JP4958941B2 - Analytical element - Google Patents

Description

  The present invention relates to an analytical element capable of measuring biological substances such as proteins with high sensitivity.

  The enzyme immunoassay, which is a typical protein measurement method, utilizes a reaction in which an antibody selectively binds to a specific protein, that is, an antigen-antibody reaction. The measurement method is roughly divided into a sandwich method and a sandwich method. There is a competitive reaction method. In the sandwich method, the amount of antigen is indirectly measured through an enzyme labeled with an antibody. After the reaction between the primary antibody previously immobilized on the solid phase and the antigen in the sample, an enzyme-labeled antibody is added to form a primary antibody-antigen-enzyme-labeled antibody. Thereafter, the bound enzyme-labeled antibody (B) and the free enzyme-labeled antibody (F) are separated (BF separation), and the product of the cyclic reaction between the enzyme and the substrate of the bound enzyme-labeled antibody (B) is luminescence or absorbance. To obtain the amount of antigen. The competitive reaction method uses the property that the binding to the immobilized antibody competes between the unlabeled antigen to be measured and the enzyme-labeled antigen of known concentration. Since the amount of the enzyme-labeled antigen bound to the antibody depends on the concentration ratio of the antigen to be measured and the enzyme-labeled antigen, the amount of the antigen to be measured can be determined by measuring the amount of the enzyme-labeled antigen bound. . Among them, the sandwich method is highly sensitive and widely used at present. It is also used to measure peptides and proteins that are present in very small amounts in the living body, which is a disease marker.

  In recent years, health awareness has increased, and countermeasures and prevention for lifestyle-related diseases have become a problem. In particular, heart disease and cerebrovascular disease are causes of death after cancer, and prevention thereof is becoming important. For example, brain natriuretic peptide (BNP) has a vasodilatory action and a diuretic action, and plays an important role in regulating body fluid volume and blood pressure. The plasma BNP concentration in healthy individuals is extremely low, but in patients with heart failure, the value increases according to the severity. Therefore, measurement of BNP is important for understanding the pathophysiology of heart failure. The blood concentration of a healthy person is 18 ppt (pg / mL), and the measurement sensitivity of an apparatus of 10 ppt or less is required. One of the inflammatory cytokines, TNF-α (Tumor Necrosis Factor) is also said to play an extremely important role in myocardial damage, particularly heart failure. The level of TNF-α is elevated in the blood or heart tissue of patients with heart failure, which is greatly involved in the pathogenesis of heart failure. Its concentration is also very low at 1ppt or less. Cytokines are multifunctional proteins that play an important role in the development and regulation of biological responses to diseases and infections, and are particularly important factors that control immunity and inflammation. For example, alcoholic hepatitis, hepatitis, rheumatoid arthritis, spinal cord disease, head disorders and the like are known as diseases in which the concentration of interleukin-1β increases. The value of healthy individuals with interleukin-1β is 10ppt or less.

  In such a background, an enzyme immunoelectrochemical measurement method combining an enzyme immunoassay and an electrochemical detection method has been proposed as a method for easily and sensitively measuring a substance present in a very small amount in a living body. In a conventional enzyme immunization method using a sandwich method, an enzyme-labeled antibody is added after a reaction between a primary antibody immobilized on a solid phase in advance and an antigen in a sample to form a primary antibody-antigen-enzyme-labeled antibody. Thereafter, the bound enzyme-labeled antibody (B) and the free enzyme-labeled antibody (F) are separated (BF separation), and the product of the cyclic reaction between the enzyme and the substrate of the bound enzyme-labeled antibody (B) is luminescence or absorbance. As a result, the amount of antigen was obtained. In the enzyme immunoelectrochemical measurement method, the steps up to BF separation are the same as before, but the product of the cyclic reaction of the enzyme-labeled antibody (B) is adsorbed and concentrated on the silver electrode, and the amount is electrochemically determined. It is measured by a technique. At this time, cholinesterase is used as an enzyme to label the antibody, thiocholine, which is a degradation product of cholinesterase, is adsorbed and concentrated on the silver electrode, and current signal generated by reductive desorption of thiocholine adsorbed on the silver electrode in a strong alkali. Was measured and detected (Japanese Patent Laid-Open No. 2004-257996). Therefore, a very small amount of enzyme reaction product can be concentrated on the silver electrode, and highly sensitive measurement is possible. This method is an application of the reductive elimination method of thiol compounds bound to the gold surface (Langmuir 7, (1991) 2687-2693). Another example of this reductive elimination is acetylcholinesterase activity. (Sensors and Actuators B 91, (2003) 148-151).

