1. Field of the Invention
The present invention is generally related to biosensors and the fabrication method thereof, and more particularly a potentiometric biosensor for detection of creatinine and urea.
2. Description of the Prior Art
Biosensor is commonly defined as an analytical device which combines energy converter with immobilized biomolecules for detecting specific chemicals via the interaction between biomolecules and such specific chemicals. The above-mentioned energy converter can be a potentiometer, a galvanometer, an optical fiber, a surface plasma resonance, a field-effect transistor, a piezoelectric quartz crystal, a surface acoustic wave, and so on. The field-effect transistor used to fabricate the miniaturized device via mature semiconductor process has become an important technique for the current market trend of developing light and portable products.
A model of a biosensor is based on an analytic method of detecting a organic compound. The analytic method was established using the specificity theory of an enzyme and its substrate. This specificity theory is proposed by Clark et al. in 1962 [Clark L. C., C. Lyois, “Electrode system for continuous monitoring in cardiovascular surgery”, Annals of the New York Academy of Sciences, vol. 102, pp. 29-33, 1962.). According to the Intechno Cunsulting investigation reports, [Zhang Chen-Sui, market demand and technology-developing tendency of sensors, Industrial Economics & Knowledge Center, 2002.], biotechnology combining with a semiconductor technology and reducing device size will have the advantages of small volumes, small weight, high reliability, high precision, good performance, low cost, and mass production.
U.S. Pat. No. 5,804,047 [Isao Karube, Susan Anne Clark, Ryohei Nagata, “Enzyme-immobilized electrode, composition for preparation of the same and electrically conductive enzyme”, 1998.] discloses an enzyme sensing system suitable for detecting a specific substance. A electrode immobilized the enzyme can immobilize a mixture which comprises a conductive enzyme and other conductive material formed by using covalent bonds to connect the enzyme and the electron transport substance, and the ways to immobilize a enzyme onto a base material are screen printing, and brushing.
U.S. Pat. No. 5,945,343 [Christiane Munkholm, “Fluorescent polymeric sensor for the detection of urea”, 1999.] discloses a fluorescent polymeric sensor for the detection of urea. The fluorescent polymeric sensor comprises three layers. The first layer is a protonated pH sensitive fluorophore immobilized on a hydrophobic polymer. The second layer is composed of urease and a polymer; and the third layer is a polymer. The structure of the sensor disclosed in the invention is simple and the sensor can be fabricated as a miniaturized and disposable device. Without improvement of the operation stability and the production of the optical sensor, the major disadvantage of the invention is high cost, as compared to voltage-mode and current-mode sensor system.
Although the concentration of urea or creatinine can be measured via spectrum analysis, but the general method is the enzyme method [C. Puig-Lleixa, C. Jimenez, J. Alonso, J. Bartroli, “Polyurethaneacrylate photocurable polymeric membrane for ion-sensitive field effect transistor based urea biosensors”, Analytica Chimica Acta, vol. 389, pp. 179-188, 1999; R. Koncki, I. Walcerz, E. Leszczynska, “Enzymatically modified ion-selective electrodes for flow injection analysis”, Journal of Pharmaceutical and Biomedical Analysis, vol. 19, pp. 633-638, 1999; A. B. Kharitonov, M. Zayats, A. Lichtenstein, E. Katz, I. Willner, “Enzyme monolayer-funtionalized field-effect transistors for biosensor applications”, Sensors and Actuators B, vol. 70, pp. 222-231, 2000.]. At present, the commercial biosensors are based on field-effect transistors and current-mode circuit. The principle of the current-mode technology is to detect a small electric current in organisms. It has fast response, but the output stage circuit needs an additional bias voltage to convert the signals. Therefore, the fabrication of current-mode biosensors is more complicated design and has higher costs. A redox reaction occurs when the current-mode biosensors detect specific chemicals and it produces a small electric current. The current flows through the surface of sensor surface and damages the biological molecules (such as enzymes), and hence affect the follow-up use of enzymes for chemical reaction.
Moreover, the biosensors based on field-effect transistors are mostly produced by the semiconductor manufacturing process that needs strict conditions (such as the need for high vacuum environment, etc.), which results in high costs of production. Since the rise of medical and health consciousness, the combination of biosensors and medical examination has become a trend (such as the measurement of creatinine concentration in human serum). How to make the biosensors having simple structure, good stability, and replaceable with low cost in medical purpose has become the current trend in sensor development.
