7% v / v glycerol (Wako, Japan) was applied to the BT474 cell surface and incubated for 1 hour. The cells were then incubated with a 1:200 diluted rabbit anti-HER2 polyclonal antibody (SAB4500789; Sigma) for 1 hour. After washing with phosphate buffer, a 1:400 diluted FITC-conjugated anti-rabbit antibody (Abcam, Cambridge, UK) was added and incubated for 1 hour.

[0072] For fluorescent staining, the cells were incubated with 4,6-diamidino-2-phenylindole (DAPI; FUJIFILM Wako Pure Chemical Corporation) diluted 1 / 1000 in phosphate buffer for 10 minutes. After washing, HER2 expression on the cell membrane surface was observed using a Nikon eclipse Ti-E confocal microscope with an LU-N series laser unit (Nikon, Tokyo, Japan).

[0073] The results are shown in Figure 7 (the bar in the figure indicates 50 µm), which confirmed that HER2 was expressed on the membrane surface of BT474 cells.

[0074] Example 5: Detection of HER2 using the device for detecting target marker molecules on the surface of a body fluid-derived sample according to the present invention In this example, it was confirmed whether HER2 could be detected using the device of the present invention for BT474 cells, which were confirmed to express HER2 on the cell surface.

[0075] The surface of the sensing region of the ISFET was chemically modified with poly-L-lysine in the same manner as in Example 2. 5 BT474 cells were captured at 1000 cells / ml and incubated overnight at 37°C in a 5% CO2 atmosphere. A 4% paraformaldehyde solution in phosphate buffer was added dropwise and incubated for 20 minutes, followed by three washes with phosphate buffer. The blocking solution was applied to the BT474 cell surface for 1 hour, followed by washing. The cells were then incubated with a 1 / 200 diluted rabbit anti-HER2 polyclonal antibody (SAB4500789; Sigma) for 1 hour. After washing, the cells were incubated with a 1 / 1600 diluted urease-conjugated anti-rabbit antibody for 90 minutes to allow binding to the anti-HER2 polyclonal antibody. The cells were then washed to remove any unbound anti-rabbit antibody.

[0076] The configuration of the target marker molecule detection device for the surface of a body fluid-derived sample constructed by the above process is shown in Figure 8. Although the basic concept of the present invention is similar to that shown in Figure 1, in this example an ISFET, which is a pH sensor, is used as the semiconductor sensor, a primary antibody that recognizes HER2 and a secondary antibody that recognizes the primary antibody are used as identification substances, and urease is used as the enzyme that binds to the secondary antibody, to detect and measure the presence of HER2 on the cell surface by changes in pH.

[0077] Urea was introduced into the sensing region of the target marker molecule detection device for the body fluid sample surface, and the reaction between urease and urea was allowed to occur on the cell surface. The resulting proton (H+) generated by the reaction led to a pH change monitored by an ISFET. The reference electrode was an Ag / AgCl electrode with a 100 mM potassium chloride solution as the internal solution, and the sample solution was contacted via a salt bridge made of agarose gel.

[0078] Figure 9 shows the response of three identical devices for detecting target marker molecules on the surface of body fluid-derived samples when a 15 mM urea solution was introduced into a glass tube. The potential response is shown converted into pH change. After the urea solution was introduced, the response of the three devices gradually shifted toward alkaline and reached a steady state in about 15 minutes. The pH changed by 0.6, from 6.3 to 6.9, over 15 minutes. The decomposition reaction of urea by urease is shown in equation (1). (NH2)2CO + 2H2O + H + → 2NH4 + + HCO3 - (1)

[0079] From this equation (1), it can be seen that hydrogen ions (H+) are consumed as the enzymatic reaction of urease progresses. In other words, it can be seen that the pH shifts to the alkaline side. Therefore, it can be understood that the change to the alkaline side observed in Figure 9 is a change that occurs as a result of the reaction between urease and urea on the cell surface. Since the amount of urease bound to the cell surface reflects the expression level of HER2, this means that the characteristics of the obtained pH change reflect the expression level of HER2 on the cell surface. Therefore, it is possible to determine the expression level of HER2 by measuring the reaction rate of the pH change or the amount of pH change.

