JP4343894B2 - Biological substance measuring apparatus and measuring method - Google Patents

Description

  The present invention relates to a measuring apparatus and a measuring method capable of easily measuring biological materials such as proteins without a washing operation.

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. The sandwich method is an antibody (labeled secondary antibody) in which the antigen to be measured is bound to an immobilized and insolubilized antibody (solid-phased antibody) on a solid surface, and this antigen is sandwiched with the solid-phased antibody. Uses the property of binding to the antigen. The amount of the antigen to be measured is determined by measuring the amount of the labeled secondary antibody bound to the antigen. The competitive reaction method uses the property that the binding to the immobilized antibody competes between the unlabeled antigen to be measured whose concentration is unknown and the enzyme-labeled antigen whose concentration is known. Since the binding amount of the labeled antigen is determined by the concentration ratio between the antigen to be measured and the labeled antigen, the amount of the antigen to be measured is determined by measuring the binding amount of the labeled antigen (Clinical Chemistry 22, (1976 ) 1243-1255). In addition to the enzyme, a fluorescent substance, a radioisotope, or the like may be used for labeling. In any case, in order to distinguish the bound antibody or antigen from the unbound antibody or antigen, the unbound antibody or antigen is used. Remove by washing operation. Then, the reaction for binding the antibody and the antigen is performed until the bound state and the unbound state reach equilibrium. Some antibodies use single-stranded DNA (aptamer DNA) that specifically binds to a specific protein instead of an antibody (Analytical Biochemistry 282, (2000) 142-146).
Clinical Chemistry 22, (1976) 1243-1255 Analytical Biochemistry 282, (2000) 142-146

  In the enzyme immunoassay, unbound antibody or antigen cannot be distinguished from bound antibody or antigen. Therefore, in order to measure bound antibody or antigen, it is necessary to remove the unbound antibody or antigen by washing. Yes, this work was complicated. Furthermore, since the washing operation may destroy the equilibrium state between the bound antibody and the antigen, skill is required for the operation. In addition, the time required for the bound state and the unbound state to reach an equilibrium state must be determined by other experiments or determined based on experience.

  An object of the present invention is to provide a biological substance measuring apparatus and measuring method using a biomolecule detection element that can be used easily without a washing operation and can measure a binding reaction process.

  In order to achieve the above object, in the present invention, a field effect transistor having a conductive electrode is used as a biomolecule detection element. A field-effect transistor having a conductive electrode responds only to molecules bound to the surface of the conductive electrode and does not respond to non-bonded molecules. Therefore, a cleaning operation is performed to remove the unbound molecules. In addition, the presence or absence of binding of the molecule to be measured to the electrode surface can be measured. Therefore, by immobilizing the biomolecule detection probe on the conductive electrode, the biomolecule detection probe of the biomolecule to be measured can be obtained without performing a cleaning operation to remove the biomolecule not bound to the biomolecule detection probe. The presence or absence of binding to can be measured.

  By calculating the difference in the output of the field-effect transistor before and after introducing the sample solution into the measurement solution, the external variation is removed, and the field-effect type by coupling the measurement target molecule to the conductive electrode or biomolecule detection probe Only the change in the output of the transistor can be measured. Furthermore, by measuring and comparing changes in the output of the field-effect transistor after introducing the sample solution multiple times, it is known that the reaction of binding of the molecule to be measured to the conductive electrode or biomolecule detection probe has reached equilibrium. be able to.

  The influence of the electric double layer on the electrode, which becomes a problem when using the conductive electrode in the solution, can be easily removed by applying an AC voltage between the conductive electrode and the reference electrode. Note that the coupling between the detection probe and the measurement object is not released by the application of the AC voltage. Further, by using a noble metal such as gold for the conductive electrode, no reaction occurs on the electrode surface in the solution. Furthermore, since this AC voltage application can suppress the surface potential from becoming unstable when the sample solution is injected, the reaction process of the binding to the biomolecule detection probe can be measured. Noise generated in the output of the field effect transistor by applying an AC voltage between the conductive electrode and the reference electrode can be removed by averaging in consideration of the frequency of the AC voltage.

  According to the present invention, by using a field effect transistor having a conductive electrode as a surface reaction detection element, by measuring the time change of the reaction on the surface of the conductive electrode caused by introducing the measurement target, The reaction process of the binding between the surface of the conductive electrode or the measurement target and the biomolecule detection probe can be measured. Therefore, no washing operation is required, the measurement can be easily performed, and the end of the reaction due to the addition of the measurement target can be detected. In addition, since a cleaning operation is not required, a device for automating cleaning and a liquid feeding device (flow device) are not required, and the device is downsized. The influence of the electric double layer on the electrode surface, which is a problem during measurement, can be easily removed by applying an AC voltage between the electrode and the reference electrode.

