JP2008209143A - Measuring method of creatinine and system for it - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a new quantitatively measuring method of creatinine in a sample, and to provide a system for the same. <P>SOLUTION: The measuring method of creatinine includes; a process of reacting creatinine in a specimen with an enzyme reactive with creatinine; a process of detecting change in ion concentration generated by the reaction using an accumulator type ISFET sensor; and a process of quantitatively measuring creatinine in the specimen based on the detected change in ion concentration. A system for the same is also disclosed. <P>COPYRIGHT: (C)2008,JPO&INPIT

Description

  The present invention relates to a novel method for quantifying creatinine in a sample and a system therefor.

  Quantification of blood and urine creatinine has been utilized in the diagnosis of kidney disease, particularly renal dysfunction. It is also used as an indicator of dialysis transfer.

  Conventionally, as a method for quantifying creatinine in body fluids, a method of coloring with alkali picric acid has been widely used, but the instability of the test solution in the analytical instrument and the adverse effect on the analytical instrument due to the alkaline reagent However, there are drawbacks such as measuring the polysaccharide and applying it to accuracy, and lacking in analytical precision. As a method for solving these problems, a method using an enzyme has been widely used in recent years. Specifically, an enzyme that specifically acts on creatinine (preferred example is creatininase [EC3.5.2.10]) and a series of following enzyme groups that generate hydrogen peroxide therefrom (preferred example) As a result, hydrogen peroxide derived by the action of creatinase [EC 3.5.3.3], sarcosine oxidase [EC 1.5.3.1], peroxidase [EC 1.11.1. It is detected by a dye group and attempts to quantify creatinine in a sample. The method using an enzyme is superior in specificity without being affected by saccharides present in a large amount in the blood like the alkali picric acid method, is not easily affected by other interfering substances, and is stable in an analytical instrument. There are advantages such as.

  Therefore, at present, a method for measuring creatinine by a continuous enzyme reaction system of creatininase-creatinase-sarcosine oxidase-peroxidase is routinely carried out as a mainstream in clinical examination (for example, Patent Documents 1 to 5). reference).

  However, this method requires a large number of enzymes and requires high specificity of each enzyme for the substance to be measured. When any enzyme reacts with a substance other than the target substance, that is, when the specificity is low, they are measured and appear as a positive error. When a large number of enzymes are used, if any of the enzymes is affected by the substance in the sample, the whole reaction system may be affected and accurate measurement may not be performed.

  In addition, since the optimum use conditions of the enzymes differ depending on the individual enzymes, in order to react several enzymes in the same reaction solution, the use of each enzyme under the optimum conditions must be given up.

  Furthermore, the method using several enzymes has a limit in reducing the cost of the reagent.

  Therefore, it has been desired that creatinine can be measured with fewer enzyme species. As one of such attempts, a method using another enzyme group has been developed. This is because creatinine deiminase [EC3.5.4.21] acts on creatinine, and the resulting ammonia acts on glutamate dehydrogenase [EC1.4.1.2], whereby nicotine corresponding to the amount of ammonia. In this method, the enzymatic conversion of amidoadenine dinucleotide from a reduced form to an oxidized form is measured by a decrease in ultraviolet absorption. Since only two types of enzymes are used in this measurement method, creatinine can be measured with fewer enzyme species than currently popular enzymatic measurement methods (see, for example, Patent Documents 6 and 7).

  However, since this method is measured by a decrease in absorbance, there is a major drawback that it is inherently insensitive. For this reason, this method is not widely used in practical use. For example, while the normal value of human serum creatinine is about 1 mg / dL, the detection sensitivity of the conventional method for measuring creatinine using creatinine deiminase was as low as about 50 mg / dL (for example, Patent Document 8, 9).

  In order to meet the desire to measure creatinine with fewer enzyme species, methods that can measure creatinine with a single enzyme, for example, research and development of an enzyme that oxidizes and reduces creatinine should be attempted. Until now, no such enzyme or measurement method has been developed.

  At present, measurement using enzyme biosensors is actively performed in fields such as medical field, analytical chemistry, and food engineering. Such a biosensor is composed of, for example, an electrochemical sensor and an enzyme-immobilized membrane, and is used for measuring a biochemical component contained in a trace amount in a sample solution (see, for example, Patent Documents 10 to 15). As an example, an enzyme sensor of a current detection type mainly for blood glucose level measurement is widely used worldwide as a glucose sensor and is widely used. For the glucose sensor, enzymes such as glucose oxidase and glucose dehydrogenase are used. In practice, the enzyme is usually carried on a test specimen for a sensor, and blood, which is a sample, is placed on the test specimen, and a current generated by the reaction of glucose in the sample with the enzyme is detected by the sensor.

As an electrochemical sensor, a current detection type enzyme sensor is generally used, but there are cases where potential detection is performed using a field effect transistor instead of current detection. For this purpose, it is conceivable to use an ion-sensitive field effect transistor (ISFET) serving as a sensor for hydrogen ion concentration (pH). The ISFET sensor uses a mechanism in which an interface potential is generated according to the ion activity in the solution when the solution comes into contact with an ion sensitive film (SiO ₂ , Si ₃ N ₄ , Ta205, etc.) on the gate of the ISFET. . An ion concentration can be detected by forming a sensing portion sensitive to ions on the gate insulating film of the ISFET, and detecting a change amount of the potential level of the channel based on a change in potential of the surface of the sensing portion. As one of the applications, the ISFET is sensitive to hydrogen ions at the sensing unit and becomes a pH sensor. The change in the interface potential between the gate insulating film and the sample is measured as a pH-dependent output voltage. Furthermore, various potentiometric sensors can be produced by depositing various sensitive films on the gate portion. Since biosensors using ISFETs are manufactured by integrated circuit manufacturing processes, they can be miniaturized and standardized, and have the advantage of being capable of mass production, so their development potential is expected. .

