JP4482856B2 - Method for detecting target substance in sample, sensor substrate, and detection kit - Google Patents

Description

  The present invention relates to a biosensor, a method for detecting a target substance in a sample using the biosensor, and a detection kit used for the detection method.

As a technique for detecting a target substance in a test sample, a method using a biosensor in which a biomolecule such as a nucleic acid or protein or an artificially synthesized functional group is immobilized on the surface of a solid support as a probe is widely used.
In this method, a target substance is immobilized on a solid phase using specific affinity of biomolecules such as complementary nucleic acid strand interaction (hybridization), antigen-antibody reaction, enzyme-substrate reaction, receptor-ligand interaction, etc. Since it is captured on the surface of the carrier, a substance having a specific affinity for the target substance is selected as a probe.
The probe is immobilized on the surface of the solid phase carrier by a method suitable for the type of probe or solid phase carrier. Alternatively, the probe can be synthesized on the surface of the solid support (for example, nucleic acid extension reaction, etc.). In either case, the surface of the solid support on which the probe is immobilized is brought into contact with the test sample and incubated under appropriate conditions. By doing so, a probe-target substance complex is formed on the surface of the solid phase carrier.

  For the detection of the interaction, a method of labeling a target molecule in advance with a substance that is easy to detect is often used, and among them, a method of optically detecting by labeling with a fluorescent molecule is mainly used. However, it is often difficult to label in advance, such as when the amount of biomolecules contained in the test sample is small or when unknown molecules are detected. The surface plasmon resonance (SPR) method that detects the presence or absence of an interaction from the change in the rate, or the presence of an interaction by measuring the change in the resonance frequency using the crystal resonator electrode as a solid phase carrier Detection systems such as the quartz crystal microbalance (QCM) method have also been developed. However, the method that does not use a label still has problems such as non-specific binding cannot be distinguished and sensitivity is not sufficient.

  In addition, biosensors (so-called biochips and microarrays) in which probes are immobilized on the surface of a solid-phase carrier as described above are expected to be useful diagnostic tools in addition to elucidating the cause of diseases and developing pharmaceuticals. Miniaturization and integration have been promoted so that a large number of test samples can be processed in parallel with high throughput. By miniaturization, it becomes possible to measure with a very small amount of a test sample or a detection reagent, so that the measurement cost can be lowered and the reaction rate can be increased. On the other hand, in order to detect the interaction of a minute amount of substance, it is necessary to further develop a highly sensitive system that can detect a slight signal. In order to avoid contamination, it is desirable that such a biosensor can be manufactured at low cost so that it can be used in a disposable manner.

  In response to the above problems and needs, various biosensors using field effect semiconductors have been developed. A biosensor using a field-effect semiconductor is manufactured using a transistor integration technique, so that a large number of minute sensors can be formed over a substrate. Further, since not only the sensor element but also a circuit can be configured, a CPU, a memory, an amplifier circuit, and the like can be formed at the same time.

  An example of such a biosensor is a biosensor (Patent Document 2) using an ion-sensitive field-effect transistor (ISFET) (see, for example, Patent Document 1). In an ISFET type biosensor, an ion sensitive region as a probe is provided above a channel region of a thin film transistor (TFT) via a gate insulating film. When the solution comes into contact with the ion sensitive region, the interface potential changes according to the amount of ions in the solution, and the phenomenon that the amount of current flowing in the ISFET changes is utilized. According to this method, it is possible to detect biomolecular interactions that occur on the ion-sensitive membrane surface without fluorescent labeling or the like.

On the other hand, a biosensor in which an electrode is connected to a source electrode of a field effect transistor and a probe is fixed on the electrode surface has been developed. In this biosensor, the interaction of biomolecules occurring on the electrode surface functions as a switch, so that the presence or absence of the interaction can be detected by a change in the amount of current flowing in the FET.
Japanese Examined Patent Publication No. 5-49187 JP 2002-296229 A JP-A-8-313521 Japanese Patent Laid-Open No. 11-176238 JP 2002-327283 A

However, an ISFET type biosensor is difficult to detect an interaction that does not cause a large change in ion concentration. In addition, when the interaction signal is amplified, noise is also amplified, so sensitivity is limited. In the case of a biosensor in which a probe is immobilized on the electrode surface, it cannot be detected unless an electrochemical change occurs due to the interaction, and the sensitivity is limited by noise as in the ISFET type.
Therefore, the present invention can detect the interaction of various biomolecules with high sensitivity in order to solve the above problems while taking advantage of the biosensor using a field effect transistor (FET). An object of the present invention is to provide a biosensor capable of processing samples in parallel with high throughput.