Embodiments of the present invention will be described below with reference to the drawings.
FIG. 1 is a diagram showing a configuration example of a biomolecule detection element according to the present invention. In the insulated gate field effect transistor 1 used in the present invention, a source 2, a drain 3 and a gate insulator 4 are formed on the surface of a silicon substrate, and a conductive electrode 5 is provided on the surface of the gate insulator between the source and drain. is there. In actual measurement, the conductive electrode 5 and the reference electrode 7 are arranged in the measurement solution 9 in the measurement cell 8, and a DC voltage (bias component) and an AC voltage are applied to the reference electrode 7 by the power source 10. By injecting the sample solution 12 with the sample injector 11 and detecting the change in the electric characteristics of the insulated gate field effect transistor 1, that is, the change in the current value flowing between the source 2 and the drain 3, Measure the process of reaction occurring on the surface. Reference numeral 13 denotes a sample injection tube. As the measurement solution, an aqueous solution of sodium sulfate, an aqueous solution of sodium chloride, a tris hydrochloric acid buffer, a phosphate buffer, or the like is used.

  By immobilizing the biomolecule detection probe 6 on the surface of the conductive electrode 5, it is possible to measure the reaction process of the measurement object such as DNA or protein contained in the sample solution 12 with the biomolecule detection probe 6. . The biomolecule detection probe 6 can use nucleic acids such as single-stranded DNA fragments, antibodies, antigens, proteins and peptides such as enzymes, saccharides, and the like. The selectivity of the biomolecule detection probe is based on the difference in specific binding force (affinity) derived from the structure unique to the biological component. The reference electrode 7 provides 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 conductive electrode 5 in the measurement solution 9. 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. By applying a predetermined voltage to the reference electrode 7, the operating point (that is, the threshold voltage) of the electrical characteristics of the insulated gate field effect transistor 1 can be adjusted.

Preferably, the insulated gate field effect transistor 1 is a metal oxide semiconductor field effect transistor (FET) using silicon oxide as an insulating film, but a thin film transistor (TFT) may also be used. No problem. Here, an example has been described in which a biomolecule detection probe is fixed on a conductive electrode, but an ion-sensitive film may be formed instead of the biomolecule detection probe. For example, in the case of pH measurement, a solid film such as silicon nitride (Si ₃ N ₄ ) or tantalum oxide (Ta ₂ O ₅ ) is formed as an ion-sensitive film, and in the case of potassium ion, a liquid film containing valinomycin is formed. That's fine.

  FIG. 2 is a block diagram showing a biomolecule measuring apparatus using the biomolecule detecting element according to the present invention. The measurement system of the present invention includes a measurement unit 20, a signal processing circuit 21, a data processing device 22, and a data display storage device 23. An insulated gate field effect transistor 1, a reference electrode 7, and a sample injector 11 are disposed in the measurement unit 20. The current value of the insulated gate field effect transistor 1 is transmitted to the signal processing circuit 21.

  The measurement procedure is as follows. First, the conductive electrode 5, the biomolecule detection probe 6 immobilized on the surface of the conductive electrode 5, and the reference electrode 7 are placed in the measurement solution 9 in the measurement cell 8, and the power supply 10 is applied to the reference electrode 7. An AC voltage including a bias component is applied. A sample solution 12 is introduced into the measurement solution 9 in the measurement cell 8 using the sample injector 11. When the molecules in the introduced sample solution 12 are bonded to the conductive electrode 5 or the biological material in the sample solution 12 is bonded to the biomolecule detection probe 6 immobilized on the surface of the conductive electrode 5, insulation is caused. The electrical characteristics of the gate field effect transistor 1 change. The electrical characteristic change is processed by the signal processing circuit 21, data processing is performed by the data processing device 22, and data is displayed and stored by the data display storage device 23.

  FIG. 3 is a block diagram showing a signal processing method in the signal processing circuit 21. The signal processing circuit 21 includes a sampling device 42 and an averaging device 43. The input signal 41 is an output value from the measurement unit 20, that is, a current value of the insulated gate field effect transistor 1.

The signal processing procedure is as follows. The input signal 41 is sampled at the sampling frequency f _s . The value obtained by each sampling is averaged n times by the averaging device 43. The averaged value is transmitted as an output signal 44 to the data processing device 22.

  This signal processing is performed in order to remove a component generated due to the response of the insulated gate field effect transistor 1 to the high frequency applied from the power supply 10. When the frequency of the high frequency applied from the power supply 10 is f and the time is t, the output from the transistor for the applied high frequency is expressed by the following equation.