  However, although ISFET is used for pH measurement, there is a problem that the sensitivity is low, the output is unstable with time, and the ion concentration cannot be detected with high accuracy. At present, ISFET is used. Biosensors are not widely used as compared with current detection type biosensors.

  In order to overcome the shortcomings of ISFETs, cumulative ISFET sensors have been developed (Patent Document 16). The cumulative ISFET sensor is configured to repeatedly transfer the change in the depth of the potential well immediately below the sensing part based on the change in the surface potential of the sensing part of the ISFET as a charge, so that the charge is accumulated in the drain. Thus, even if the change in the surface potential of the sensing unit is very small, it can be detected reliably, and the change in the ion concentration can be detected with high sensitivity. In addition, as an application of this technology, it is possible to detect the presence or absence of hybridization between a nucleic acid in a sample and a single-stranded nucleic acid with high sensitivity, and without increasing the amount of DNA by the PCR method, An FET type sensor and a base sequence detection method that can determine a base sequence in a short time and at low cost have already been invented (Patent Document 17). However, it is unclear whether such a cumulative ISFET can be used for an enzyme biosensor for measuring creatinine or the like with high sensitivity.

JP-A-6-217800 JP 2002-306197 A JP-A-10-215874 JP 2000-157279 A JP-A-11-346771 JP-A-10-136997 JP-A-7-143881 JP-A-7-327695 JP-A-7-322898 JP 10-293112 A JP 2002-350383 A JP-A-8-327587 JP 2004-294087 A JP-A-8-31478 JP-A-8-29389 Japanese Patent No. 3623728 International Publication No. 03/042683 Pamphlet

  An object of the present invention is to provide a novel method for quantifying creatinine in a sample and a system therefor. More specifically, it is to provide a method capable of measuring creatinine with a single enzyme, and a system therefor.

  As a result of diligent investigations to achieve the above object, the present inventors have found that a sensor system using a cumulative ion-sensitive field effect transistor (ISFET) is effective as a change in ion concentration. We found that enzyme reactions can be traced well. In addition, by using creatinine deiminase as an enzyme, it is possible to quantify creatinine in a sample by detecting changes in the concentration of ions generated by reaction with creatinine in the sample with a cumulative ISFET sensor. As a result, the present invention has been completed.

That is, the present invention has the following configuration.
[1] A method for measuring creatinine, comprising the following steps:
(1) a step of reacting creatinine in a sample with an enzyme that reacts with creatinine;
(2) A step of detecting a concentration change of ions generated by the reaction using a cumulative ion sensitive field effect transistor (ISFET) sensor, wherein the cumulative ISFET sensor changes according to the ion concentration acting on the sensing unit. An ISFET sensor that measures the amount of charge accumulated after resetting the potential as a potential change according to the depth of the potential well and the number of pumps from the potential well; and (3) In the sample based on the detected ion concentration change Quantifying the amount of creatinine.
[2] The measuring method of [1], wherein the ion generated by the reaction is an ammonium ion.
[3] The measuring method of [1], wherein the enzyme that reacts with creatinine is a single enzyme.
[4] The measurement method according to [3], wherein the enzyme that reacts with creatinine is creatinine deiminase [EC3.5.4.21].
[5] The measuring method of [1], wherein an enzyme that reacts with creatinine is present in a free state in the reaction mixture.
[6] The measuring method according to [1], wherein creatinine contained in the sample at a concentration of 0.1 mg / dL can be measured.
[7] The method according to [1], wherein an enzyme that reacts with creatinine is reacted with creatinine in a sample in the presence of an alkali metal salt and / or an alkaline earth metal salt.
[8] The measuring method according to [1], wherein an enzyme that reacts with creatinine is reacted with creatinine in a sample in the presence of the buffer component, and the concentration of the buffer component in the reaction mixture is 0.5 to 5 mM.
[9] The measuring method according to [1], wherein the surface area of the sensing part of the cumulative ISFET sensor is 1 mm ² or less.
[10] The measuring method of [1], wherein an enzyme that reacts with creatinine is reacted with creatinine in a sample in a reaction mixture having a volume of 30 μL or less.
[11] A system for quantifying creatinine in a sample, comprising at least the following elements (a) and (b):
(A) Cumulative ion-sensitive field effect transistor (ISFET) sensor, where the cumulative ISFET sensor determines the depth of the potential well that has changed according to the ion concentration acting on the sensing unit and the number of pumps from the potential well. Accordingly, an ISFET sensor that measures the amount of charge accumulated after potential reset as a potential change; and (b) an enzyme that reacts with creatinine.
[12] The system according to [11], wherein the enzyme that reacts with creatinine in (b) is creatinine deiminase [EC3.5.4.21].

  According to the present invention, creatinine in a sample can be quantified using only a single enzyme. In addition, it is possible to construct a system for measuring creatinine in a sample with a very simple configuration.