JP 2004-257996 A

Langmuir 7, (1991) 2687-2693 Sensors and Actuators B 91, (2003) 148-151

  In the enzyme immunoelectrochemical measurement method using the electrochemical method, it is necessary to carry out the reductive elimination reaction of thiocholine adsorbed on the silver electrode in a strong alkaline solution (for example, 0.5 M KOH solution). Therefore, in the electrochemical measurement step, the solution for performing the steps up to BF separation must be replaced with a strong alkaline solution. The use of a strong alkaline solution has a safety problem, and therefore needs to be handled with care. Furthermore, in this measurement method, since the thiocholine, which is the enzyme reaction product of cholinesterase, is measured after being adsorbed to the silver electrode, the end point of the enzyme reaction is measured, and the reaction rate of the enzyme cannot be measured directly. That is, there is a problem that the measurement accuracy is lowered because the progress of the enzymatic reaction of cholinesterase cannot be measured. Even in the method of reducing and desorbing thiocholine adsorbed on the silver electrode, the amount of adsorption is estimated from the peak area of the reduction current. .

  An object of the present invention is to provide a highly sensitive and highly accurate measuring apparatus and measuring method that can be used easily without a solution exchange operation and that directly measures the enzyme reaction rate of a labeled enzyme.

  In order to achieve the above object, in the present invention, the amount or generation rate of a thiol compound that is a product of a cyclic reaction of an enzyme-labeled antibody is measured with a field effect transistor. The field effect transistor has, for example, a gold electrode in a sensing portion, and is connected to the gate of the insulated gate field effect transistor by a conductive wiring. The potential change accompanying the adsorption of the thiol compound to the gold electrode in the sensing portion is measured as a change in drain current between the source and drain of the field effect transistor. That is, an enzyme reaction for generating a thiol compound is performed in the same container, and an adsorption reaction of the thiol compound generated at that time to the gold electrode is measured as a drain current change. In addition, even when the enzyme reaction for generating the thiol compound and the adsorption reaction of the generated thiol compound to the gold electrode are performed in separate containers, the solution in the container containing the generated thiol compound is subjected to the adsorption reaction to the gold electrode. After transferring to a container, the amount of thiol compounds is similarly measured as a drain current change. At the time of measurement, an AC voltage is applied between the gold electrode and the reference electrode. Furthermore, it is preferable to use a linear polymer polymer by physical adsorption on a gold electrode in order to reduce drift during measurement.

  According to the present invention, by using a field effect transistor having a gold electrode, by measuring the drain current change accompanying the adsorption of the thiol compound, which is a product of the cyclic reaction of the enzyme labeled with the antibody, to the gold electrode, The amount of the thiol compound that is a product of the enzyme reaction or the production rate can be measured. At that time, by performing an enzyme reaction for generating a thiol compound and an adsorption reaction of the generated thiol compound to a gold electrode in the same container, the enzyme reaction can be measured in real time, and highly accurate quantification becomes possible. In addition, even when the enzyme reaction for generating the thiol compound and the adsorption reaction of the generated thiol compound to the gold electrode are performed in separate containers, the solution in the container containing the generated thiol compound is subjected to the adsorption reaction to the gold electrode. The amount of thiol compound produced can be measured with a simple operation by simply transferring it to the container. Therefore, operations up to B / F separation and enzyme reaction can be performed using conventional techniques, and there is no need to perform new operations for the measurement of the present invention. When using gold electrodes in solution, the effects of the adsorption of contaminants on the gold electrode and the drift due to ions in the solution can be easily achieved by applying an AC voltage between the gold electrode and the reference electrode. Can be excluded. Alternatively, drift can be reduced by physically adsorbing the linear polymer on the gold electrode.

1 is a block diagram showing an example of an immune analyzer according to the present invention. It is a figure which shows the structural example of the insulated gate field effect transistor used for the immunoassay apparatus of this invention, (a) is sectional drawing, (b) is a top view. The figure which shows the example of the immunoassay apparatus by this invention Block diagram showing an example of an immune analysis system according to the present invention The figure which shows the reaction flow of the immunoassay by this invention. The figure which shows the effect of the high frequency superposition method by this invention (The frequency of the applied voltage, (a): DC, (b): 10Hz, (c): 100Hz, (d): 1KHz, (e): 10KHz, (f ): 100 kHz, (g): 1 MHz, (h): 10 MHz). The figure which shows the result of having measured the thiol compound solution of a different density | concentration (The density | concentration (a) of a thiol compound: 50 micromol, (b): 25 micromol, (c): 10 micromol, (d): 5 micromol, (e): 2.5 micromol, ( f): 0.5 μM, (g): 0.1 μM). The figure which shows the relationship between the density | concentration of a thiol compound, and the adsorption rate to a gold electrode. The figure which shows the measurement result by the immunoassay apparatus of this invention. The figure which shows the relationship between the density | concentration in the sample solution measured with the apparatus of this invention, and adsorption rate. The figure which shows the relationship between the density | concentration in the sample solution measured with the apparatus of this invention, and adsorption rate. The figure which shows the effect of the linear polymer physically adsorbed on the gold electrode. The figure which shows the effect of the linear polymer physically adsorbed on the gold electrode. The figure which shows the effect at the time of combining the gold electrode which physically adsorbed the linear polymer, and the high frequency superposition method.