When the component represented by the equation (1) is given to the input signal 41, the output signal 44 is obtained by the equation (2) using the sampling frequency f _s of the sampling device 42 and the averaging number n of the averaging device 43. expressed.

Here, if the following expression (3) holds, the value of expression (2) becomes 0 and the output signal 44 becomes 0. By setting the sampling frequency f _s of the sampling device 42 and the averaging number n of the averaging device 43 so that the condition of the expression (3) is satisfied, the insulated gate field effect type is applied to the high frequency applied from the power supply 10. A noise component generated due to the response of the transistor 1 can be completely removed.

However, as shown in FIG. 4, the drain current value of the transistor is not completely linear with respect to the gate voltage. Here, the current / voltage characteristics of the transistors were measured with a configuration shown in FIG. 5 using a semiconductor parameter analyzer (Agilent 4155C Semiconductor Parameter Analyzer). The drain-source voltage 107 is applied between the source 2 and the drain 3 of the insulated gate field effect transistor 1, the gate voltage 6 (V _G ) is applied between the source 2 and the conductive electrode 5, and the drain 3 The current I _D flowing through was measured with an ammeter 108. The drain-source voltage 107 was 0.5V.

  When the amplitude of the applied high frequency is small, there is no problem even if the drain current value is considered to be substantially linear with respect to the gate voltage. Therefore, even if the drain current value is used as the output from the measurement unit 20, the expression (1) is established. However, when the amplitude of the applied high frequency is so great that the non-linearity of the drain current value with respect to the gate voltage cannot be ignored, Equation (1) does not hold if the drain current value is used as the output from the measurement unit 20. Therefore, by converting the drain current value of the output from the measurement unit 20 into a gate voltage using the relationship of FIG. 4, Equation (1) is established.

  As a method of converting the drain current value into the gate voltage, there is a method of using an electric circuit in addition to a method of converting by calculation of a computer using the relationship of FIG. FIG. 6 is a block diagram illustrating a method for converting a drain current value into a gate voltage using an electric circuit. It comprises a measurement insulated gate field effect transistor 51, a measurement unit 53, a comparison insulated gate field effect transistor 52, a comparison unit 54, and a feedback circuit 55. The feedback circuit 55 is used for the conductivity of the comparison insulated gate field effect transistor 52 so that the drain current value output from the measurement unit 53 and the drain current value output from the comparison unit 54 have the same value. A voltage is applied to the electrode 56. Then, by using the applied voltage as an output, the drain current value can be converted into a gate voltage.

FIG. 7A is a flowchart showing a data processing method in the data processing apparatus. In step 11, an alternating voltage of 1 KHz or more including a bias component is applied to the reference electrode 7 by the power source 10. In step 12, wait for at least 5 minutes after applying the voltage. Records output from the signal processing circuit 21 at step 13, and x _0. In step 14, the sample solution 12 is introduced into the measurement solution 9 in the measurement cell 8 using the sample injector 11. This time is time t = 0. In step 15, the output from the signal processing circuit 21 is recorded as x ₁ and the time is recorded as t ₁ . The output from the signal processing circuit 21 and recorded x ₂ in step 16, records the time and t _2. In step 17, the slope s of the values recorded in steps 15 and 16 is calculated using the following equation.

  In step 18, it is determined whether the reaction on the conductive electrode 5 has reached equilibrium and the output from the signal processing circuit 21 has become stable. The determination is performed using the following equation.

  If this formula is true, the process proceeds to step 19, and if this formula is not true, the process returns to step 15. In step 19, a change Δx in output from the signal processing circuit 21 due to the introduction of the sample solution 12 into the measurement solution 9 is calculated using the following equation and output.

  The determination in step 18 is based on the following grounds. It is assumed that the change Δx in the output from the signal processing circuit 21 due to the introduction of the sample solution 12 into the measurement solution 9 follows the relaxation curve of the following equation.

Here, Δx _T is the change in output when the equilibrium state is reached, t is the time when the sample injection time is 0, and τ is the relaxation time. At this time, the time change of the output is expressed by Equation (8). If Equation (9) is set, Equation (10) is derived from Equations (7) and (8).

The expression on the right side of Expression (10) is shown in FIG. In conventional enzyme immunoassay, because practically a sufficient quantitative accuracy of about 80%, practically it is sufficient determines that the [Delta] x reaches 80% of the [Delta] x _T. The [Delta] x reaches 80% of the [Delta] x _T is expressed by equation (11), this time, an expression (12). When formula (13) is used in formula (12), the above formula (5) is obtained.

Also, based on the following equation (14) obtained from equation (10) and equation (13), find u and substitute it into the following equation (15) obtained from equations (6) and (7) to obtain Δx _T You can also.