  As a result of extensive research, the present inventors have caused a change in the characteristics of the FET when the probe and the target substance interact with each other at a position in direct contact with the semiconductor layer of the field effect transistor (FET). In addition, the inventors have found that the interaction between the probe and the target substance can be detected by measuring the change in the characteristics of the FET, and the present invention has been completed.

  In order to achieve the above object, a biosensor according to a first aspect of the present invention is a biosensor for detecting a target substance in a test sample, and is formed on one side of an insulating film and an insulating film. The gate electrode and the other surface side of the insulating film are fixed to the surface of the other surface side of the insulating film between the source electrode and the drain electrode, and the source electrode and the drain electrode formed with a portion facing the gate electrode therebetween. And a probe that specifically interacts with the target substance.

  In the biosensor having such a configuration, when the test sample is brought into contact with the probe fixed on the surface of the insulating film and incubated, the target substance contained in the test sample is captured on the surface of the insulating film by binding to the probe. The As will be described in detail later, when a semiconductor layer is formed between the source / drain regions so as to be in contact with the insulating film in this state, an FET is formed. In this FET, the probe and the target substance interact at a position in direct contact with the semiconductor layer, causing a change in the characteristics of the FET. In the present invention, “change in FET characteristics” refers to threshold voltage (Vth), Vg-Id characteristics, and S value [V / dec] (a gate voltage required for the drain current to change by one digit). It means change.

The FET characteristics change as the potential applied to the surface of the semiconductor layer by the voltage applied to the gate electrode changes. Therefore, when the surface potential changes due to the interaction between the probe and the target substance, this interaction can be detected by measuring the change in the characteristics of the FET.
The surface potential is easily changed by the presence of a dipole near the surface. For example, when the probe is a single-stranded nucleic acid, the probe partially forms a double strand in the molecule and does not extend linearly, and the dipole of the entire molecule is relatively small. On the other hand, when a target nucleic acid having a base sequence complementary to the probe nucleic acid is hybridized to the probe nucleic acid, the probe nucleic acid changes to a linear probe-target substance complex having a double helix structure. Is known to grow. Such a change of the dipole is not limited to the interaction between the nucleic acids, and the dipole moment is changed by the change of the conformation even in the interaction including the protein or the synthesized compound. Therefore, it is possible to detect the presence or absence of formation of the probe-target substance complex by measuring the change in the FET characteristics.

  In the biosensor of this embodiment, since the probe-target substance complex is contained in the semiconductor layer, if this complex has conductivity, a trap level is created in the band gap of the semiconductor. This also changes the FET characteristics. When a large number of trap levels exist in the band gap, phenomena such as a decrease in ON current value due to a decrease in carrier mobility and an increase in threshold voltage occur at various levels depending on the depth of the trap level. Therefore, when there is a trap level, in the Vg-Id characteristic curve of the FET with the drain voltage (Vg) on the horizontal axis and the drain current (Id) on the vertical axis, the curve shifts to the right and the level decreases. Changes occur. In addition, since each of the molecules having conductivity forms a trap level, it is possible to perform detection with a very high sensitivity of almost one molecule level.

  When the probe-target substance complex itself does not have conductivity and does not form a trap level, an appropriate label may be attached to the target substance. As the label molecule in this case, a molecule having a function of deteriorating semiconductor characteristics as an impurity is suitable. In addition, when the probe and target substance are single-stranded nucleic acids, an intercalator that forms a trap level is added after the double-stranded nucleic acid that is the probe-target substance complex is formed. The formation of double strands can be detected quantitatively. This method will be described in detail later.

  Since the change in the characteristics of the FET is obtained as a change in a plurality of parameters such as threshold voltage, drain current, and S value, it is useful for detecting various binding modes of biomolecules. By recording in the database information on each parameter change due to the combination of binding, when an unknown target substance contained in the test sample binds to the probe and changes the FET characteristics, this target is referred to the database. It becomes possible to identify the substance.

  In the FET, the drain current flows near the interface between the semiconductor layer and the insulating film. In the biosensor according to this aspect, since the probe-target substance complex is formed in the semiconductor layer in a region very close to this interface, the complex easily affects the drain current and is highly sensitive to detect changes. There is also an advantage that a sensor can be obtained.