  The present invention is described in detail below.

The present invention provides a method for measuring creatinine, which comprises the following steps:
(1) a step of reacting creatinine in a sample with an enzyme that reacts with creatinine;
(2) A step of detecting a concentration change of ions generated by the reaction using a cumulative ion sensitive field effect transistor (ISFET) sensor, wherein the cumulative ISFET sensor changes according to the ion concentration acting on the sensing unit. An ISFET sensor that measures the amount of charge accumulated after resetting the potential as a potential change according to the depth of the potential well and the number of pumps from the potential well; and (3) In the sample based on the detected ion concentration change Quantifying the amount of creatinine.

  Creatinine is a substance also called methylglycocyanidin and is produced by dehydration and cyclization of creatine. The normal value of human serum creatinine is about 1 mg / dL, but it increases with kidney disease.

  In the method of the present invention, the sample used for creatinine measurement is not particularly limited. Examples of biological samples include blood, serum, plasma, saliva, cerebrospinal fluid, urine, feces, feces, lymph, semen, tears, and various organs such as mammals (human, cow, horse, pig, wild boar, sheep) , Rabbits, especially humans), birds (chicken, ducks, quail, turkeys, ducks, pheasants, etc.), invertebrates (insects; silkworms, bees, ants, stag beetles, beetles, crustaceans; shrimps, crabs, etc.), plants Examples include biological samples such as mulberry, red beans, broad bean, tomato, eggplant, cucumber, melon, tobacco, chrysanthemum, lily, rose, and microorganisms (bacteria, yeast, mold, etc.). Moreover, food, a river, soil, etc. are illustrated as a sample derived from an environment.

  The sample may be subjected to measurement after being crushed by ultrasonic treatment, surfactant treatment or the like. By doing so, it may be possible to improve the measurement sensitivity.

  In the present invention, the ions to be measured are not particularly limited as long as they cause changes in the hydrogen ion concentration directly or indirectly. Examples of such ions include hydrogen ions, ammonium ions, and phosphate ions. A suitable measurement object of the present invention includes ammonium ions.

  In the present invention, any enzyme that reacts with creatinine to cause the above-described ion concentration change can be used. The enzyme that reacts with creatinine may be a single enzyme or a combination of enzymes (enzyme group) as long as it reacts with creatinine to cause a change in ion concentration. Examples of such enzymes (enzyme group) include creatininase-creatinase-sarcosine oxidase enzyme group, creatininase-creatine kinase [EC 2.7.3.2] enzyme group, and the like. In one embodiment, the present invention is characterized in that creatinine in a sample can be accurately quantified only by using a single enzyme. Therefore, in order to take advantage of the present invention, it is preferable to react creatinine deiminase with creatinine. By using creatinine deiminase, the method of the present invention can detect the concentration change of the generated ions and quantify creatinine.

  Creatinine deiminase is an enzyme that catalyzes the reaction of hydrolyzing creatinine into N-methylhydantoin and ammonia. Examples of the supply source of creatinine deiminase include Bacillus bacteria (JP-A-61-219383), Corynebacterium bacteria (JP-A 52-34976), Flavobacterium bacteria (The Journal of Biological Chemistry). Vol.260 No.7 p3915-3922, 1985) is known. The creatinine deiminase used in the present invention is not limited in any way, such as its origin.

  In the present invention, the form of the enzyme or the composition containing the enzyme is not particularly limited. It may be in a liquid state or in a solid state, and water may be supplied by mixing with a sample. Placed in an appropriate container or mounted on an appropriate device, for example, an analytical reagent for molecular biology, an analytical reagent for biochemistry, an in vitro diagnostic agent, a liquid in vitro diagnostic agent, a chip or a slit Various forms such as in vitro diagnostics, biosensors, pharmaceuticals, foods and beverages can be taken.

  In the present invention, the enzyme may be immobilized on a carrier such as a membrane or may exist in a free state. Generally, a conventional biosensor using a field effect transistor (FET) is used by immobilizing and supporting an enzyme directly or indirectly on a membrane of a sensing unit. The enzyme immobilized on the membrane and the enzyme not immobilized may have different pH characteristics. This is due to the fact that the three-dimensional structure of the enzyme changes and the characteristics change due to the immobilization, but the following two points can be considered. First, a pH gradient is generated by the enzyme-immobilized membrane, and there is a difference between the pH in the solution and the pH in the enzyme-immobilized membrane. Second, the pH in the vicinity of the enzyme by an acid generated by an enzyme reaction or the like. Is changing. Therefore, in such a case, it is preferable that the enzyme exists in a free state. In addition, in the case of multi-item simultaneous measurement in which a sample is used for measurement of substances other than creatinine, it is advantageous that the enzyme exists in a free state.

  In the present invention, the concentration of the enzyme used in the sample-enzyme reaction mixture is not particularly limited, but it is preferably contained so that the final concentration after mixing with the sample is 0.001 to 10 mg / mL. More preferably, the final concentration is 0.01 to 1 mg / mL.

  The present invention is a highly sensitive method for measuring creatinine. According to the measurement method of the present invention, for example, 1 mg / dL or more, preferably 0.5 mg / dL or more, more preferably 0.25 mg / dL or more, further preferably 0.1 mg / dL or more, and, for example, 2 mg / dL or less. The creatinine contained in the sample can be measured at a concentration of preferably 5 mg / dL or less, more preferably 10 mg / dL or less, and even more preferably 20 mg / dL or less.