  Instead of steps 15 to 17 in FIG. 7A, the inclination may be determined based on a plurality of recordings such as steps 21 to 24 in FIG. 7B. In this case, it is preferable to perform fitting by a least square method or the like using a polynomial or an exponential function for determining the inclination. The output in step 19 may use averaging or fitting based on a plurality of recordings.

  FIG. 9 is a diagram showing a structure example of an insulated gate field effect transistor according to another embodiment of the present invention. FIGS. 9A and 9B show a cross-sectional structure and a planar structure, respectively. In the insulated gate field effect transistor 81, a source 82, a drain 83, and a gate insulator 84 are formed on the surface of a silicon substrate, and a conductive electrode 85 is provided. A conductive electrode 85 for fixing the detection probe and a gate 86 of the insulated gate field effect transistor are connected by a conductive wiring 87. By adopting this structure, the conductive electrode 85 for fixing the probe can be formed at any location and in any size. Depending on the measurement object, it is easy to increase the probe-immobilized electrode area in order to improve the measurement sensitivity. In addition, when manufacturing various sensor chips with different measurement targets, it is not necessary to manufacture them individually. Instead, using a conventional semiconductor process, the parts other than the probe-immobilized electrode are manufactured in common, and finally the measurement target is matched. Thus, the probe-immobilized electrode and the measurement object can be immobilized, and the manufacturing cost can be greatly reduced.

  FIGS. 10A and 10B are diagrams showing the effect of removing the influence of the electric double layer on the conductive electrode 5 by applying an AC voltage to the reference electrode 7 shown in FIG. Measurements of transistor current / voltage characteristics, impedance, capacitance, etc. were performed using a semiconductor parameter analyzer (Agilent 4155C Semiconductor Parameter Analyzer) and an impedance analyzer (Agilent 4294A Precision Impedance Analyzer). Evaluation in solution was performed using a reference electrode (Ag / AgCl reference electrode) on the gate side. The AC voltage was applied to the gate at a center voltage of 50 mV and an amplitude voltage of 50 mV. FIG. 10B is an enlarged view of the dotted circle in FIG.

  The conductive electrode 5 is used as a floating gate. In the solution, an electric double layer is formed on the surface of the conductive electrode 5, which affects the change in the electric characteristics of the insulated gate field effect transistor 1 and becomes a large background noise. This is particularly noticeable when a noble metal such as gold or silver is used as the conductive electrode 5. In this embodiment, gold is used as the conductive electrode 5, and an alternating voltage is applied to the reference electrode 7 in order to remove the influence of the electric double layer. As shown in FIG. 10 (b), it can be understood that the drain current value (ID) decreases when an AC voltage is applied, compared with the case where a direct current (DC) is applied, and the effect of removing the influence of the electric double layer is obtained. . At this time, it can be seen that by increasing the frequency of the applied AC voltage, the drain current value (ID) decreases monotonously and the effect of applying the AC voltage is great.

  The size of the electric double layer on the conductive electrode is proportional to the size of the electric capacity. FIG. 11 shows the applied voltage frequency dependence of the electric capacitance of the electric double layer on the gold electrode surface actually measured. FIG. 11A shows the capacitance of the insulated gate field effect transistor, which is almost constant without frequency dependency. On the other hand, FIG. 11B shows the value in the solution, that is, the sum of the electric capacity of the insulated gate field effect transistor alone and the electric capacity of the electric double layer on the gold electrode surface. In this case, since the electrical equivalent circuit can be regarded as a capacitor, the reciprocal of the measured value represents the sum of the reciprocal of each value. As shown in FIG. 11B, by applying an alternating voltage, the electric capacity approached the value of the original electric capacity of the insulated gate field effect transistor alone and became substantially the same at 100 kHz or more. That is, it is shown that if an alternating voltage of 100 kHz or higher is applied, the influence of the electric double layer on the gold electrode surface can be almost completely eliminated.

The effect of applying the alternating voltage will be described using another embodiment. FIGS. 12A and 12B are diagrams showing changes over time from the start of measurement until the drain current is stabilized. On the surface of the gold electrode of the insulated gate field effect transistor used, a 21-base single-stranded DNA (5'-HS- (CH ₂ ) ₆ -TACGC CACCA GCTCC AACTA C-3 'is complementary to the k-ras coden12 gene. Is fixed by a bond between thiol and gold via six carbon chains. FIG. 12A shows a change over time in drain current when a positive potential is applied, and FIG. 12B shows a drain current when a negative potential is applied. In normal measurement, a direct current is used to apply a voltage to the reference electrode. However, when a sample solution is introduced into an insulated gate field effect transistor, the potential on the gold electrode surface changes and takes more than 30 minutes to stabilize. . However, as shown in FIGS. 12A and 12B, in either case of applying a positive potential or applying a negative potential, increasing the frequency of the applied voltage shortens the time until the drain current value is stabilized. I understand.