[0080] As a negative control, a 15 mM urea solution was introduced into the sensing region of an ISFET without BT474 cells, as described above. Figure 10 shows the response. After the introduction of urea, the pH instantly changed by 0.2, from 6.3 to 6.5, and then remained constant. This change was smaller than the pH change observed in the presence of cells, as shown in Figure 9. Since the pH of the measurement solution and the 15 mM urea solution introduced later were different, this was expected to be a pH change simply caused by mixing.

[0081] Figure 11 shows the difference between the response of the target marker molecule detection device on the surface of a body fluid-derived sample (Figure 9) and the response of the ISFET (negative control) without cell capture (Figure 10) to a 15 mM urea solution. This differential measurement allows us to subtract background effects such as changes in the pH of the entire mixture due to differences in the pH of the two solutions and changes in the ISFET device response due to differences in temperature. As a result, we can extract only the pH change that reflects the expression level of HER2 on the cell surface due to the enzyme-substrate reaction. Based on the expression level of HER2 on the BT474 cell membrane surface, we found that the reaction between urease and urea caused a pH change of approximately 0.35 over 15 minutes. This demonstrates that we successfully monitored the pH increase resulting from the urease enzymatic reaction, targeting HER2 expressed on the BT474 cell surface.

[0082] Thus, it was demonstrated that the target marker molecule, HER2, expressed on the surface of a body fluid-derived sample, BT474, can be monitored by the increase in pH resulting from the urease enzyme reaction, and the expression level of the target marker molecule can be detected and measured using the target marker molecule detection device for the body fluid-derived sample of the present invention. Although cells were used as the body fluid-derived sample in this example, it was shown that this detection principle can also be applied to other body fluid-derived samples (e.g., exosomes and extracellular vesicles).

[0083] Example 6: Quantitative detection of cell surface protein expression levels In this example, an example of quantitatively detecting differences in the expression level of epidermal growth factor receptor (EGFR) protein on the surface of breast cancer cells was examined.

[0084] (6-1) Experimental protocol An extracellular matrix (ECM) gel layer for capturing biological vesicles was formed on the gate insulating film surface of a pH-sensitive field-effect transistor, and one of four established cell lines, BT474, MM231, MM468, and MM453, was individually seeded on top of it (see Figure 12).

[0085] Specifically, BT474, MM231, MM468, and MM453 were cultured in advance, and then detached from the culture dish by trypsin treatment. The cells were then collected by centrifugation at 3000 rpm for 5 minutes, and 1 × 10 6 A cell suspension was prepared at a concentration of 1000 cells / mL.

[0086] 200 μL of cell suspension (0.2 × 10 6 Cells (containing 100 cells) were seeded onto the reaction chamber (glass ring) on ​​the ISFET surface where ECM gel had formed and cultured overnight at 37°C in a 5% CO2 atmosphere. A 2% bovine serum albumin (BSA) solution prepared in bis-tris propane (BTP) buffer was then added to the reaction chamber as a blocking solution. After incubation at 37°C in a 5% CO2 atmosphere for 1 hour, the cells were washed three times with BTP buffer. This process allowed the BSA to adsorb to the cell surface, ECM gel surface, and reaction chamber surface, suppressing nonspecific adsorption.

[0087] To form an antibody-enzyme complex corresponding to the expression level of EGFR on the cell surface, an anti-human EGFR antibody (BioLegend Inc.) was first added as the primary antibody and incubated at 37°C for 30 minutes. Next, 50 μL of a secondary antibody complexed with the enzyme glucose oxidase was added and incubated at 37°C for 30 minutes. Finally, the supernatant was discarded, and the cells were washed three times with BTP buffer.