Embodiments of the present invention will be described below with reference to the drawings.
FIG. 1 is a block diagram showing an example of an immune analyzer using an FET sensor according to the present invention. The measurement system of the present invention includes a measurement unit 1, a signal processing circuit 2, and a data processing device 3. In the measurement unit 1, an insulated gate field effect transistor 4, a reference electrode 5, a power source 6 for applying a high frequency voltage to the reference electrode 5, a sample solution injector 7 for supplying a sample solution containing a measurement substance, and a thiol compound are generated. An enzyme-labeled antibody solution injector 8 for supplying an enzyme-labeled antibody obtained by binding an enzyme to be measured and an antibody against a measurement substance; a substrate solution injector 9 for supplying a substrate of the enzyme that generates the thiol compound; and a measurement cell 10 . In the reaction solution 11 in the measurement cell 10, an antibody fixing plate 13 on which the antibody 12 is fixed, a gold electrode 14 formed on the insulated gate field effect transistor 4, and a reference electrode 5 are arranged.

  The measurement procedure is as follows. First, the sample solution is injected into the reaction solution 11 in the measurement cell 10 using the sample solution injector 7, and the antigen and the antibody 12 in the sample solution are combined. After a certain time, an enzyme-labeled antibody solution is injected into the reaction solution 11 using the enzyme-labeled antibody solution injector 8 to cause an antigen-antibody reaction to form an antibody-antigen-enzyme-labeled antibody. Thereafter, the bound enzyme-labeled antibody and the free enzyme-labeled antibody are separated by washing the measurement cell 10 and exchanging the reaction solution 11 in the measurement cell 10. When the substrate of the labeled enzyme is injected using the substrate solution injector 9 after washing and solution exchange in the measurement cell 10, the substrate is decomposed by the enzyme, and a thiol compound is generated. The generated thiol compound is adsorbed on the gold electrode 14 formed on the insulated gate field effect transistor 4 to form a self-assembled film. As a result, the potential on the gold electrode 14 changes. In the measurement, the current between the source 15 and the drain 16 in the insulated gate field effect transistor 4 changing before and after the substrate injection by the substrate solution injector 9 is monitored in real time and recorded by the signal processing circuit 2 and the data processing device 3. To do. The adsorption rate of the thiol compound to the gold electrode 14 is proportional to the production rate of the thiol compound, that is, the amount of antibody-antigen-enzyme labeled antibody. Therefore, by measuring the adsorption rate of the thiol compound to the gold electrode 14, the amount of bound labeled enzyme, that is, the amount of antigen in the sample solution can be obtained. At that time, a high frequency voltage is applied to the reference electrode 5 by the power source 6 during the measurement in order to reduce the influence due to the measurement external variation.

  For the sample solution injector 7, enzyme-labeled antibody solution injector 8, and substrate solution injector 9, a syringe pump or a pressurized liquid delivery device can be used. If the volume to be injected is 1 μL or more, both a syringe pump and a pressurized liquid delivery device can be used, but if the injection volume is 1 μL or less, a pressurized liquid delivery device using a capillary for the resistance tube is desirable. For example, when the injection volume is 0.2 μL, a flow rate control capillary having an inner diameter of 25 μm and a length of 20 mm can be used, and the injection can be accurately performed under the conditions of a pressure of 2 atm and a pressure time of 2 seconds. .

  The reference electrode 5 gives a reference potential in order to stably measure a potential change based on an equilibrium reaction or a chemical reaction occurring on the surface of the gold electrode 14 in the reaction solution 11. Usually, the reference electrode is a silver / silver chloride electrode or calomel electrode using saturated potassium chloride as the internal solution. However, if the composition of the sample solution to be measured is constant, There is no problem even if only a silver / silver chloride electrode is used as an electrode.