  The biosensor according to the present invention is usually formed on a substrate. The material of the substrate is not particularly limited, and examples thereof include glass, silicon, metal (for example, gold, silver, copper, aluminum, platinum), and resin (for example, polyethylene terephthalate, polycarbonate). But you can. If an inexpensive material such as a glass substrate or a resin is used, it is suitable for disposable use.

  In addition, as described above, the biosensor according to the present invention can form a large number of sensors in an array on a single substrate using a normal transistor integration technique. Samples can be processed with high throughput.

  A biosensor according to a second aspect of the present invention is a biosensor for detecting a target substance in a test sample, and includes an insulating film, a gate electrode formed on one side of the insulating film, and an insulating film The source and drain electrodes are formed on the other side of the insulating film with a portion facing the gate electrode therebetween, and at least on the other side of the insulating film so as to face the gate electrode and be in contact with the source and drain electrodes. And a probe that specifically interacts with a target substance immobilized on the surface of the semiconductor layer in a region facing the gate electrode.

  In the biosensor having such a configuration, when the test sample is brought into contact with the surface of the semiconductor layer on which the probe is immobilized and incubated, the target substance contained in the test sample is captured on the surface of the semiconductor layer by binding to the probe. Is done. Therefore, similarly to the biosensor according to the first aspect, the probe and the target substance interact at a position in direct contact with the semiconductor layer, causing a change in the characteristics of the FET. Therefore, by measuring the change in the characteristics of the FET, the interaction between the probe and the target substance can be quantitatively detected.

  In addition, since the biosensor according to this aspect has an FET structure formed from the beginning, there is also an advantage that a post-treatment of forming a semiconductor layer after the incubation step is unnecessary.

  In the biosensor according to the present invention, it is desirable that the insulating film is laminated on the gate electrode, and the source electrode and the drain electrode are formed on the insulating film. Thus, by using a so-called bottom-gate field effect transistor, the surface of the insulating film or the surface of the semiconductor layer between the source electrode and the drain electrode (channel region) can be used as a sensor region. It becomes easy to contact the test sample.

  The semiconductor layer of the biosensor according to the present invention is preferably an organic semiconductor. By using an organic semiconductor material, it is possible to reduce the manufacturing cost and incinerate, so that a biosensor suitable for disposable use can be created. Further, since processing at room temperature is possible, a manufacturing facility such as a large-scale vacuum apparatus or a furnace for producing a single crystal is not required. If an organic semiconductor material is used using a flexible resin substrate, there is also an advantage that a durable biosensor that is hard to break can be produced.

  In particular, since the biosensor according to the first aspect forms the semiconductor layer after the probe and the target substance interact, the semiconductor layer is formed under the user of the biosensor. In such a case, if an organic semiconductor is used, a large-scale facility is not necessary, and the semiconductor layer can be easily formed by a spin coating method or a method of supplying by an inkjet discharge head or the like. In addition, the high-temperature treatment may denature the biomaterial, and there is a possibility that the binding between the probe and the target substance may not be maintained during the process of forming the semiconductor layer, but such a problem does not occur when an organic semiconductor material is used.

  As long as the probe immobilized on the biosensor according to the present invention specifically binds to the target substance, it may be a biologically derived substance or a synthesized substance, and is not particularly limited. , Sugars, organelles, microorganisms, animal and plant cells, various functional groups, and the like, and most preferably nucleic acids or proteins. On the other hand, the target substance to be detected by the biosensor according to the present invention may be a living substance or a synthesized substance as long as it specifically binds to the probe, and is not particularly limited. By selecting an appropriate probe, in addition to biomolecules, for example, harmful chemical substances contained in the environment can be detected.

  In the present invention, “specific interaction” means, for example, a binding between complementary polynucleotides, an enzyme and a substrate, an enzyme and a coenzyme, an antigen and an antibody, or a receptor and a ligand. An action that occurs only in a specific combination due to the selective affinity of a substance, and mainly means binding. Such an interaction can occur if at least one biological material is included in the interacting combination, and even if the combination includes an artificially synthesized compound, Any interaction that occurs selectively due to biological affinity is included in a specific interaction.