In accordance with the present invention, a potentiometric biosensor for detection of creatinine and urea is provided for commercial need.
The present invention further discloses a potentiometric biosensor for detection of creatinine and urea. The potentiometric biosensor revealed in this invention is for detecting the content of creatinine in serum and urea in urine which are important indicators for the renal, thyroid and muscle function of human body.
The present invention discloses a potentiometric biosensor based on field-effect transistors which can be fabricated to form the miniaturized component via semiconductor process. The potentiometric biosensor of the present invention doesn't need an additional bias voltage to convert the signals. The disclosed biosensor comprises a substrate, at least two working electrode on the substrate, at least one reference electrode on the substrate, an internal reference electrode on the substrate, and a packaging structure which separates the adjacent electrodes. The working electrode comprises urease or creatinine iminohydrolase (CIH). The detection signal is transmitted for further processing through a wire or an exposed surface on the biosensor. The disclosed biosensor is replaceable.
What is probed into the invention is a potentiometric biosensor for detection of creatinine and urea. Detail descriptions of the structure and elements will be provided in the following in order to make the invention thoroughly understood. Obviously, the application of the invention is not confined to specific details familiar to those who are skilled in the art. On the other hand, the common structures and elements that are known to everyone are not described in details to avoid unnecessary limits of the invention. Some preferred embodiments of the present invention will now be described in greater detail in the following specification. However, it should be recognized that the present invention can be practiced in a wide range of other embodiments besides those explicitly described, that is, this invention can also be applied extensively to other embodiments, and the scope of the present invention is expressly not limited except as specified in the accompanying claims.
U.S. Pat. No. 5,858,186 [Robert S. Glass, “Urea biosensor for hemodialysis monitoring”, 1999.] discloses an electrochemical sensor for quantitatively detecting the urea concentration of the dialysis waste liquid in the process of blood dialysis. The sensor uses an enzyme to hydrolyze the urea and detects the variation of pH generated by the hydrolysis. The structure of the sensor is good for mass production and reducing production cost, so the structure has an advantage for developing a disposable sensor. For a typical application, this sensor is usually used to diagnose the stop point of the blood dialysis at an inspection center or used with an appropriate computer system. This sensor can also be used at home by a dialysis patient, and only requires a few drop of blood sample to perform detection.
U.S. Pat. No. 4,691,167 [Hendrik H. v. d. Vlekkert, and Nicolaas F. de Rooy, “Apparatus for determining the activity of an ion (pIon) in a liquid”, 1987) discloses an apparatus determining the reactivity of an ion in a liquid. The system comprises a measuring circuit, an ion sensitive field effect transistor (ISFET), a reference electrode, a temperature sensor, amplifiers, a controller, computing circuits, and a memory. Since the sensitivity is a function of temperature and drain current and is decided by a variable of gate voltage, the sensitivity can be obtained by calculating formulas stored in the memory.
U.S. Pat. No. 5,474,660 [Ian Robins, John E. A. Shaw, “Method and apparatus for determining the concentration of ammonium ions in solution”, 1995.] discloses an apparatus and a method thereof for detecting ammonium ion concentration, wherein an ammonia gas sensor is placed into a container, and a solution containing ammonium ions is placed into a partial region of the container; hydroxyl ions are generated from the solution by an electrochemical generator at the vicinity of the container placing the ammonia gas sensor, and then the sensor detects the ammonia gas through a film, transformed by the ammonium ions in the solution. The sensor disclosed by this patent thus using the above-mentioned method to detect the ammonium ion concentration in a solution.
U.S. Pat. No. 6,021,339 [Atsushi Saito, Soichi Saito, Masako Miyazaki, “Urine testing apparatus capable of simply and accurately measuring a partial urine to indicate urinary glucose value of total urine”, 2000.] discloses a uric acid multiple sensor which comprising a sensing device for measuring urea and at least one component for detecting sodium and chlorine ions in uric acid. As far as we know, the specific weight of uric acid is based on the detected signals generated from the concentration of each device. Besides, a component for detecting the units of glucose must be added herein and then finally the particular specific weight in sugar can be used to correct the measured sugar [that is, glucose base line]. After that, after all uric acid secreted 24 hours, the detected conditions can be understood simply and accurately from a partial uric acid.