  An enzyme that reacts with creatinine may be reacted with creatinine in a sample in the presence of an alkali metal salt and / or an alkaline earth metal salt. For example, sodium chloride can be used as the alkali metal salt. These metal salts were expected to stabilize the detection signal, and an effect of increasing the signal was seen as an unexpected effect (see Examples). The addition concentration of the alkali metal salt and / or alkaline earth metal salt used is not particularly limited. For example, the effect is expected in the range of 1 mM to 1 M, preferably 10 to 200 mM, more preferably 50 to 100 mM. In Examples of the present invention, 50 to 100 mM sodium chloride was added, and a dramatic signal increase effect was observed.

  In the present invention, the sample-enzyme reaction mixture may be further mixed with other substances for the purpose of pH buffering action of the solution, protein stabilization and the like. In one implementation, an enzyme that reacts with creatinine may be reacted with creatinine in a sample in the presence of a buffer component. For example, examples of the buffer component include phosphate buffer, acetate buffer, borate buffer, citrate buffer, Tris buffer, Good buffer such as PIPES, MES, TES, MOPS, and HEPES. The concentration of the buffer component in the reaction mixture is not particularly limited, but if the concentration is too high, the ion concentration is difficult to change, and if it is too low, the change is excessive, so that an appropriate concentration is achieved to achieve high detection sensitivity. Select. The concentration of the buffer component is, for example, 0.1 to 20 mM, preferably 0.2 to 10 mM, more preferably 0.5 to 5 mM. In addition, alcohols such as ethanol, methanol, and propanol may be added.

  In the present invention, the sample-enzyme reaction mixture may contain other additives as necessary. Examples of additives include surfactants, stabilizers, preservatives, chelating agents, and activators. These are not limited in any way, but specific examples include the following. As the surfactant, any of nonion, anion, and cation may be used. Examples of nonionic surfactants include polyoxyethylene alkylphenyl ethers (trade name Triton X-100 and the like) and polyoxyethylene alkyl ethers, and anionic surfactants include alkylbenzene sulfonate and linear alkylbenzene sulfone. Examples of cationic surfactants such as acid salts include tetraalkylammonium salts. Stabilizers include saccharides such as sucrose, trehalose, cyclodextrin and their derivatives, amino acids such as arginine, lysine, histidine, glutamic acid, proteins such as albumin, salts such as alkali metals and alkaline earth metals, ammonia ions And gluconic acid. Antiseptics include antibiotics, azide compounds, other antibacterial agents and fungicides, and the like. Examples of the chelating agent include ethylenediaminetetraacetic acid and its salt. Examples of the activator include alkali metals and alkaline earth metals.

  In the method of the present invention, as a quantification method for creatinine, the endpoint method (measures the increase in signal derived from creatinine in a sample at a certain time), the kinetic method (signals the initial rate of enzyme derived from creatinine in the sample) Any of the above may be employed. Preferably, the end point method is adopted.

  In the present invention, the sample-enzyme reaction mixture comes into direct or indirect contact with the sensing part of the cumulative ISFET sensor. The optimal volume of the sample-enzyme reaction mixture depends on the system configuration, particularly the area of the sensing unit, but is preferably 0.1 to 500 μL. Furthermore, 0.1-30 microliters is mentioned as a suitable range.

  In the present invention, a cumulative ISFET sensor is used. In order to overcome the drawbacks of the conventional ISFET, the cumulative ISFET sensor repeatedly transfers the change in the depth of the potential well immediately below the sensing portion based on the change in the surface potential of the sensing portion of the ISFET as a charge to the drain. By configuring so that electric charges are accumulated, it is possible to reliably detect even a slight change in the surface potential of the sensing unit, and to detect a change in ion concentration with high sensitivity.

  As a form of the cumulative ISFET sensor used in the present invention, an input diode portion formed of a diffusion region opposite to the substrate and a floating diffusion formed on the surface side of a P-type or N-type semiconductor substrate at a predetermined interval. And an input gate and an output gate fixed via an insulating film at positions on the substrate surface corresponding to the start and end of the conduction channel to be formed between the input diode portion and the floating diffusion portion, respectively. A sensing portion made of an ion sensitive film fixed via an insulating film at a position on the substrate surface corresponding to an intermediate portion of the channel, and the substrate surface connected to a side of the floating diffusion portion away from the channel. A reset gate fixed through an insulating film at an upper position, and a substrate formed on the substrate surface portion on the side of the reset gate away from the floating diffusion portion; and And a reset diode section composed of a diffusion region of a mold, and the potential well changed depending on the concentration of ions acting on the sensing section as a result of adding and mixing an enzyme or a composition containing the enzyme and a sample to the sensing section The FET type sensor is configured to detect the amount of charge accumulated in the floating diffusion portion after the potential reset as a potential change in accordance with the depth of the current and the number of pumping out of the potential well.

  Hereinafter, embodiments according to the present invention will be described with reference to the drawings.