[0088] (6-2) Confirmation of the correlation between the generated potential and the expression level of the cell membrane surface protein (EGFR) A schematic diagram of the cell-immobilized ISFET prepared as described above is shown in Figure 12. As an example, 50 μL of BTP buffer was added to the reaction chamber of an ISFET with MM468 cells immobilized, and the output voltage of the ISFET was measured. Once the output voltage stabilized, 5 mM, 10 mM, 25 mM, and 50 mM glucose solutions were added. The change in output potential of the cell-immobilized ISFET at this time is shown in Figure 13. The response of an ISFET without immobilized cells (■) is also shown as a control.

[0089] The output potential (●) of the cell-immobilized ISFET increases depending on the glucose concentration. This increase in output potential indicates that the pH near the gate has shifted toward acidity. This is thought to be due to the enzyme glucose oxidase binding to EGFR expressed on the cell surface via the primary and secondary antibodies, which decomposes glucose in an enzymatic reaction depending on the added glucose concentration to produce gluconic acid, causing a local change in pH near the gate (sensing part) of the ISFET.

[0090] The reason why the change in output potential tends to saturate when the glucose concentration becomes high is thought to be that dissolved oxygen is consumed during the enzyme reaction, causing a temporary shortage of dissolved oxygen near the cell surface in the high concentration region, making the enzyme reaction rate-limited by the dissolved oxygen concentration.

[0091] The response of the ISFET without immobilized cells (■) also showed a small change in potential, but a glucose concentration-dependent response. This is thought to be because the primary and secondary antibodies were added without cells on the ISFET surface, resulting in nonspecific adsorption of the secondary antibody / enzyme complex to the ECM gel surface, etc., and subsequent enzymatic decomposition by the addition of glucose, resulting in a pH change. The difference in output potential between the cell-immobilized ISFET (●) and the ISFET without immobilized cells (■) is thought to be a signal reflecting the expression level of EGFR. Based on this result, in the following Example (6-3), the output potential from the secondary antibody / enzyme (glucose oxidase) complex was measured at a glucose concentration of 10 mM.

[0092] (6-3) Correlation between the magnitude of potential change and the expression level of cell membrane surface protein (EGFR) Next, we demonstrated that the expression levels of EGFR on the cell membrane surface of four cell lines, BT474, MM231, MM468, and MM453, differed.

[0093] BT474, MM231, and MM453 cells were captured on the gate surface of an ISFET. The enzyme glucose oxidase was immobilized on the cell membrane surface via primary and secondary antibodies using the same protocol as for capturing MM468 cells in (6-2) above. The response (output potential) was measured when a 10 mM glucose solution was added. Figure 14 shows a comparison of the ISFET responses when capturing four types of cells (MM453, BT474, MM231, and MM468). The magnitude of the potential change increased in the order MM453, BT474, MM231, and MM468, which is thought to reflect the level of EGFR expression on the cell membrane surface.

[0094] To confirm the correlation between the magnitude of the potential change and the amount of EGFR expression on the cell membrane surface, the expression levels of EGFR on the surface of the four types of cells were analyzed using Fluorescence-Activated Cell Sorting (FACS). A fluorescently labeled anti-human EGFR antibody was used to bind and stain the EGFR on the cell surface, and the fluorescence intensities were compared. The results of measuring the fluorescence intensity of MM453, BT474, MM231, and MM468 cells are shown in the figure. The magnitude of the fluorescence intensity, i.e., the amount of EGFR expression, increased in the order of MM453, BT474, MM231, and MM468, from lowest to highest, indicating a consistent trend with ISFET measurements.