FIG. 2 is a diagram showing the structure of an insulated gate field effect transistor used in the immunoassay apparatus of the present invention. 2A and 2B show a cross-sectional structure and a planar structure, respectively. The insulated gate field effect transistor 21 has a source 22, a drain 23, and a gate insulator 24 formed on the surface of a silicon substrate, and a gold electrode 25. The gold electrode 25 and the gate 26 of the insulated gate field effect transistor are connected by a conductive wiring 27. Preferably, the insulated gate field effect transistor is a metal-oxide semiconductor field effect transistor (FET) using silicon oxide as an insulating film, but there is no problem even if a thin film transistor (TFT) is used. By adopting this structure, the gold electrode 25 can be formed at any location and in any size, and the volume of the measurement cell can be changed according to the amount of the sample solution to be measured. The insulated gate field effect transistor used in the present invention is a depletion type FET having an insulating layer using SiO ₂ (thickness: 17.5 nm), and a gold electrode is fabricated in a size of 400 μm × 400 μm. . Since normal measurement uses an aqueous solution, the device must operate in solution. When measuring in a solution, it is necessary to operate in an electrode potential range of −0.5 to 0.5 V which hardly causes an electrochemical reaction. Therefore, in this embodiment, the fabrication condition of the depletion type n-channel FET, that is, the ion implantation condition for adjusting the threshold voltage (Vt) is adjusted, and the threshold voltage of the FET is set to around −0.5V. In place of the gold electrode, an electrode made of another noble metal such as silver may be used.

  FIG. 3 is a diagram showing another configuration example of the immune analyzer using the FET sensor according to the present invention. In the insulated gate field effect transistor 31 used in this embodiment, a source 32, a drain 33, and a gate insulator 34 are formed on the surface of a silicon substrate, and a gold electrode 35 is formed on the surface of the gate insulator between the source 32 and the drain 33. It is provided. An antibody 36 is immobilized on the surface of the gold electrode 35.

  In actual measurement, the gold electrode 35, the antibody 36 immobilized on the surface of the gold electrode 35, and the reference electrode 37 are placed in the reaction solution 39 in the measurement cell 38, and the power source 40 is connected to the reference electrode 37. A high frequency voltage is applied to detect an enzyme reaction product occurring in the reaction solution 39 as a change in electrical characteristics of the insulated gate field effect transistor 31 that changes before and after substrate injection, that is, a change in current between the source 32 and the drain 33. Thus, the amount of antigen in the sample solution bound to the antibody 36 can be measured.

  The measurement procedure is as follows. First, the sample solution is injected into the reaction solution 39 in the measurement cell 38 by using the sample solution injector 7 to bind the antigen 36 and the antibody 36 in the sample solution. After a certain time, an enzyme-labeled antibody solution is injected into the reaction solution 39 using the enzyme-labeled antibody solution injector 8 to cause an antigen-antibody reaction, thereby forming an antibody-antigen-enzyme-labeled antibody. Thereafter, the bound enzyme-labeled antibody and the free enzyme-labeled antibody are separated by exchanging the reaction solution 39 in the measurement cell 38 and washing the measurement cell 38. When the substrate of the labeled enzyme is injected using the substrate solution injector 9 after exchanging and washing the solution in the measurement cell 38, the substrate is decomposed by the enzyme and a thiol compound is generated. The generated thiol compound is adsorbed on the gold electrode 35 formed on the insulated gate field effect transistor 31 to form a self-assembled film. As a result, the potential on the gold electrode 35 changes. The adsorption rate of the thiol compound to the gold electrode 35 is proportional to the production rate of the thiol compound, that is, the amount of antibody-antigen-enzyme labeled antibody. Therefore, by measuring the adsorption rate of the thiol compound to the gold electrode 35, the amount of bound labeled enzyme, that is, the amount of antigen in the sample solution can be obtained. At that time, the immobilization onto the gold electrode 35 can be achieved by using an Fab ′ fragment that is a part of the antibody or an aptamer that is a single-stranded DNA in addition to the antibody.

  FIG. 4 is a block diagram showing an example of an immune analysis system using an FET sensor according to the present invention. This analysis system includes a measurement unit 41, a signal processing circuit 42, a data processing device 43, and a reaction vessel 44 for a thiol compound generation reaction. In the measurement unit 41, an insulated gate field effect transistor 45, a reference electrode 46, a power source 47 for applying a high-frequency voltage to the reference electrode 46, and a thiol compound solution injector 48 for supplying a solution in the reaction vessel 44 are provided. . A gold electrode 51 and a reference electrode 46 formed on the insulated gate field effect transistor 45 are disposed in the reaction solution 50 in the measurement cell 49. An antibody 52 is immobilized on an antibody fixing plate 53 in a reaction container 44 for thiol compound generation reaction. The antibody 52 may be directly immobilized inside the reaction container 44.