  FIG. 1A is a longitudinal sectional view showing a configuration of an FET type sensor element 20A that is a biosensor element according to an embodiment of the present invention, and FIG. 1B is an FET type sensor element of FIG. It is a schematic diagram which shows the energy level of 20A. FIG. 2 is a block diagram showing a configuration of a biosensor measurement apparatus 30 for measuring a sensor output signal of the FET type sensor element 20A of FIG. 3 is an operation timing chart when measuring the output characteristics of the FET type sensor element 20A of FIG. 1A by the biosensor measuring device 30 of FIG. Further, FIG. 4 is a diagram showing the operation of the FET type sensor element 20A of FIG. 1A, and is a schematic diagram showing the energy level similar to FIG. 1B, showing the transition of the potential state and accumulated charge. FIG.

Referring to FIG. 1A, an input diode 2 and a floating diffusion portion as a charge supply portion made of an n + -type diffusion layer are spaced apart from each other on the front side of a p-type semiconductor substrate 1 typically made of silicon. 3 is formed, and a reset diode 4 is formed at a small interval from the floating diffusion portion 3. In this case, an insulating film 5 made of SiO ₂ or Si ₃ N ₄ is formed on the semiconductor substrate 1 including the n + type diffusion layer. A conduction channel (n-type inversion layer) is formed on the surface of the semiconductor substrate 1 between the input diode 2 and the floating diffusion portion 3 in relation to the gate structure described below. As a result, the input diode 2 serves as a source and floats. An FET type sensor 20 having the diffusion unit 3 as a drain is configured. On the insulating film 5, the input gate 6 is adjacent to the input diode 2 corresponding to the channel start end portion, and the output gate 7 is adjacent to the floating diffusion portion 3 corresponding to the channel end portion is polysilicon, or A reset gate 8 is formed of a similar vapor deposition layer between the floating diffusion portion 3 and the reset diode 4. On the upper surfaces of the input gate 6, the output gate 7 and the reset gate 8, and the insulating film 5 outside the gate supporting these gates, a deposition film 10 typically formed of a Si ₃ N ₄ vapor deposition layer is formed. The Since the Si ₃ N ₄ film is denser than the SiO ₂ film and has a small oxygen diffusion coefficient, the recess formed between the input gate 6 and the output gate 7 itself is used as the sensing unit 9, so that good ion sensitivity is achieved. Construct a membrane. In addition to Si ₃ N ₄ , SiO ₂ , Au, Ta205, etc. can be used as the ion sensitive film. The ion-sensitive membrane of the sensing unit 9 may be immobilized with substances such as enzymes, antibodies, microorganisms, and nucleic acids that react with or bind to the analyte in the sample or that serve as a catalyst for the analyte reaction. It may exist in a state (not shown).

  Note that a relatively thick mask layer 11 made of a silicon oxide film or the like similar to the insulating film 5 is formed on the surface of the semiconductor substrate 1 outside the input diode 2 and the reset diode 4 to form the sensing unit 9 described above. The deposited film 10, which is a deposited film, covers the mask layer 11, and the sensing unit 9 is excluded from the deposited film 10, for example, on the protective film 12 made of phosphor glass and the protective film 12. An exterior film 13 with the outer surface flush is deposited. From the left side of FIG. 1A, electrode leads made of aluminum or the like are formed on the top surfaces of the input diode 2, the input gate electrode 6 and the output gate electrode 7, the floating diffusion portion 3, the reset gate 8 and the reset diode 4, respectively. A voltage according to the measurement sequence is applied through these electrode leads, or the potential of the floating diffusion part 3 is detected. The floating diffusion unit 3 is connected to an output voltage Vout of the electrode lead terminal (in this specification, the output signal voltage is also referred to as a signal) to a biosensor measurement device 30 including a source follower amplifier.

  In the present embodiment, the FET type sensor element 20A is an input diode portion formed of an N + type diffusion region which is formed on the surface side of the P− type semiconductor substrate 1 at a predetermined interval and is opposite to the semiconductor substrate 1. 2 and the floating diffusion portion 3 and the reset diode 4, and the reset gate 8 is also provided on the insulating film 5 between the floating diffusion portion 3 and the reset diode 4. A reset transistor is configured. First, the input gate 6 and the output gate 7 are fixed on the insulating film 9 corresponding to the intermediate part and the terminal part of the conduction channel to be formed between the input diode part 2 and the floating diffusion part 3, respectively. Since the gates 6 and 7 are adjacent to each other, the output gate 7 having a relatively narrow width and height is insulated from the input gate 6 by the deposited film 10 covering the adjacent side of the relatively wide input gate 6. It has a structured. The output gate 7 has a bottom surface in contact with the insulating film on the semiconductor substrate 1, an upper portion thereof penetrating the deposition film 10 and the protective film 12, and an upper end positioned in the exterior film 13.

  At the position on the substrate surface between the input diode 2 and the input gate 6 formed in this way, that is, at the position corresponding to the input end of the inversion channel to be formed, the ion-sensitive film is formed together with the insulating film 5 forming the bottom surface. A sensing portion 9 is formed which is sandwiched between the insulating film 5 on the input gate 6 side, which is a film, and the fault of the deposited film 10 and the protective film 12 on the input diode 2 side.