  The “nucleic acid” used as a probe in the biosensor of the present invention is an oligonucleotide or polynucleotide that may be partially or completely modified (including substitution) and that is further single-stranded or double-stranded. It is preferably a single-stranded oligonucleotide or polynucleotide that may be partially or completely modified (including substitution). Preferred examples of the “nucleic acid” used in the present invention include DNA, RNA, PNA (peptide nucleic acid), CNA (aminocyclohexyl ethanoic acid nucleic acid), HNA (hexitol nucleic acid), p-RNA (pyranosyl RNA), and the nucleic acid molecule A nucleic acid selected from the group consisting of oligonucleotides consisting of the above, polynucleotides consisting of the nucleic acid molecules, and the like. The nucleic acid may be extracted from a biological sample or may be artificially synthesized such as cDNA or aptamer. If a biosensor in which these nucleic acids are immobilized as probes is used, hybridization occurs when the test sample contains nucleic acids (target substances) partially or entirely complementary to the probe nucleic acids. This target substance can be captured on the biosensor surface. By this method, determination of the base sequence of the target nucleic acid, detection of mutations and polymorphisms related to various diseases, gene expression analysis, and the like can be performed.

  The nucleic acid can be immobilized on the surface of the insulating film or semiconductor layer according to a method known per se. For example, a method of adsorbing a specific group by binding a specific group with a linker sandwiched between the ends of the nucleic acid molecule is known. At this time, a self-assembled monolayer can be formed by selecting a group suitable for the material of the surface to be fixed. For example, when the insulating film is a silicon oxide film, it is possible to fix the monomolecular film of nucleic acid on the surface of the insulating film according to the method of Patent Document 5. In addition, the insulating film surface is treated with a silicon coupling agent to which various functional groups (for example, amino group, aldehyde group, epoxy group) are bonded, and a functional group that can be bonded to any of the functional groups (for example, amino group, aldehyde group) There is also a method in which a nucleic acid molecule is uniformly fixed on an insulating film by bonding a group, a thiol group, biotin, etc.) to the end of the nucleic acid molecule.

The “protein” used in the biosensor of the present invention means a protein in which at least two amino acids are covalently bonded, and examples thereof include enzymes, substrates for enzymes, antibodies, epitopes, antigens, various receptors, and ligands. By using these, the target substance can be captured on the biosensor surface using enzyme-substrate reaction, enzyme-coenzyme binding, antigen-antibody reaction, receptor-ligand reaction, and the like. The probe may be a part of the protein or a complex protein to which a sugar chain or the like is bound.
For example, the surface of a silicon oxide film is treated with a silicon coupling agent that is modified with various functional groups, and the amino terminus, carboxy terminus of the protein, residues in the middle of the polypeptide chain, and various spacers are used to form the silicon. By bonding with the functional group of the coupling agent, it can be fixed on the insulating film. Further, the surface to be immobilized can be bound via an amino group of a protein by coating with an aldehyde or polylysine. As an anchor molecule, it may be indirectly bound via protein G, protein A, biotin, streptavidin or the like.

  The above-described nucleic acids and proteins may be fixed to each sensor one by one, or multiple types of probes may be fixed to one sensor, and a biosensor designed for each detection purpose can be designed. it can.

  A first method for detecting a target substance in a test sample according to the present invention includes an insulating film, a gate electrode formed on one side of the insulating film, and the gate electrode on the other side of the insulating film. And a probe that specifically interacts with the target substance fixed on the other surface side surface of the insulating film between the source electrode and the drain electrode. And a step of bringing the sample into contact with the probe for incubation, a step of washing the surface of the insulating film on which the probe is fixed, and a source electrode and a drain electrode on the surface on which the probe is fixed Forming a semiconductor layer so as to be in contact with each other, forming a field effect transistor, and measuring a characteristic of the field effect transistor.

  In such a method, first, a probe and a target substance having specific affinity for the probe are bound by an incubation step. In the subsequent washing step, the test sample that does not bind to the probe is washed away, and only the probe-target substance complex is left on the surface of the insulating film. An FET is formed by forming a semiconductor layer on the surface of the insulating film. Since the probe-target substance complex is formed at a position directly in contact with the semiconductor layer of the TFT, the probe-target substance complex can be detected by measuring the change in the characteristics of the FET as described above.