  The result is shown in FIG. 13 as the relationship between the frequency and the time until stabilization. As shown in FIG. 13, the time until stabilization is almost constant when the frequency of the applied voltage is 1 kHz or more in both cases of positive potential application (indicated by ●) and negative potential application (indicated by ○). It was. In the case of a low frequency applied voltage, the stabilization time of the negative potential voltage application is shorter than that of the positive potential voltage application because the DNA is negatively charged and repels the gold electrode surface, that is, the DNA fragment is It seems to be caused by standing up.

  The effect of the AC voltage application of the present invention will be described using another embodiment. In general, it is known that a compound having a thiol reacts with a gold surface to form an Au—S bond to form self-assembled monolayers (SAMs) having high density and high orientation. Using this property, the surface state can be easily changed by an alkyl group, a terminal functional group, a hydrophilic group of the main chain, or the like. For example, as shown in FIG. 14, if an amino group is used as the functional group at the end of the alkanethiol 91, the surface of the gold electrode 92 becomes a positive charge 93, and a carboxyl group is used as the functional group at the end of the alkanethiol 94. The surface of the gold electrode 95 becomes a negative charge 96. Taking advantage of this property, a sample in which the surface charge state on the gold electrode of the transistor of the present invention was changed was produced, and the effect of applying an alternating voltage was examined. Samples with different charge states on the surface of the gold electrode were alkanethiol having different terminal functional groups; amino group (11-amino-1-undecanethiol; 11-AUT), hydroxyl group (11-hydroxy-1-undecanethiol; 11- HUT) and a carboxyl group (10-carboxy-1-decanethiol; 10-CDT) were used. For immobilization on the gold electrode, the gold electrode was immersed in an alkanethiol ethanol solution for about 1 hour, and then washed with ethanol and pure water before use.

  The results of measuring the difference in charge state are shown in FIGS. 15 (a) and 15 (b). Also in this experiment, as described with reference to FIGS. 12A and 12B, when a DC voltage was applied to the reference electrode, it took 1 hour or more until the drain current value was stabilized. Therefore, when DC voltage is applied to the reference electrode, the data is 1 hour after immersion in the measurement solution, and when 1 MHz AC voltage is applied, the data is 5 minutes after immersion in the measurement solution. Immobilization of alkanethiol reduced the drain current compared to the untreated gold electrode (indicated as bare in the figure). Furthermore, reflecting the difference in the terminal functional group, the drain current became easier to flow in the order of amino group (positive charge; +1), hydroxyl group (neutral charge; ± 0), carboxyl group (negative charge; -1). . That is, if a positive charge is present on the gold electrode surface, the drain current tends to flow. On the other hand, if a negative charge is present on the gold electrode surface, it becomes difficult for the drain current to flow (FIG. 15A). This tendency well represents the characteristics of the FET sensor, indicating that the sensor is operating normally.

As shown in FIG. 15B, when an AC voltage (1 MHz) is applied to the gate (that is, the reference electrode), the entire drain current is reduced, and the difference in the drain current depending on the functional group at the end is also increased. It can be seen that there is an effect of applying an alternating voltage. Using the area of the gold electrode used in this experiment (0.16 mm ² ; 0.4 × 0.4 mm) and the density of alkanethiol on the gold electrode (4 / nm ² ), about 1 pmol of molecules were used in this measurement. That is, the difference in charge between the two can be detected. The density of alkanethiol on the gold electrode was measured by voltammetry under strong alkali conditions.

The effect of the AC voltage application of the present invention will be described using another embodiment. FIG. 16A is a diagram showing a change with time in drain current when an AC voltage is applied to the reference electrode 7 shown in FIG. 2 and the sample solution 12 is introduced into the measurement solution 9. FIG. 16 (b) is an enlarged view of the sample solution before and after introduction of FIG. 16 (a). As the reference electrode 7, an Ag / AgCl reference electrode was used. The AC voltage was applied to the reference electrode 7 at a center voltage of 100 mV and an amplitude voltage of 100 mV. An untreated gold electrode was used as the conductive electrode 5 of the insulated gate field effect transistor 1 at the start of measurement. As the measurement solution 9, 1.9 ml of 0.1M Na ₂ SO ₄ aqueous solution was used. As the sample solution 12, 0.1 ml of 1 mM 6-HHT (6-hydroxy-1-hexanethiol; HS- (CH ₂ ) ₆ —OH) aqueous solution was used as the alkanethiol aqueous solution.