  Next, the configuration and operation of the biosensor measurement device 30 will be described below with reference to FIG. In FIG. 2, a DC voltage + Vdd is applied to the reset diode 4 from the DC voltage source 21 to the FET type sensor element 20A. Further, the controller 40 changes the predetermined voltages Vin, Vgin, Vsen, Vgout, Vgr having a predetermined pulse width, period, etc., which change in a timing chart, which will be described in detail later, to the FET type via the D / A converters 34 to 38. Applied to the sensor element 20A. The sensor output voltage Vout output from the FET type sensor element 20A is input to the A / D converter 32 via the source follower amplifier 31, and then from the clock generator 33 whose clock frequency and timing are controlled by the controller 40. A / D conversion is performed using the sampling clock. The digital data after A / D conversion is input to the controller 40 and then stored in the RAM 42.

  The controller 40 includes a ROM 41 for storing an operation program of the controller 40 and data necessary for executing the controller 40, a RAM 42 for storing digital data, and a flexible disk drive 43 for storing the digital data on a flexible disk. It is connected. The controller 40 is connected to a keyboard 45 through which an operator inputs instruction data such as instruction commands and measurement conditions via a keyboard interface 44, and instruction data and measurement results via a display interface 46. A liquid crystal display 47 for displaying is connected. In the present embodiment, the controller 30 has various display functions for displaying a comparison display, a graph display, a table display, and the like of the measured waveform on the liquid crystal display 47 based on the measured digital data.

  Further, a method for detecting the ion concentration using the FET-type sensor element 20A and the biosensor measurement device 30 of FIG. 1A will be described below with reference to FIGS.

  First, a voltage Vgin of 2 V, for example, is applied to the aqueous solution in the sensing unit 9 to make the surface potential of the semiconductor substrate 1 immediately below the sensing unit 5 constant. This is the initial setting for the potential well entrance. Next, as shown in FIG. 3, an appropriate DC voltage Vsen (for example, 2.0 V) is applied to the input gate 6 to fix the surface potential of the semiconductor substrate 1 immediately below the input gate 6 and input as a charge supply unit. A reverse bias voltage Vin = 5V is applied to the diode 2 and a reset voltage Vgr having a predetermined pulse width is applied to the reset gate 8 to set an initial value of the potential of the floating diffusion portion 3. At this time, the voltage Vgout of the output gate 7 is zero volts.

  The voltage Vin = 5V of the input diode 2 is set as a sufficient reverse bias, and the charge remaining in the input gate is suppressed very slightly as shown in FIG. 1B, and the upper end of the charge pool is at the level of the sensing unit 9. And does not enter after the sensing unit 9 (on the right side of the sensing unit 9 in FIG. 1B). In this case, when the negative ion concentration in the aqueous solution increases, the surface potential of the sensing unit 9 changes, and the surface potential of the semiconductor substrate 1 immediately below the sensing unit 9 further increases from the initial set value b0. When the negative ion concentration decreases or when the positive ions increase, the surface potential falls below b0.

  When the voltage Vin applied to the input diode 2 decreases from 5 V to 1.0 V, the amount of charge pool increases as the reverse bias is relaxed, and in this case, the level is the substrate surface potential b0 (potential) immediately below the sensing unit 9. The charge from the input diode 2 is supplied to the potential well immediately below the input gate 6 (see FIG. 4A).

  When the voltage Vin applied to the input diode 2 rises to 5 V again, charges are exhausted at the level of the surface potential immediately below the sensing unit 9, charges remain at the potential well capacity below that level, and other charges are The portion remaining in the diode 2 is passed through the input diode 2 and is returned to the DC voltage source 21 (see FIG. 4B). Also in this case, the amount of charge remaining in the potential well varies depending on the negative ion concentration, and the amount of change in the surface potential of the sensing unit 9 is converted into this amount of charge.

  Next, when the voltage Vgout of 5 V is applied to the output gate 7, the output gate 7 is opened, and the charge is transferred to the floating diffusion portion 3 that has been previously maintained at the reset potential (see FIG. 4C). . After this charge transfer, the voltage Vgout applied to the output gate 7 is lowered to 0 V, and the output gate 7 is closed (see FIG. 4D). Further, after the potential is read, the floating diffusion unit 3 applies the reset gate voltage Vgr to the reset gate 8 and the charge of the floating diffusion unit 3 is connected to the DC voltage source 21 of the power supply voltage + Vdd. 4 and further absorbed by the DC voltage source 21 to reset the initial output voltage Vout.

  As described above, the amount of change in the surface potential of the sensing unit 9 is accumulated as the amount of charge in the floating diffusion unit 3 by repeatedly performing the processes shown in FIGS. 4B to 4D. Then, the amount of potential change accumulated in the floating diffusion unit 3 is input to the biosensor measurement device 30 of FIG. 2 as the output voltage Vout, and then displayed and used for recording and other processing.

In the present invention, the surface area of the sensing part of the cumulative ISFET sensor is preferably smaller in order to reduce the size of the sensor and increase convenience, and to increase the volume of the sample-enzyme reaction mixture and increase the signal level. Larger is preferred. Therefore, the surface area of the sensing unit may be optimized according to the purpose of use, the target sample type, and the like. Generally, the surface area of the sensing portion is preferably 5 mm ² or less, more preferably 1 mm ² or less, preferably 0.0001 mm ² or more, 0.005 mm ² or more is more preferable.

The present invention provides a system for quantifying creatinine in a sample, comprising at least the following elements (a) and (b):
(A) Cumulative ion-sensitive field effect transistor (ISFET) sensor, where the cumulative ISFET sensor determines the depth of the potential well that has changed according to the concentration of ions acting on the sensing unit and the number of pumps out of the potential well. In response, an ISFET sensor that measures the amount of charge accumulated after a potential reset as a potential change; and (b) an enzyme that reacts with creatinine.