  The measurement procedure is as follows. The sample solution is injected into the reaction container 44 for the thiol compound generation reaction, and the antigen and the antibody 52 in the sample solution are combined. After a certain period of time, an enzyme-labeled antibody solution is injected into the reaction container 44 to cause an antigen-antibody reaction to form an antibody-antigen-enzyme-labeled antibody. Thereafter, the bound enzyme-labeled antibody and the free enzyme-labeled antibody are separated by exchanging the solution in the reaction container 44 and washing the reaction container 44. When the substrate of the labeling enzyme is injected after the solution in the reaction container 44 is exchanged and washed, the substrate is decomposed by the enzyme to generate a thiol compound. After the reaction for a certain time, the generated thiol compound is introduced into the reaction solution 50 in the measurement cell 49 using a thiol compound solution injector 48. The thiol compound introduced into the reaction solution 50 in the measurement cell 49 is adsorbed to the gold electrode 51 formed on the insulated gate field effect transistor 45 to form a self-assembled film. As a result, the potential on the gold electrode 51 changes. In the measurement, the current between the source 54 and the drain 55 in the insulated gate field effect transistor 45 that changes before and after the injection of the thiol compound generated by the thiol compound solution injector 48 is monitored in real time, and the signal processing circuit 42 and the data processing device are monitored. This is done by recording at 43. The adsorption rate of the thiol compound to the gold electrode 51 is proportional to the concentration of the thiol compound, that is, the amount of antibody-antigen-enzyme labeled antibody. Therefore, by measuring the adsorption rate of the thiol compound to the gold electrode 51, the amount of bound labeled enzyme, that is, the amount of antigen in the sample solution can be obtained.

FIG. 5 shows a reaction flow in the immunoassay apparatus using the FET sensor according to the present invention.
In the immunoassay, a specific binding reaction between an antigen and an antibody is used to measure the amount of binding between the antigen and the antibody to obtain the amount of antigen. In the present invention, the amount of antigen is indirectly measured through an enzyme labeled with an antibody using a sandwich method generally used in conventional immunoassay. After the reaction between the antibody 61 previously immobilized on the solid phase and the antigen 62 in the sample, an enzyme-labeled antibody 63 is added to form an antibody-antigen-enzyme-labeled antibody 64. Thereafter, the bound enzyme-labeled antibody 65, the free enzyme-labeled antibody 63, and the free antigen 62 are separated, and the thiol compound 68, which is a product of the cyclic reaction of the enzyme 66 and the substrate 67 of the bound enzyme-labeled antibody, is converted into an FET sensor. The amount of antigen can be obtained by measuring with

The effect of the AC voltage application of the present invention will be described using another embodiment. 6A to 6H show the change in drain current over time when an AC voltage is applied to the reference electrode 46 shown in FIG. 4 and the sample solution is introduced into the reaction solution 50 in the measurement cell 49. FIG. FIG. FIGS. 6B to 6H show the results when AC voltages having frequencies of 10 Hz, 100 Hz, 1 KHz, 10 KHz, 100 KHz, 1 MHz, and 10 MHz are applied, respectively. In order to see the effect of alternating voltage application, direct current (DC) application data is shown in FIG. 6A as a reference experiment. As the reference electrode 46, an Ag / AgCl reference electrode was used. The AC voltage was applied to the reference electrode 46 at a center voltage of 100 mV and an amplitude voltage of 100 mV. The reaction solution used was 1.9 ml of 0.1M Na ₂ SO ₄ aqueous solution. For the sample solution, 1 mM 6-hydroxy-1-hexanethiol (6-HHT) aqueous solution was used as the alkanethiol compound. The current / voltage characteristics of the transistor were measured using a semiconductor parameter analyzer (Agilent 4155C Semiconductor Parameter Analyzer).

  After 600 seconds from the start of measurement (arrow in the figure), 0.1 mL of the sample solution was introduced into the reaction solution. After introducing the sample solution, a decrease in drain current was observed in all cases. When DC was applied, the stability was poor and the drift from the start of measurement to the introduction of the sample solution was large. Furthermore, after the sample solution was introduced, the drain current value increased again after decreasing, and it took more than 10 minutes to stabilize. This tendency was seen when the frequency was 1 KHz or less. On the other hand, when the frequency was 10 KHz or more, the drain current value hardly increased after the decrease, and the time to stabilization was shorter than that when the frequency was 1 KHz or less. Also, the drain current from the start of measurement to the introduction of the sample solution was stable with little drift. These results show that the surface potential of the gold electrode becomes unstable due to the introduction of the sample solution, but when the frequency of the alternating current applied to the reference electrode is 10 KHz or higher, the surface potential of the gold electrode is disturbed quickly. This is thought to be due to the greater recovery effect. Thus, by superimposing a high frequency of 10 KHz or more on the reference electrode during measurement, the reaction process occurring on the gold electrode can be accurately measured.