  600 seconds after the start of measurement, the entire amount of the sample solution 12 was introduced into the measurement solution 9. At any frequency, a decrease in drain current value was observed after the sample solution 12 was introduced. When the frequency was 1 KHz or less, the drain current value increased again after decreasing, and it took more than 10 minutes to stabilize. 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. This is because the surface potential of the conductive electrode 5 becomes unstable due to the introduction of the sample solution 12 when the frequency is 1 KHz or less, but the sample solution 12 is introduced when the frequency is 10 KHz or more. This is probably because the disturbance of the surface potential of the conductive electrode 5 is eliminated. Due to the effect of stabilizing the surface potential by applying this alternating voltage, the process of the reaction occurring on the surface of the conductive electrode 5 can be measured by setting the frequency to 10 KHz or more.

The decrease in the drain current value due to the introduction of the sample solution 12 can be explained with reference to FIG. In FIG. 14, the current value of the gold electrode in which 11-HUT is fixed is smaller than the current value of the untreated gold electrode. Similarly, the current value of the gold electrode to which 6-HHT is fixed is smaller than the current value of the untreated gold electrode. That is, since 6-HHT was immobilized on the conductive electrode 5 by introducing the sample solution 12, the drain current value decreased. This immobilization is expressed by the following chemical reaction.
Au + HS- (CH ₂ ) ₆ -OH → Au-S- (CH ₂ ) ₆ -OH + 1 / 2H ₂

FIG. 17 is a diagram showing that the reaction rate is different when the concentration of the alkanethiol aqueous solution used for the sample solution 12 is different. As the alkanethiol aqueous solution, 0.1 ml of 0.05 mM 6-HHT aqueous solution and 0.1 ml of 1 mM 6-HHT aqueous solution were used. The AC voltage was applied to the reference electrode 7 at a center voltage of 100 mV, an amplitude voltage of 100 mV, and a frequency of 1 MHz. An untreated gold electrode was used as the conductive electrode 5 of the insulated gate field effect transistor 1 at the start of measurement. As the measurement solution 9, 1.9 ml of 0.1M Na ₂ SO ₄ aqueous solution was used.

  In the case of 1 mM, the reaction was completed within 1 minute, but in the case of 0.05 mM, it took 10 minutes or more to complete the reaction. Thus, it was able to measure that the reaction rate was different depending on the concentration because the measurement was performed before and after the introduction of the sample solution, and an AC voltage was applied. Conventionally, it was necessary to estimate the time until the reaction was completed or to examine it separately, but it was changed by the reaction by measuring before and after introducing the sample solution and judging whether the drain current value was stable. The difference in drain current value can be obtained quickly and reliably.

  A protein detection method using the surface state detection element will be described below. In the configuration of FIG. 2, a gold electrode is used for the conductive electrode 5, a protein detection probe is used for the biomolecule detection probe 6, an Ag / AgCl reference electrode is used for the reference electrode 7, and a central voltage is applied to the AC voltage applied to the reference electrode 7. 50 mV, amplitude voltage 50 mV, frequency 1 MHz, 450 μL of 0.1 M phosphate buffer (pH 7) was used for the measurement solution 9, and 50 μL of 0.1 M phosphate buffer (pH 7) containing 750 nM thrombin was used for the sample solution 12.

As a protein detection probe, a single-stranded DNA (aptamer DNA; HS- (CH ₂ ) ₆ -5′-TTTTT TTTTT GGTTG GTGTG GTTGG-3 ′) that specifically binds to a kind of protein thrombin was used. When immobilizing the protein detection probe, the immobilization density of the protein detection probe was controlled by using a mixed solution with alkanethiol. As the alkanethiol, 6-hydroxy-1-hexanethiol was used, the concentration ratio of the protein detection probe to the alkanethiol was 1:10, and the density of the protein detection probe was controlled to 2.6 × 10 ¹² cells / cm ² . Assuming that the protein detection probes are immobilized in a grid, the protein detection probe spacing is 6.2 nm, which is greater than the thrombin diameter of about 5 nm. The density of the protein detection probe on the gold electrode was measured by voltammetry under strong alkaline conditions.

  The measurement results are shown in FIG. The drain current value decreased about 30 seconds after the introduction of the sample solution. The delay of about 30 seconds is due to the time required for the diffusion of the introduced sample solution. The decrease in the drain current value is due to a change in the potential of the gold thin film surface due to the binding of the protein detection probe and thrombin. That is, by measuring the time change of the drain current due to sample solution injection, the binding between the protein detection probe and the protein to be measured (thrombin) in the sample solution can be measured without removing unbound protein by a washing operation. did it. As a result, it was clarified that the protein to be measured was contained in the sample solution without removing unbound protein by washing operation.