[0095] The expression levels of EGFR on the cell membrane surface of four types of cell lines, MM453, BT474, MM231, and MM468, were measured using ISFET on the vertical axis, and the results measured using FACS on the horizontal axis. The correlation between the two is shown in Figure 16. This shows that there is a good correlation between the two, demonstrating that the cell-immobilized ISFET of the present invention can be used to quantitatively analyze the expression levels of EGFR on the cell membrane surface.

[0096] This method can be applied to cell surface markers other than EGFR by changing the antibody. Furthermore, by using it in combination with enzymes other than glucose oxidase, such as the urease already described, it is possible to continuously analyze multiple cell surface markers while the cells are captured on the ISFET surface.

[0097] Example 7: Measurement of protein expression levels on bead-captured biological vesicles This example was carried out with the aim of demonstrating that the configuration of the present invention can be applied not only to the cells shown in Example 6 but also to other biological vesicles.

[0098] Human umbilical vein endothelial cells (HUVEC)-derived MVs were bound to magnetic beads, and a magnet was placed on the backside of the ISFET to localize the MVs on the gate surface of the ISFET. The output potential of the ISFET was then measured by adding glucose.

[0099] Specifically, 100 μL of a 30 μg / mL goat anti-mouse IgG secondary antibody conjugated with the enzyme glucose oxidase (Glucose Oxidase Conjugation Kit - Lightning-Link, Goat anti-Mouse IgG H&L; Abcam) and a 20 μg / mL primary anti-human CD105 antibody (mouse IgG) (Purified anti-human CD105; BioLegend) were added to a 1.5 mL tube, mixed, gently tapped, and then allowed to stand in the refrigerator for 30 minutes. CD105 is expressed on the MV surface as one of several surface molecules, to which the primary antibody-secondary antibody-enzyme complex binds.

[0100] 4×10 5 Washed CD31 antibody-immobilized beads (Dynabeads) at 1000 beads / mL TMOne microliter of CD31 Endothelial Cell (Invitrogen) was added and the beads were incubated at 4°C for 30 minutes. The magnetic beads had CD31 antibodies immobilized on them, which allowed MVs to bind to the surface of the beads. CD31 is the abbreviation for PECAM-1 (platelet endothelial cell adhesion molecule), and is known as a marker for HUVECs.

[0101] Then, a magnet (DynaMag TM The beads were then incubated in a 100 μL PBS-2 (Invitrogen) for 2 minutes, collected, and the supernatant discarded. The beads were then removed from the magnet, washed with 100 μL of 1 mM phosphate buffer, pipetted several times, and incubated again for 2 minutes. The supernatant was discarded. This washing process was repeated three times. Figure 17 shows a schematic diagram of the anti-CD31 antibody-containing magnetic beads, the anti-CD105 antibody-containing primary antibody-secondary antibody-enzyme complex, and their complexation with MV.

[0102] Next, a neodymium magnet (1 x 1 x 2 mm) was placed behind the gate of the ISFET with the chamber attached. The MV solution prepared as described above was added dropwise and left to stand for approximately 20 minutes. ISFET measurement was started, and glucose solution (0 mM, 5 mM, 10 mM, 25 mM) or phosphate buffer solution was added, and the potential response was recorded.

[0103] Figure 18 shows the relationship between the change in ISFET output potential after 40 minutes and the glucose concentration. The output potential increased as the glucose concentration increased, confirming that this was due to the reaction of the enzyme bound to CD105. Furthermore, compared to Figure 16, when the glucose concentration was 25 mM, the output potential change in the liquid phase reaction was 10 mV, whereas the output potential change increased to 26 mV when beads were used to collect MVs at the ISFET gate before the enzyme reaction.

[0104] Example 8: Measurement of protein expression levels on biological vesicles captured with oleyl groups This example was carried out to examine a method for efficiently capturing the MVs shown in Example 7 in the vicinity of the gate of the ISFET.