FIG. 7 is a diagram showing the results of measuring sample solutions having different concentrations. In the sample solution, 6-HHT aqueous solution was used as the alkanethiol compound. 7A to 7G show the results when the final concentrations in the reaction solution are 50 μM, 25 μM, 10 μM, 5 μM, 2.5 μM, 0.5 μM, and 0.1 μM, respectively. In order to see the effect of physical adsorption on the gold electrode, an HDO (hexanethidiol) aqueous solution was used as a reference experiment (final concentration: 50 μM). The arrow in the figure represents the time when the use solution was introduced. As the reference electrode 46, an Ag / AgCl reference electrode was used. The AC voltage was applied to the reference electrode 46 at a center voltage of 100 mV, an amplitude voltage of 100 mV, and a frequency of 1 MHz. The reaction solution used was 1.9 ml of 0.1M Na ₂ SO ₄ aqueous solution. As shown in FIGS. 7A and 7B, when the concentration in the reaction solution was 50 μM or 25 μM, the drain current value became constant within a few seconds, and the adsorption reaction to the gold electrode was completed. Further, as shown in FIGS. 7D and 7E, when the concentration was 5 μM or 2.5 μM, it took about 5 minutes and 10 minutes for the drain current value to be constant and the reaction to be completed, respectively. On the other hand, as shown in FIGS. 7F and 7G, when the concentration was 0.5 μM or 0.1 μM, it took 1 hour or more for the drain current value to become constant. Further, as shown in FIG. 7 (h), in the case of the HDO aqueous solution, the drain current value hardly changes, indicating that there is no influence of physical adsorption on the gold electrode.

  Thus, since the time until the drain current value becomes constant changes according to the concentration in the reaction solution, if the time until the drain current value after the sample solution introduction becomes constant is measured or estimated, The concentration of the alkanethiol compound in the sample solution can be obtained. Also, instead of measuring or estimating the time until the drain current value becomes constant after the sample solution is introduced, the rate of adsorption of the alkanethiol compound to the gold electrode is used from the amount of change in the drain current value immediately after the sample solution is introduced. The concentration in the solution can also be obtained. The adsorption rate to the gold electrode may be a function y = F (x) obtained by fitting the drain current value immediately after introduction of the sample solution by a least square method using a polynomial or exponential function. Alternatively, the differential value of F (x) or the amount of change over a certain time may be used as the initial rate of the adsorption reaction immediately after the sample solution is introduced.

  FIG. 8 is a diagram showing the relationship between the concentration in the sample solution and the adsorption rate of the measurement result on the gold electrode. The measurement was performed three times, and the average was plotted in the figure. The adsorption rate was determined by calculating the tangent of the drain current curve immediately after introduction of the sample solution as the initial rate. As shown in FIG. 8, the concentration in the sample solution and the adsorption rate showed good linearity. Thus, instead of measuring the time until the drain current value becomes constant and using the time until the adsorption reaction is completed, the concentration in the sample solution can be obtained quickly and accurately from the adsorption rate on the gold electrode. be able to.

The enzyme immunoassay method using the apparatus of the present invention will be described below. In this example, the amount of antigen was indirectly measured through an enzyme labeled with an antibody using a sandwich method generally used in conventional immunoassay. After the reaction between the antibody immobilized on the plate and the antigen in the sample, an enzyme-labeled antibody was added to form an antibody-antigen-enzyme-labeled antibody. Thereafter, the bound enzyme-labeled antibody was separated from the free enzyme-labeled antibody and the free antigen, and the thiol compound that was a product of the cyclic reaction between the enzyme of the bound enzyme-labeled antibody and the substrate was measured with an FET sensor. Samples and reagents used in this example are shown below.
Immobilized antibody: Interleukin 1β antibody Sample: Human plasma
Measurement object: Interleukin 1β
Enzyme-labeled antibody: Acetylcholinesterase (AChE): Interleukin-1β Fab 'Conjugate
Substrate: 2.5 mM Acetylthiocholine
Reaction solution: 0.1 M phosphate buffer (pH 7.4), 0.15 M NaCl, 1 mM EDTA
Note that the reaction conditions and reagent concentrations used here are merely examples, and can be appropriately changed according to the apparatus configuration and the measurement target.

  The measurement procedure is as follows. First, 100 μL of sample solution (Human plasma) and 100 μL of enzyme-labeled antibody (AChE: Interleukin-1β Fab 'Conjugate) are added to the wells of the plate on which the antibody of Interleukin 1β is immobilized, and the plate is covered with a plastic film. React overnight at 4 ° C. Thereafter, the plate well solution is discarded and washed 5-6 times with wash buffer. An acetylthiocholine solution, which is a substrate for acetylcholinesterase, is added to each well and allowed to react for about 30 minutes. The reaction solution containing the thiol compound produced by the reaction is introduced into the reaction cell immersed in the FET sensor, and the concentration of the thiol compound produced by the reaction is obtained by measuring the adsorption rate of the thiol compound to the gold electrode. . Since the concentration of the produced thiol compound is proportional to the enzyme concentration of the antibody-antigen-enzyme labeled antibody, the amount of antigen can be quantified.