  The second method for detecting a target substance in a test sample according to the present invention includes an insulating film, a gate electrode formed on one side of the insulating film, and the gate on the other side of the insulating film. A source electrode and a drain electrode formed with a portion facing the electrode separated from each other; and a semiconductor formed at least on the other surface side of the insulating film so as to face the gate electrode and to be in contact with the source electrode and the drain electrode And a probe that specifically interacts with the target substance immobilized on the surface of the semiconductor layer in a region facing the gate electrode. A step of bringing the sample into contact with the test sample, a step of washing the surface of the semiconductor layer on which the probe is fixed, and a step of measuring the characteristics of the field effect transistor. Including the.

  According to such a method, the probe and the target substance having specific affinity for the probe are first bound by the incubation step, and the test sample that has not bound to the probe is washed away by the subsequent washing step. Only the target substance complex is left on the surface of the semiconductor film. Since the probe-target substance complex is formed at a position in direct contact with the semiconductor layer, it changes the dipole moment on the surface of the semiconductor layer, resulting in a change in FET characteristics. Therefore, the probe-target substance complex can be detected by determining the characteristic change of the FET.

  Further, in the detection method according to the present invention, the FET characteristics measured by the first or second method according to the present invention described above, and the field effect transistor characteristics measured when the probe does not interact with the target substance, It is preferable to further include a step of comparing. According to such a method, it is possible to quantify the change in characteristics of the FET between the state in which the target substance is not bound to the probe and the state in which the target substance is bound, and quantitatively detect the formation of the probe-target substance complex. Is possible. Specifically, the test sample not containing the target substance is brought into contact with the probe, the other steps are performed in the same manner, and then the difference in the measurement results is taken to determine the change in FET characteristics. By doing this, the influence of the buffer used for the test sample and the characteristic change that occurs when the intercalator is adsorbed to the single-stranded nucleic acid by electrostatic action when using the intercalator are canceled as the background, It is possible to determine the change in characteristics of the FET only by forming the probe-target substance.

In the detection method according to the present invention, the semiconductor layer is preferably an organic semiconductor layer. The advantages of using the organic semiconductor layer are as described above. Examples of organic semiconductor materials include low molecular compounds such as naphthalene, anthracene, tetracene, pentacene, hexacene, oxydiazole derivatives (PBD), oxydiazole dimers (OXD-8), beryllium-benzoquinolinol complexes (Bebq), tri Phenylamine derivatives (MTDATA) and triarylamine derivatives, triazole derivatives, polyphenylene, polyalkylfluorene, polyalkylthiophene (P ₃ HT), polyvinylpyrene, polyvinylanthracene, F8T2 (poly (9,9-dioctylfluorene-co-bithiophene) ) And the like.

  Preferably, a polymer organic semiconductor material such as F8T2 is dissolved in a nonpolar organic solvent such as xylene, toluene, or trimethylbenzene, and supplied by an ink jet type ejection head. Low molecular organic semiconductor materials such as tetracene and pentacene can also be formed by, for example, vapor deposition. Also in this case, since it is not necessary to heat the whole biosensor, it is also suitable for biomaterials.

  In the detection method according to the present invention, as described above, the target substance can be labeled in advance. For the labeling, it is preferable to use a conductive substance that forms a trap level. However, when the probe-target substance complex is formed, the dipole moment changes greatly compared to the case of the probe alone. Any material can be used. By such labeling, the interaction between the probe and the target substance greatly changes the characteristics of the field effect transistor, and can be used as a highly sensitive biosensor.

  When a single-stranded nucleic acid is used as a probe and the nucleic acid in the test sample is captured by hybridization, an intercalator can be added and detected after the incubation step. As the conductive intercalator, for example, an intercalator imparted with surface activity disclosed in Japanese Patent Laid-Open No. 8-313521 (Patent Document 3) or Japanese Patent Laid-Open No. 11-176238 (Patent Document 4) may be used. it can. The intercalator is preferably conductive, but any substance that can greatly change the dipole moment of the double-stranded nucleic acid as compared to that of the single-stranded nucleic acid may be used as in the above-described label. Thereby, the change of the dipole moment due to the formation of the double-stranded nucleic acid can be increased, and the interaction detection of a small amount of sample is also possible.

  The present invention relates to an insulating film, a gate electrode formed on one side of the insulating film, a source electrode and a drain electrode formed on the other side of the insulating film with a portion facing the gate electrode, and a source electrode A target substance in a test sample including a biosensor including a probe that specifically interacts with a target substance immobilized on the other surface side surface of the insulating film between the drain electrode and the drain electrode; and an organic semiconductor material. A detection kit for detection is also included. If such a kit is used, the organic semiconductor layer can be easily formed under the user's condition.