  The system of the present invention can be used in the above-described method for measuring creatinine of the present invention. Accordingly, the above-described matters concerning the method for measuring creatinine of the present invention can be applied to the system of the present invention. For example, creatinine deiminase [EC3.5.4.21] can be used as the enzyme that reacts with creatinine in (b).

  A biosensor measuring machine using a cumulative ISFET as a sensor element can measure changes in ion concentration caused by the reaction of creatine and an enzyme that reacts with creatine. FIG. 3 shows a block diagram of an exemplary biosensor measuring machine. In FIG. 3, the measuring instrument portion is indicated by a dotted line. The configuration of the measuring machine is as follows. 2: Amplify signal from sensor (1); 3: Convert analog signal to digital signal; 4: Store digital signal; 5: Display digital signal in graphs and tables; 6: Past for comparison 7: Extract display data as needed; 8: Supply power to elements and circuits; Receive signal at 9: 3 and feed back measurement timing; 10: Sensor voltage 11: Set measurement conditions such as pulse width and voltage; 12: Instruct cumulative measurement; 13: Save measurement data; 14: Print measurement data.

  EXAMPLES Hereinafter, although an Example demonstrates this invention further in detail, this invention is not limited by these Examples.

[Preparation of reagent composition]
The following reagent composition was prepared.
1 mM Tris-HCl buffer (pH 7.5)
1.39 U / mL Creatinine deiminase (Toyobo: CNI-311, 0.1 mg / mL)
0-100 mM sodium chloride

[Preparation of measurement sample]
The following sample solution was prepared.
1 mM Tris-HCl buffer (pH 7.5)
0-10mg / dL Creatinine (Tokyo Kasei)

[Preparation of control reagent]
The following reagents were prepared.
1 mM Tris-HCl buffer (pH 7.5)

[Device used]
A measuring device (a measuring device for biosensor development, manufactured by Bio-X) using a cumulative ISFET sensor (AMIS sensor, manufactured by Bio-X) was used. The apparatus configuration is based on the prior art (Patent Document 17), and its block diagram is shown in FIG.
Tantalum pentoxide was used for the ion sensitive film on the gate of the ISFET. The surface area of the sensing part was 0.05 mm ² .

[Measurement condition]
The amount of creatinine in the sample was measured under the following measurement conditions.
18 μL of the reagent composition was added to the sensing part A of the sensor, and 18 μL of a control reagent not containing creatine deiminase was added to the sensing part B (for control). Preheating was performed at 37 ° C. for 5 minutes to stabilize the temperature, and then the signal was measured for 50 seconds to check the state of the sensor. Then, 2 μL each of the measurement samples prepared at a creatinine concentration of 0 to 10 mg / dL were added to and mixed with the reagents of the sensing parts A and B, and after the signal was stabilized (1 to 2 minutes later), the signal ( Signal increase compared to the signal from the sensing part B for control) was measured every 5 seconds for 4-5 minutes. The cumulative number of cumulative ISFET sensors was set to 10 times.

[Example 1]
Measurement was performed on each measurement sample having a creatinine concentration of 5 mg / dL, 1 mg / dL, and 0 mg / dL with a reagent not containing sodium chloride, and the time course of each signal increase was measured. The results are shown in FIG. It can be seen that when the creatinine deiminase in the reagent reacts with the creatinine in the sample, the ion concentration increases, a signal increase according to the creatinine concentration is observed, and measurement is good.

[Example 2]
In the same manner as in Example 1, measurement was performed on each measurement sample having a creatinine concentration of 0, 1, 2, 3, 4, 5, 6, 7, 8, 10 mg / dL, and each signal increase was increased for 245 seconds. It was measured. From the measurement data at 245 seconds, the linearity between the creatinine concentration and the measured value was observed. The results are shown in FIG. A very good correlation is observed between the creatinine concentration and the measured value, and it can be seen that the amount of creatinine is measured well by the method of the present invention.

[Example 3]
Measurement was performed on each measurement sample having a creatinine concentration of 0.5 mg / dL, 0.25 mg / dL, and 0 mg / dL with a liquid reagent containing 100 mM sodium chloride, and the time course of each signal increase was measured. The results are shown in FIG. It can be seen that when the creatinine deiminase in the reagent reacts with the creatinine in the sample, the ion concentration increases, a signal increase according to the creatinine concentration is observed, and measurement is good.

[Example 4]
In the same manner as in Example 3, measurement was performed on each measurement sample having a creatinine concentration of 0, 0.1, 0.25, 0.5, 0.75, 1, 1.5, 2 mg / dL, Signal increase was measured for 245 seconds. From the measurement data at 245 seconds, the linearity between the creatinine concentration and the measured value was observed. The results are shown in FIG. A good correlation is observed between the creatinine concentration and the measured value, and it can be seen that the amount of creatinine is measured well by the method of the present invention.

[Example 5]
Measurement was performed on each measurement sample having a creatinine concentration of 0.5 mg / dL, 0.25 mg / dL, 0.1 mg / dL, and 0 mg / dL with a liquid reagent containing 50 mM sodium chloride, and the time of each signal increase The course was measured. The results are shown in FIG. It can be seen that when the creatinine deiminase in the reagent reacts with the creatinine in the sample, the ion concentration increases, a signal increase according to the creatinine concentration is observed, and measurement is good.