[0105] Specifically, the oleyl group-containing molecules prepared in Example 2 were immobilized on the gate surface of the ISFET, and MVs were captured even closer to the gate of the ISFET (Figure 19). The oleyl group has a high affinity for the lipid membrane of MVs, allowing them to directly fuse with the lipid membrane and capture MVs on the gate surface. The oleyl group was immobilized by a coupling reaction between the hydroxyl group of the gate oxide film and the phosphonic acid group of the oleyl derivative. 20 μL of 1 mM oleyl derivative dissolved in isopropyl alcohol (IPA) was dropped onto an ISFET equipped with a chamber, air-dried for 1 hour, and then completely dried with nitrogen gas. The ISFET was annealed in an oven at 120 °C for 1 hour and rinsed three times with 200 μL of IPA.

[0106] The MV solution, a goat anti-mouse IgG secondary antibody conjugated with the enzyme glucose oxidase (Glucose Oxidase Conjugation Kit - Lightning-Link, Goat anti-Mouse IgG H&L; Abcam), and a mouse anti-human CD105 primary antibody (purified anti-human CD105; BioLegend) were mixed and left to stand in a refrigerator for 30 minutes. After centrifugation at 20,000 g and 4°C for 30 minutes, the supernatant was discarded and the precipitate was dissolved in phosphate buffer.

[0107] 100 μL of MV solution was placed in the chamber of the ISFET with the immobilized oleyl derivative and allowed to stand at room temperature for 1 hour to immobilize MV on the ISFET. The supernatant was discarded and replaced with phosphate buffer. Sequentially, phosphate buffer was added to the ISFET to initiate potential measurement, and glucose solutions (5 mM, 10 mM, 25 mM) or phosphate buffer were added, and the potential response was recorded.

[0108] Figure 20 shows the relationship between the change in ISFET output potential after 40 minutes and the glucose concentration. The output potential increased as the glucose concentration increased, confirming that this was due to a reaction caused by the enzyme bound to CD105. The data marked with ■ in the figure represent the results of first applying MV to an ISFET with an oleyl derivative immobilized on it to capture the MV on the gate, followed by the addition of an enzyme-conjugated antibody. The data marked with ● represent the results of first reacting MV with an antibody and an enzyme-conjugated antibody in the liquid phase to form an antibody-enzyme complex on the surface of the MV, followed by capture by the oleyl group on the gate surface.

[0109] Both data showed that the output potential changed depending on the glucose concentration, indicating that CD105 on the MV surface could be detected. The difference in the magnitude of the output potential was thought to be due to differences in nonspecific adsorption caused by differences in the detection protocol. [Industrial Applicability]

[0110] In the detection and measurement method of the present invention, a body fluid-derived sample carrying multiple target marker molecules on its surface is immobilized on a support, resulting in the capture of multiple target marker molecules on a single support. Each target marker molecule individually undergoes a pH or optical change, and these changes are detected or measured to detect or measure the target marker molecules in the body fluid-derived sample. Furthermore, by capturing each biological vesicle on a single sensor in a one-to-one relationship and integrating these sensors densely on a substrate, the number of biological vesicles expressing high levels of target marker molecules can be counted, allowing the ratio of biological vesicles expressing high levels of target marker molecules to those expressing normal levels to be determined with high accuracy. Using this method or device, the quantitative accuracy and efficiency of liquid biopsies using body fluid-derived samples can be significantly improved.

Claims

1. a step of capturing a cell sample or a biological vesicle sample containing a target marker molecule without using an antibody onto a support whose surface has been modified by applying a surface modifier selected from a cationic polymer, an extracellular matrix protein, a lipid and a heteromolecule having a phosphate group, a phosphonate group or a silanol group; binding the captured sample with an identification substance that specifically recognizes the target marker molecule in the sample; detecting or measuring the pH change caused by the identifying substance bound to the target marker molecule by a semiconductor sensor present in the same structure as the support; A method for detecting or measuring a target marker molecule in a cell sample or a biological vesicle sample based on a pH change, comprising:

2. 2. The method of claim 1, wherein the cell sample or biological vesicle sample is a single cell sample or biological vesicle sample and comprises multiple target marker molecules therein.