The best mode for carrying out the present invention will be described below with reference to the drawings. However, the present invention is not limited to the following specific examples. Those skilled in the art can implement the present invention by making various modifications to the examples shown below, and such modifications are included in the scope of the claims of the present application.
<First embodiment>
FIG. 1 shows a cross-sectional view of a biosensor 100 according to the first embodiment. In this example, a case of a biosensor for detecting a nucleic acid in which a single-stranded DNA is immobilized in a channel region will be described as an example.
In the biosensor 100, a gate electrode 12 made of a gold thin film is formed on a glass substrate 10, and a silicon oxide film 14 is formed on the gate electrode 12 as a gate insulating film. On the insulating film, the source electrode 16 and the drain electrode 18 are formed of a gold thin film with the upper portion of the gate electrode 12 therebetween, whereby the surface of the insulating film between the two electrodes, that is, the upper portion of the gate electrode 12 is the channel region 20. It becomes. Single-stranded DNA having a sequence complementary to the target nucleic acid is immobilized on the channel region 20 as a probe.
(Production method)
Next, a method for manufacturing the biosensor according to the present embodiment will be described. FIG. 2 is a cross-sectional view of the manufacturing process of the biosensor 100. As shown in FIG. 2A, a gold thin film is formed to a thickness of about 20 μm on a glass substrate 10 by, eg, vapor deposition, and then patterned to form a gate electrode 12. Tantalum, aluminum, platinum, chromium, or the like may be used for the gate electrode.
Subsequently, as shown in FIG. 2B, a silicon oxide film 14 is formed on the glass substrate 10 and the gate electrode 12 so as to have a film thickness of about 500 nm by, for example, PECVD. The silicon oxide film 14 functions as a gate insulating film. As shown in FIG. 2B, the gate insulating film may be planarized by an etch back method.
Further, as shown in FIG. 3C, the gold thin film 28 is formed to a thickness of about 20 μm by, for example, sputtering. Then, as shown in FIG. 2D, patterning is performed to form a channel region 20 above the gate electrode. The size of the channel region can be appropriately changed according to the type of the target substance, but is preferably about 5 μm to 20 μm square.
Next, as shown in FIG. 5E, the single-stranded DNA 22 as a probe is fixed to the surface of the insulating film 14 in the channel region 20. DNA can be fixed in accordance with the method described above. For example, a functional group that forms a SAM film on a silicon oxide film is bonded to the end of DNA, and a solution containing the functional group is supplied to the channel region by an ink jet discharge head.
(Detection method)
Next, a detection method using this biosensor will be described with reference to FIGS. 3 (a) to 1 (d). First, for example, a sample containing a target substance is diluted with a PBS buffer (50 mM KPO ₄ , 5 mM EDTA, 1M NaCl, pH 7.0) or the like so that the nucleic acid concentration becomes 1 μM, for example, and used as a test sample.
Next, as shown in FIG. 3B, the test sample 24 is discharged into the channel region 20 by an ink jet discharge head, and incubated to perform a hybridization reaction. The nucleic acid in the target sample having a sequence complementary to the base sequence of the probe nucleic acid 22 hybridizes with the probe nucleic acid 22 to form a double strand. The composition of the buffer solution is not particularly limited, and the stringency can be adjusted by changing the salt concentration (ionic strength) or temperature.
Subsequently, a surface active dye intercalator 25 of the following formula disclosed in JP-A-8-313521 is added to the channel region 20 as a conductive intercalator. As shown in FIG. 3C, the intercalator 25 selectively inserts and binds to the portion where the double-stranded nucleic acid is formed. As the conductive intercalator, in addition to methylene blue, ethidium bromide, actinomycin D, acridine dye, etc., or those provided with surface activity or conductive functional group may be used.
Next, the channel region 20 is washed, and a test sample containing nucleic acid or the like not bound to the probe is removed from the channel region. For washing, a solution obtained by appropriately diluting an SSC (saline-sodium citrate) solution with ultrapure water and adding SDS (Sodium dodecyl sulfate) as necessary can be used.
As shown in FIG. 3D, F8T2 (poly (9,9-dioctylfluorene-co-bithiophene)) was dissolved in xylene as the organic semiconductor layer 26 on the channel region 20, the source electrode 16, and the drain electrode 18. These are stacked to form a field effect transistor. The organic semiconductor layer can be supplied by an inkjet discharge head. If discharged by an ink jet discharge head, the double-stranded nucleic acid and the intercalator can be effectively contained in the semiconductor layer. In addition to the inkjet method, the organic semiconductor layer may be formed by vapor deposition, spin coating, LB, or the like. Note that the semiconductor layer can also be formed using a water-soluble polymer.
Finally, with the positive voltage applied to the drain electrode, the positive gate voltage is gradually increased, the threshold voltage (Vth) at which current starts to flow between the source / drain electrodes, and the drain flowing between the source / drain electrodes The current (Id) is measured.
On the other hand, as a comparative example, only the PBS buffer used for the preparation of the test sample is added to the biosensor shown in FIG. 3A, washed through an incubation step, and the threshold voltage and the drain current are similarly measured. .