  In addition to this example, various antibodies and antigens can be measured without washing operations by immobilizing antibodies, antigens, or portions of antibodies (for example, Fab ′) as protein detection probes.

FIG. 19 is a diagram showing another configuration example of the biomolecule detection element according to the present invention. The biomolecule detection element of the present example is an element in which a sample measurement electrode and a reference electrode are mixedly mounted on the same element. The element of this embodiment has a structure in which a sample measuring electrode 111, a reference electrode 112, and a temperature measuring diode 113 are mixedly mounted. The sample measurement electrode 111 and the reference electrode 112 of this element are connected to the gates 114 and 115 of the insulated gate field effect transistor and the conductive wirings 116 and 117, respectively. That is, this element has an extended gate type structure. The insulating layer is made of SiO ₂ (thickness 17.5 nm), and a gold electrode (400 μm × 400 μm) is used as an electrode. Since this element also uses an aqueous solution, the threshold voltage of the FET is set to around -0.5V. An n ⁺ / p junction type was used as the temperature measurement diode mixedly mounted on this element.

FIG. 20 is a diagram showing another configuration example of the biomolecule detection element according to the present invention. The operation principle of the element used in the present invention is that the current between the source and the drain is caused by the potential change of the gate surface that occurs when the measurement object is bonded to the probe fixed to the surface of the gate or floating gate (that is, the conductive electrode). Is based on the principle of changing. The drain current _{ID at} that time is expressed by the following equation.

Here, W is the channel width, L is the channel length, μ _c is the mobility, C _G is the coupling capacitance between the gate insulating film and the gold surface, V _G is the gate voltage, V _t is the threshold voltage at which the channel is formed, V _DS is a voltage between the source and the drain.

  Therefore, in order to improve the measurement sensitivity of this element, the current change rate, that is, W / L may be increased. Conventionally, in order to improve the measurement sensitivity, the channel width is lengthened and the channel shape is a vertically long structure (for example, W / L = 100/1) in order to shorten the channel length. In this embodiment, as shown in FIG. 20, the source 121 and the drain 122 are comb-shaped, the channel between them is a zigzag shape 123, and the ratio of the length and width of the channel between the source 121 and the drain 122 is increased ( W / L = 480/1). In this structure, a comb-shaped structure is formed in a 400 × 400 μm shape, and the sensitivity is about 6 times higher than that of a conventional structure (for example, 400 × 5 μm) formed in the same area. It is. The conductive electrode 124 of this element and the gate of the insulated gate field effect transistor (the upper layer portion of the zigzag channel 123) are connected by the conductive wiring 125.

The figure which shows the structural example of the biomolecule detection element by this invention. The block diagram which shows the biomolecule measuring apparatus using the biomolecule detection element by this invention. The block diagram which shows the signal processing method by this invention. The figure which shows the measurement result of the electrical property of the insulated gate field effect transistor by this invention. The figure which shows the measuring method of the electrical property of the insulated gate field effect transistor by this invention. The figure which shows the method of converting the drain current value of the insulated gate field effect transistor by this invention into a gate voltage value. The flowchart which shows the biomolecule measuring method using the biomolecule detection element by this invention. The figure which shows the relationship of Formula (10). It is a figure which shows the structural example of the insulated gate field effect transistor by this invention, (a) is a top view, (b) is sectional drawing. (A) is a figure which shows the drain current value in each frequency at the time of applying an alternating voltage to a reference electrode, (b) is an enlarged view in the dotted-line circle in (a). It is a figure which shows the measurement result of the frequency dependence of the electric capacity of an insulated gate field effect transistor, (a) is a single-piece | unit (namely, in the air) measurement result, (b) is a figure which shows the measurement result in a solution. It is a figure which shows a time-dependent change until a drain current is stabilized after a measurement start, (a) is a figure at the time of positive potential application, (b) is a figure at the time of negative potential application. The figure which shows the relationship between time and frequency until a drain current is stabilized after a measurement start. The figure which shows the charge state control method of the gold | metal | money surface using alkanethiol. The figure which shows an example which detected the difference in the electric charge of the surface by a biomolecule detection element. The figure which shows the frequency dependence of the time-dependent change of the drain current value after sample solution introduction. The figure which shows the sample solution density | concentration dependence of the time-dependent change of the drain current value after sample solution introduction. The figure which shows the example which measured the binding process of the thrombin with respect to a probe using aptamer DNA for a biomolecule detection probe. The figure which shows the other structural example of the biomolecule detection element by this invention. The figure which shows the other structural example of the biomolecule detection element by this invention.