3. The method according to claim 2, wherein the step of binding one type of identification substance to a single cell sample or biological vesicle sample and the step of detecting or measuring the resulting pH change are repeated multiple times.

4. 4. The method according to claim 3, wherein one type of enzyme is bound to a plurality of recognition substances used in a plurality of steps.

5. The method according to claim 2, wherein the step of binding multiple identification substances to a single cell sample or biological vesicle sample and then detecting or measuring the pH change caused by each of the identification substances is repeated multiple times.

6. 6. The method according to claim 5, wherein the enzymes bound to the identification substances are different enzymes for each of a plurality of different identification substances.

7. The method according to any one of claims 1 to 6, wherein the identifying substance is an antibody that specifically binds to the target marker molecule.

8. The method according to any one of claims 1 to 7, wherein the biological vesicles are exosomes or extracellular vesicles.

9. The method according to any one of claims 1 to 8, wherein the pH change is caused by a substrate-enzyme reaction.

10. a support for capturing a cell sample or a biologically derived vesicle sample, the surface of which is modified by applying a surface modifier selected from cationic polymers, extracellular matrix proteins, lipids and heteromolecules having a phosphate group, a phosphonate group or a silanol group; a semiconductor sensor that detects or measures pH changes caused by an identification substance that specifically recognizes a target marker molecule in a cell sample or a biological vesicle sample, the identification substance being present in the same structure as the support; A device for detecting or measuring a target marker molecule in a cell sample or a biologically-derived vesicle sample, comprising:

11. 11. The device of claim 10, wherein the semiconductor sensor is a pH sensor selected from an ISFET, an extended gate FET, and a floating gate FET.

12. 12. The device according to claim 10 or 11, wherein the support is selected from the group consisting of a pH-sensitive film selected from the group consisting of tantalum oxide, iridium oxide, silicon nitride, and hafnium oxide, glass, silicon element, a carbon material selected from the group consisting of carbon nanotubes and graphite, a two-dimensional material selected from the group consisting of graphene and molybdenum sulfide, and GaAs.

13. The device of any one of claims 10 to 12, wherein the semiconductor sensor electrically detects or measures pH changes caused by a substrate-enzyme reaction.

14. (1) a site that binds to a cell sample or a biologically derived vesicle sample; and (2) a moiety consisting of a phosphate or phosphonate group that binds to the surface of a metal oxide support material; A substance for capturing a cell sample or a biological vesicle sample on a support, comprising:

15. (1) The substance according to claim 14, wherein the binding moiety to a cell sample or a biological vesicle sample is a hydrocarbon chain or a long-chain fatty acid that binds to the lipid membrane of the cell or biological vesicle.

16. 16. The substance according to claim 15, wherein the hydrocarbon chain or long-chain fatty acid is a hydrocarbon chain or long-chain fatty acid characterized by having 4 or more carbon atoms.

17. The substance is the following compound 【Chemistry 1】 17. The substance according to any one of claims 14 to 16, wherein

18. a semiconductor sensor in the device, the surface of which has been modified by applying a surface modifier selected from a cationic polymer, an extracellular matrix protein, a lipid, and a heteromolecule having a phosphate group, a phosphonate group, or a silanol group, and capturing a single cell sample or a biological vesicle sample on each of a plurality of sensing regions of the semiconductor sensor in the device without using an antibody; binding a target marker molecule in a sample to an identification substance that specifically recognizes the target marker molecule; introducing a substrate to induce a pH change near the sensing region; digitally counting the number of cells or biological vesicles in the sample that express a high amount of the target marker molecule based on the pH change; Counting the number of semiconductor sensors on which the cell sample or biological vesicle sample expressing a high amount of the target marker molecule is captured; A method for analyzing the expression level of a target marker molecule by

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