An example of the Vg-Id characteristic curve thus obtained is shown in FIG. In the figure, a solid line indicates a Vg-Id characteristic curve in the case where a double-stranded nucleic acid is formed and an intercalator is bound to this double-strand. A dotted line indicates a comparison. It is an example characteristic curve. In the comparative example, since the organic semiconductor layer includes only a single-stranded nucleic acid having an insulating property and a small dipole moment, a sharply rising curve close to the Vg-Id characteristic of the intrinsic semiconductor is drawn.
In contrast, in an embodiment in which an intercalator is included in the organic semiconductor layer, a double strand is formed and the dipole moment of the probe-target substance is increased, and the intercalator functions as a trap. The absolute value of Vth) increases and the drain current (Id) decreases. In addition, the slope of the curve becomes gentle and the S value becomes large.
As described above, according to the biosensor according to the present embodiment, the binding between the probe and the target substance can be measured with high sensitivity and quantitatively using a plurality of parameters, and various biomolecule interactions can be detected. Useful. In addition, a large number of biosensors can be formed on one substrate, and high-throughput measurement is possible. Furthermore, since the manufacturing cost is relatively low and incineration can be performed, it is suitable for disposable use and can prevent contamination between test samples.
<Second Embodiment>
FIG. 5 shows a cross-sectional view of a biosensor 200 according to the second embodiment. In this embodiment, an example of a biosensor for antigen detection in which an antibody is immobilized on the surface of a semiconductor layer above the channel region will be described.
A gate electrode 52 is formed of a gold thin film on the glass substrate 50, and a silicon oxide film is formed on the gate electrode 52 as a gate insulating film 54. On the insulating film 54, a source electrode 56 and a drain electrode 58 are formed of a gold thin film across the upper portion of the gate electrode 52, whereby the surface of the semiconductor thin film between the two electrodes, that is, the upper portion of the gate electrode 52 is channeled. It becomes area 60. A semiconductor layer 66 is formed on the source electrode 56, the drain electrode 58, and the channel region 60, and an antibody having affinity for the target substance is immobilized as a probe on the surface of the semiconductor layer 66 above the gate electrode 52. ing.
(Production method)
The biosensor according to the present embodiment can be manufactured by manufacturing a field effect transistor according to a normal manufacturing method and immobilizing an antibody on the surface of the semiconductor film. First, the gate electrode 52, the insulating film 54, the source electrode 56, and the drain electrode 58 are formed on the substrate 50 by the same method as the biosensor according to the first embodiment, and the source electrode, the drain electrode, and the channel region are formed. Then, an organic semiconductor layer 66 is formed. The organic semiconductor layer 66 can be formed by vapor deposition using, for example, F8T2 dissolved in xylene. In order to obtain sufficient sensitivity, the organic semiconductor layer 66 is preferably formed as thin as possible, and may be formed to a thickness of 5 nm to 50 nm. By forming the organic semiconductor thin, a recess is formed in the upper part of the channel region 60, so that this recess can be used as a sensing unit.
Next, the antibody 62 is fixed to the surface of the organic semiconductor layer in the recess. The antibody can be immobilized in accordance with the above-described protein immobilization method. For example, the surface of the semiconductor layer 66 can be coated with an aldehyde to be bound via the amino group of the antibody. In any of the methods, fixation is performed so that the binding ability is not impaired by loss of steric hindrance or higher order structure, and the orientation for binding to the target substance is maintained.
(Detection method)
Next, a method for detecting a target substance using the biosensor according to the present embodiment will be described with reference to FIG. A test sample 64 that is considered to contain the target substance is added after being diluted so that the protein concentration becomes 0.1 to 0.5 mg / ml, for example. After the incubation, the sample is washed with PBS, and the test sample 64 containing the unbound substance is washed away. Due to the binding between the antibody 62 and the target substance, the dipole moment of the whole molecule changes, and the surface potential of the semiconductor film 66 changes, thereby changing the FET characteristics. By measuring the characteristics of the biosensor 200 as an FET in advance, the formation of the complex can be quantitatively measured.
When changes in characteristics that differ depending on the target substance are expressed by a plurality of parameters such as threshold voltage, Vg-Id characteristics, and S value, and a database in which each value is recorded is created, when an unknown substance binds to the probe, It is also possible to identify this substance with reference to a database.
As described above, when the biosensor according to the present embodiment is used, the process of forming the semiconductor layer after the incubation is unnecessary, and the process of bringing the test sample into contact with the surface of the semiconductor layer, the incubation process, and the cleaning process can be easily performed. The probe-target substance complex can be detected. In addition, similar to the biosensor according to the first embodiment, it is possible to measure interaction with high sensitivity, quantitative and high throughput, and it is suitable for disposable use.