Explanation of symbols

  1, 51, 52, 81, 101 ... insulated gate field effect transistor, 2, 82, 102 ... source, 3, 83, 103 ... drain, 4, 84, 104 ... gate insulator, 5, 85, 105 ... conductive Electrode, 6 ... probe for biomolecule detection, 7 ... reference electrode, 8 ... measurement cell, 9 ... measurement solution, 10 ... power source, 11 ... sample injector, 12 ... sample solution, 13 ... sample injection tube, 20, 53 ... Measuring unit, 21 ... signal processing device, 22 ... data processing device, 23 ... data display device, 41 ... input signal, 42 ... sampling device, 43 ... averaging device, 44 ... output signal, 54 ... comparison unit, 55 ... feedback Circuit 86 86 Gate of insulated gate field effect transistor 87 Conductive wiring 91 94 94 Alkanethiol 92 95 Gold electrode 93 Positive charge 96 Negative charge 106 107 DC power supply

Claims (20)

A field effect transistor having a conductive electrode provided on the surface with a substance that contacts the measurement solution and reacts with the sample;
A reference electrode in contact with the measurement solution;
Means for applying an alternating voltage between the conductive electrode and the reference electrode;
And a means for measuring the output of the field effect transistor twice or more and calculating a difference.   2. The measuring apparatus according to claim 1, further comprising means for injecting a sample solution into the measuring solution.   3. The measuring apparatus according to claim 2, wherein the sample solution contains a biomolecule, a substrate, a nucleic acid, or a protein.   3. The measuring apparatus according to claim 2, wherein the output of the field effect transistor is measured before and after the sample solution injection.   2. The measuring apparatus according to claim 1, further comprising means for determining that the output of the field effect transistor is stable.   6. The measuring apparatus according to claim 5, wherein means for determining that the output of the field effect transistor is stable is an elapsed time t after the sample solution injection, and the field effect transistor after the sample solution injection. A measuring apparatus that makes a judgment based on whether or not st / Δx <r is established among an output change value Δx, a time differential value s of an output of the field effect transistor, and a reference value r of 0.4 or less.   2. The measuring apparatus according to claim 1, further comprising means for averaging the output using the frequency of the AC voltage. 8. The measuring apparatus according to claim 7, wherein the means for averaging the output includes a frequency f of an AC voltage, a sampling frequency f _s , an averaging count n, and any natural numbers m and l, nf = mf _s and f A measuring apparatus comprising means for sampling and averaging so that a relational expression of / f _s ≠ l holds.   2. The measuring apparatus according to claim 1, wherein the conductive electrode is connected to a gate of the field effect transistor by a conductive wiring.   2. The measuring apparatus according to claim 1, wherein the frequency of the AC voltage is 10 kHz or more.   The measuring apparatus according to claim 1, wherein the conductive electrode is made of gold.   2. The measuring apparatus according to claim 1, wherein the field effect transistor has a zigzag shape in a channel for electrically coupling a source to a drain.   The measuring apparatus according to claim 1, wherein the reference electrode is formed on the same substrate as the field effect transistor.   2. The apparatus according to claim 1, wherein a biomolecule detection probe is immobilized on the surface of the conductive electrode. The conductive electrode of a field effect transistor having a conductive electrode provided on the surface with a substance that reacts with a sample is brought into contact with a measurement solution, and an alternating current is generated between the conductive electrode and a reference electrode in contact with the measurement solution. Applying voltage, measuring the electrical characteristics of the transistor before and after the reaction between the sample and the substance on the surface of the conductive electrode, and calculating the difference between the measured value before the reaction and the measured value after the reaction Measuring method characterized by   16. The measurement method according to claim 15, wherein an AC voltage of 1 kHz or more is applied between the conductive electrode and the reference electrode before measuring at least electrical characteristics of the transistor before reaction. Method.   16. The measurement method according to claim 15, wherein an AC voltage of 10 kHz or more is applied between the conductive electrode and a reference electrode in contact with the measurement solution before measuring at least electrical characteristics of the transistor after reaction. Measuring method characterized by 16. The measurement method according to claim 15, wherein a biomolecule detection probe is immobilized on a surface of the conductive electrode.   16. The measurement method according to claim 15, wherein the measurement of the electrical characteristics of the transistor is performed a plurality of times, the elapsed time t after the sample solution injection, the change value Δx of the output of the field effect transistor after the sample solution injection, A measurement method characterized in that a measured value measured when st / Δx <r is established between a time differential value s of an output of a field effect transistor and a reference value r of 0.4 or less, and the measured value after the reaction . 18. The measurement method according to claim 17, wherein a relational expression of nf = mf _s and f / f _s ≠ l is set between the frequency f of the AC voltage, the sampling frequency f _s , the number n of averaging, and the arbitrary natural numbers m and l. A measurement method comprising sampling and averaging the measurement values so as to hold.

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