[Example 6]
In the same manner as in Example 5, measurement was performed on each measurement sample having a creatinine concentration of 0, 0.1, 0.25, 0.5, 0.75, 1 mg / dL, and each signal increase was increased for 245 seconds. It was measured. From the measurement data at 245 seconds, the linearity between the creatinine concentration and the measured value was observed. The results are shown in FIG. A very good correlation is observed between the creatinine concentration and the measured value, and it can be seen that the amount of creatinine is measured well by the method of the present invention.

  As described above, according to the present invention, it has been achieved to provide a novel method for quantifying creatinine in a sample, and a system therefor, in contrast to conventional methods and systems. In particular, it has been achieved to provide a method and a system therefor that can measure creatinine with a single enzyme.

(A) is a longitudinal cross-sectional view which shows the structure of FET type sensor 20 which concerns on one Embodiment to this invention, (B) is a schematic diagram which shows the energy level of FET type sensor 20 of FIG. 1 (A). . It is a block diagram which shows the structure of the biosensor measuring apparatus 30 for measuring the sensor output signal of the FET type sensor 20 of Fig.1 (a). 3 is an operation timing chart when measuring output characteristics of the FET type sensor 20 of FIG. 1A by the biosensor measuring device 30 of FIG. 2. It is a figure which shows operation | movement of FET type | mold sensor 20 of FIG. 1 (A), Comprising: It is a schematic diagram which shows transition of a potential state and an accumulation charge, and shows the energy level similar to FIG. 1 (B). The time course of the measurement of creatinine in a sample in the method of the present invention (Example 1) is shown. The creatinine dilution linearity in the method of the present invention (Example 2) is shown. The time course of the measurement of creatinine in a sample in the method of the present invention (Example 3) is shown. The creatinine dilution linearity in the method of the present invention (Example 4) is shown. The time course of the measurement of creatinine in a sample in the method of the present invention (Example 5) is shown. The creatinine dilution linearity in the method of the present invention (Example 6) is shown.

Explanation of symbols

1 ... Semiconductor substrate,
2 ... Input diode,
3 ... floating diffusion part,
4 ... Reset diode,
5 ... Insulating film,
6 ... Input gate,
7 ... Output gate,
8 ... Reset gate,
9 ... Sensing part,
10 ... Deposited film 11 ... Mask layer,
12 ... Protective film,
13 ... exterior membrane,
20 ... FET type sensor,
20A: FET type sensor element,
21 ... DC voltage source,
30 ... Biosensor measuring device,
31 ... Source follower amplifier,
32 ... A / D converter,
33 ... Clock generator,
34 to 38 ... D / A converter,
40 ... Controller,
41 ... ROM,
42 ... RAM,
43. Flexible disk drive,
44 ... Keyboard interface,
45 ... Keyboard,
46 ... Display interface,
47 ... Liquid crystal display.

Claims (12)

A method for measuring creatinine, comprising the following steps:
(1) a step of reacting creatinine in a sample with an enzyme that reacts with creatinine;
(2) A step of detecting a concentration change of ions generated by the reaction using a cumulative ion sensitive field effect transistor (ISFET) sensor, wherein the cumulative ISFET sensor changes according to the ion concentration acting on the sensing unit. An ISFET sensor that measures the amount of charge accumulated after resetting the potential as a potential change according to the depth of the potential well and the number of pumps from the potential well; and (3) In the sample based on the detected ion concentration change Quantifying the amount of creatinine.   The measurement method according to claim 1, wherein the ion generated by the reaction is an ammonium ion.   The measuring method according to claim 1, wherein the enzyme that reacts with creatinine comprises a single enzyme.   The measurement method according to claim 3, wherein the enzyme that reacts with creatinine is creatinine deiminase [EC3.5.4.21].   The measuring method according to claim 1, wherein an enzyme that reacts with creatinine is present in a free state in the reaction mixture.   The measurement method according to claim 1, wherein creatinine contained in the sample at a concentration of 0.1 mg / dL or more can be measured.   The measurement method according to claim 1, wherein an enzyme that reacts with creatinine is reacted with creatinine in a sample in the presence of an alkali metal salt and / or an alkaline earth metal salt.   The measurement method according to claim 1, wherein an enzyme that reacts with creatinine is reacted with creatinine in the sample in the presence of the buffer component, and the concentration of the buffer component in the reaction mixture is 0.5 to 5 mM. The measuring method according to claim 1, wherein the surface area of the sensing part of the cumulative ISFET sensor is 1 mm ² or less.   The measurement method according to claim 1, wherein an enzyme that reacts with creatinine is reacted with creatinine in a sample in a reaction mixture having a volume of 30 µL or less. A system for quantifying creatinine in a sample, comprising at least the following elements (a) and (b):
(A) Cumulative ion-sensitive field effect transistor (ISFET) sensor, where the cumulative ISFET sensor determines the depth of the potential well that has changed according to the concentration of ions acting on the sensing unit and the number of pumps out of the potential well. In response, an ISFET sensor that measures the amount of charge accumulated after a potential reset as a potential change; and (b) an enzyme that reacts with creatinine.   12. The system according to claim 11, wherein the enzyme that reacts with creatinine in (b) is creatinine deiminase [EC 3.5.4.21].

Images