  In the measurement with the FET sensor, an Ag / AgCl reference electrode was used as a reference electrode. An AC voltage having a center voltage of 100 mV, an amplitude voltage of 100 mV, and a frequency of 1 MHz was applied to the reference electrode. The measurement result by the FET sensor is shown in FIG. The data in FIGS. 9A to 9D show the measurement results of samples having concentrations of Interleukin 1β as an antigen of 200, 20, 2.0, and 1.0 pg / mL, respectively. The data in FIG. 9 (e) is a blank with a concentration of Interleukin 1β of 0 pg / mL. In the present invention, the concentration of Interleukin 1β in the sample was determined as the adsorption rate to the gold electrode. The adsorption rate was determined by calculating the initial rate of the tangent line (dotted line in the figure) of the drain current curve immediately after the introduction of the sample solution (arrow in the figure).

  As a result, as shown in FIG. 10, the concentration in the sample solution and the adsorption rate showed good linearity and could be measured up to 1.0 pg / mL. This detection sensitivity is one digit higher than the conventional method. In this way, the FET sensor can be used without using a spectrophotometer by measuring the adsorption rate of the thiol compound produced by the sandwich-type enzyme reaction commonly used in conventional immunoassay to the gold electrode. Thus, the concentration in the sample solution can be measured with high sensitivity.

Another embodiment of the enzyme immunoassay using the apparatus of the present invention will be described below. In this example, the thiol compound of the enzyme immune reaction product was measured in real time in the same reaction cell with the apparatus configuration shown in FIG. In the measurement with the FET sensor, an Ag / AgCl reference electrode was used as a reference electrode. An AC voltage having a center voltage of 100 mV, an amplitude voltage of 100 mV, and a frequency of 1 MHz was applied to the reference electrode. Samples and reagents used in this example are shown below.
Immobilized antibody: Interleukin 1β antibody Sample: Human plasma
Measurement object: Interleukin 1β
Enzyme-labeled antibody: Acetylcholinesterase (AChE): Interleukin-1β Fab 'Conjugate
Substrate: 2.5 mM Acetylthiocholine
Reaction solution: 0.1 M phosphate buffer (pH 7.4), 0.15 M NaCl, 1 mM EDTA

  The measurement procedure is as follows. First, an antibody immobilization plate on which an antibody of Interleukin 1β is immobilized is placed in the reaction cell. 100 μL of sample solution (Human plasma) and 100 μL of enzyme-labeled antibody (AChE: Interleukin-1β Fab ′ Conjugate) are added to the reaction cell, and the reaction cell is covered with a plastic film and allowed to react at 4 ° C. overnight. Thereafter, the solution in the reaction cell is discarded and washed 5 to 6 times with a washing buffer. Add 1.8 mL of the reaction solution (0.1 M phosphate buffer (pH 7.4), 0.15 M NaCl, 1 mM EDTA) to the reaction cell, and start the measurement. When 0.2 mL of acetylthiocholine solution is introduced into the reaction cell 600 seconds after the start of measurement, a thiol compound is generated by an enzymatic reaction. The drain current value of the FET sensor changes due to adsorption of the thiol compound generated by the reaction to the gold electrode. By measuring the drain current value in real time, the generation rate of the thiol compound can be accurately measured. Since the production rate of the thiol compound depends on the enzyme amount of the antibody-antigen-enzyme labeled antibody, the amount of the labeled enzyme bound, that is, the antigen amount can be obtained from the production rate of the thiol compound.

  In this example, the concentration of Interleukin 1β in the sample was determined by calculating the adsorption rate on the gold electrode with the tangent to the drain current curve immediately after introducing the sample solution as the initial rate. FIG. 11 shows the measurement result by the FET sensor. As shown in FIG. 11, the concentration in the sample solution and the adsorption rate showed good linearity.

Next, the effect of physically adsorbing the linear polymer on the gold electrode will be described. As the linear polymer, polyethylene glycols having different molecular weights were used. FIG. 12 is a diagram showing the change with time of the drain current accompanying the adsorption of the thiol compound to the electrode on which polyethylene glycol is physically adsorbed. FIGS. 12B to 12F are data of gold electrodes coated with 0.5% polyethylene glycol aqueous solutions having molecular weights of 1000, 2000, 8000, 500000, and 2000000, respectively. FIG. 12A shows a case where an untreated gold electrode is used as a reference. The thiol compound used is a 1 mM 6-HHT aqueous solution. As the reaction solution, 1.9 ml of 0.1M Na ₂ SO ₄ aqueous solution was used, and 0.1 mL of 6-HHT aqueous solution was introduced 600 seconds after the start of measurement. An Ag / AgCl reference electrode was used as the reference electrode. In order to see the effect of the linear polymer, no DC voltage was applied to the reference electrode, and DC 100 mV was applied. The current / voltage characteristics of the transistor were measured using a semiconductor parameter analyzer (Agilent 4155C Semiconductor Parameter Analyzer).