FIG. 1 is a cross-sectional view of the biosensor according to the first embodiment. FIG. 2 is a manufacturing process cross-sectional view of the biosensor according to the first embodiment. FIG. 3 is a cross-sectional view showing a target substance detection method using the biosensor according to the first embodiment. FIG. 4 shows an example of a Vg-Id curve of a field-effect transistor that changes due to the probe-test substance interaction. FIG. 5 is a cross-sectional view of the biosensor according to the second embodiment. FIG. 6 is a cross-sectional view showing a method for detecting a target substance using the biosensor according to the second embodiment.

Explanation of symbols

DESCRIPTION OF SYMBOLS 10, 50 ... Glass substrate, 12, 52 ... Gate electrode, 14, 54 ... Silicon oxide film, 16, 56 ... Source electrode, 18, 58 ... Drain electrode, 20, 60 ... Channel region, 22 ... Probe nucleic acid, 24 ... Test sample, 25 ... intercalator, 26, 66 ... organic semiconductor layer, 28 ... gold thin film, 62 ... antibody, 100 ... biosensor, 200 ... biosensor

Claims (9)

A method for detecting a target substance in a test sample,
An insulating film; a gate electrode formed on one surface of the insulating film; a source electrode and a drain electrode formed on the other surface of the insulating film with a portion facing the gate electrode; and the source electrode and using a sensor substrate comprising, a probe in which the interact specifically with the target substance fixed on the other side surface of the insulating film between said drain electrode each other,
Incubating the probe with a test sample; and
Cleaning the surface of the insulating film on which the probe is fixed;
Forming a field effect transistor by forming a semiconductor layer on the surface on which the probe is fixed so as to be in contact with the source electrode and the drain electrode;
Measuring the characteristics of the field effect transistor;
Including methods. A method for detecting a substance contained in a sample,
A gate electrode, a source electrode, a drain electrode, an insulator in contact with the gate electrode, and a probe provided on a surface of the insulator on the side facing the gate electrode; Allowing the probe and the substance to interact to form a first substance;
Removing the sample not interacting with the probe on the surface of the sensor base on the side of the insulator facing the gate electrode;
Forming a semiconductor so as to be in contact with the source electrode and the drain electrode and encapsulate the first substance;
Measuring characteristics of a field effect transistor including the gate electrode, the source electrode, the drain electrode, the insulator, and the semiconductor;
Including methods.   2. The method of claim 1, further comprising the step of comparing the field effect transistor characteristics measured by the method of claim 1 with the field effect transistor characteristics measured when the probe does not interact with the target substance. the method of.   The method of claim 1, wherein the semiconductor layer is an organic semiconductor layer.   The method according to claim 1, wherein a label molecule is bound to the target substance before the incubating step.   The method according to claim 1, further comprising the step of adding a conductive intercalator after the incubation step, wherein the target substance and the probe are single-stranded nucleic acids.   The method according to claim 1, wherein the characteristic of the field effect transistor is one of a threshold voltage, a Vg-Id characteristic, and an S value. A sensor substrate used in the method according to claim 1 ,
A gate electrode;
A source electrode;
A drain electrode;
An insulator in contact with the gate electrode;
The probe provided on the surface of the insulator facing the gate electrode;
A sensor substrate. A detection kit for detecting a target substance in a sample, comprising the sensor substrate according to claim 8 .