  As a result, as shown in FIG. 12, in the untreated gold electrode, the baseline was greatly fluctuated before introducing the 6-HHT aqueous solution, but the baseline was stable when polyethylene glycol was physically adsorbed. In particular, when the molecular weight was 2000 or more, the baseline fluctuation disappeared and became stable.

Moreover, the measurement result at the time of using dextran as another linear polymer is shown in FIG. FIG. 13 is a diagram showing a change with time in drain current accompanying adsorption of a thiol compound to an electrode on which dextran is physically adsorbed. FIGS. 13B to 13D are data of gold electrodes coated with 0.5% dextran solutions having molecular weights of 40,000, 90000, and 200,000, respectively. FIG. 13A shows an untreated gold electrode as a reference. The thiol compound used is a 1 mM 6-HHT aqueous solution. As the reaction solution, 1.9 ml of 0.1 M Na ₂ SO ₄ aqueous solution was used, and 0.1 mL of 6-HHT aqueous solution was introduced 600 seconds after the start of measurement. An Ag / AgCl reference electrode was used as the reference electrode. In order to see the effect of the linear polymer, no DC voltage was applied to the reference electrode, and DC 100 mV was applied. The current / voltage characteristics of the transistor were measured using a semiconductor parameter analyzer (Agilent 4155C Semiconductor Parameter Analyzer). As a result, as in the case of FIG. 12, in the untreated gold electrode, the baseline was greatly fluctuated before the introduction of the 6-HHT aqueous solution, but when dextran was physically adsorbed to the gold electrode, the baseline was stabilized. .

Moreover, the measurement result at the time of combining the gold electrode which physically adsorbed these linear polymers and the high frequency superposition method is shown in FIG. FIGS. 14A, 14B, and 14C correspond to an untreated gold electrode, a gold electrode coated with dextran (molecular weight: 2000000), and a gold electrode coated with polyethylene glycol (molecular weight: 500000), respectively. A 0.5% aqueous solution was used for physical adsorption of the polymer on the gold electrode. The thiol compound used is a 1 mM 6-HHT aqueous solution. As the reaction solution, 1.9 ml of 0.1M Na ₂ SO ₄ aqueous solution was used, and 0.1 mL of 6-HHT aqueous solution was introduced 600 seconds after the start of measurement. An Ag / AgCl reference electrode was used as the reference electrode. The AC voltage applied to the reference electrode was a center voltage of 100 mV, an amplitude voltage of 100 mV, and a frequency of 1 MHz. The current / voltage characteristics of the transistor were measured using a semiconductor parameter analyzer (Agilent 4155C Semiconductor Parameter Analyzer).

  As a result, as shown in FIG. 14 surrounded by a broken line, even in the case of an untreated gold electrode, the baseline fluctuation before introduction of the 6-HHT aqueous solution was small and stable. When dextran or polyethylene glycol was coated, there was almost no baseline fluctuation, and there was a synergistic effect of the linear polymer coating and the high-frequency superposition method.

  DESCRIPTION OF SYMBOLS 1,41 ... Measuring part, 2,42 ... Signal processing circuit, 3,43 ... Data processing device, 4, 21, 31, 45 ... Insulated gate field effect transistor, 5, 37 ... Reference electrode, 6, 40, 47 ... Power source, 7 ... Sample solution injector, 8 ... Enzyme-labeled antibody solution injector, 9 ... Substrate solution injector, 10,38,49 ... Measurement cell, 11,39,50 ... Reaction solution, 12,36,52,61 ... antibody, 13, 53 ... antibody fixing plate, 14, 25, 35, 51 ... gold electrode, 15, 22, 32, 54 ... source, 16, 23, 33, 55 ... drain, 24, 34 ... gate insulator, 26 ... gate of insulated gate field effect transistor, 27 ... conductive wiring, 44 ... reaction vessel, 48 ... thiol compound solution injector, 62 ... antigen, 63 ... enzyme labeled antibody, 64 ... antibody-antigen-enzyme labeled antibody, 65 ... anti-conjugated enzyme label, 6 ... the enzyme-labeled antibody binding, 67 ... substrate, 68 ... thiol compound.

Claims (2)

An analytical element for measuring a thiol compound,
A field effect transistor;
An electrode made of a noble metal physically adsorbed on the surface of a linear polymer,
The gate of the field effect transistor and the electrode are connected by a conductive wiring ,
The thiol compound measuring element , wherein the linear polymer is dextran or polyethylene glycol . Before SL noble metals thiol compound measuring device according to claim 1, wherein the gold or silver.

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