U.S. Pat. No. 4,970,145 [Hung P. Bennetto, Gerard M. Delaney, Jeremy R. Mason, Chrispother F. Thurston, John L. Stirling, David R. DeKeyzer, “Immobilized enzyme electrodes”, 1990.) discloses an enzyme electrode fabricated using a carbon electrode as a base structure. The enzyme electrode with this structure allows the enzyme [such as glucose oxidized enzyme] attach on the electrode to fabricate an amperometric sensor with good response and stability. The substrate material of the electrode is a thin carbon electrode plated with platinum seldom and can perform detection with the condition that the dissolved oxygen at low level. The enzyme sensor runs measurement in a 10 mM glucose solution, and the reaction result is a current density having several hundreds microampere per square centimeter with a short response time. While preserved under a humid environment at room temperature, the sensor still has a good stability and several months of its working life.
U.S. Pat. No. 5,397,451 [Mitsugi Senda, Katsumi Hamamoto, Hisashi Okuda, “Current-detecting type dry-operative ion-selective electrode”, 1995.] discloses an amperometric and dry-operated ion-selective electrode which comprising a work electrode and an auxiliary electrode, both are fabricated on an insulating substrate. A first layer is a hydrophilic polymer, but the ion-selective membrane using a hydrophobic polymer.
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The present invention discloses a method of forming a potentiometric biosensor. First, provide a substrate, and then form an internal reference electrode on said substrate. Then form at least one counter electrode on said substrate. Then form at least two working electrodes on said substrate. Finally, form a packaging structure to separate the adjacent electrodes. A better method further comprises forming a conducting layer between these electrodes and the substrate and a wire connected to the second conducting layer to facilitate transmission of the detection signal, before the electrodes are formed on the substrate. Another better method further comprises forming an exposed surface on said at least two working electrodes, at least one counter electrode, and internal reference electrode to electrically couple with the external electrical devices and transmits the detection signal.
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The present invention discloses a measuring method using a potentiometric biosensor, comprising: measuring a reference voltage via putting at least two working electrodes into a buffer solution, and measuring the reference voltage. Next, amplify the readout signal of at least two working electrodes using at least two instrumental amplifiers, and measure a reaction voltage via putting at least two working electrodes into the tested solution. These at least two instrumental amplifiers electrically couple with a signal measurement module separately, and the signal measurement module measures the output signals from instrumental amplifiers to produce plural measured values, and each measured value corresponds to each output signal of the instrumental amplifiers.
The present invention discloses a potentiometric biosensor comprising a substrate, at least two working electrodes on the substrate, at least one reference electrode on the substrate, an internal reference electrode on the substrate, and a packaging structure which separates the above-mentioned at least four electrodes. The substrate comprises one selected from the group consisting of the following: insulating glass, non-insulated indium-tin oxide glass, non-insulated tin dioxide glass, and polyethylene terephthalate (PET). About the condition of forming an internal reference electrode, please refer to the condition of forming a tin dioxide/indium-tin oxide/glass-extension ion biosensor or a tin dioxide/carbon/PET-extension ion biosensor presented as follows. About the condition of forming a counter electrode, please refer to the condition of forming ammonium ion-selective electrode presented as follows. About the condition of forming at least two working electrodes, please refer to the condition of forming a potentiometric urea sensing film and a potentiometric creatinine sensing film presented as follows.
(A) The condition of forming a tin dioxide/indium-tin oxide/glass-extension ion biosensor:
(1) an Indium-tin oxide glass, wherein the thickness of indium-tin oxide film is 230 Å;
(2) a sensing window (2×2 mm2); and
(3) the condition of forming tin dioxide sensing film: the thickness of tin oxide sensing film is 2000A, which is formed via the sputtering tin dioxide using a tin dioxide target in a gas mixtures of Ar and O2 (4:1) with the air pressure of 20 mtorr, the radio frequency power of 50 Watt, and the substrate temperature of 150° C.
(B) The condition of forming a tin dioxide/carbon/PET-extension ion biosensor:
(1) a carbon/PET substrate with a 2 mm diameter sensing window; and
(2) a tin dioxide sensing film: the thickness of tin oxide sensing film is 2000 Å, which is formed via sputtering tin dioxide using a tin dioxide target in a gas mixtures of Ar and O2 (4:1) with the air pressure of 20 mtorr, the radio frequency power of 50 Watt and the substrate temperature of 150° C.
(C) The condition of forming an ammonium ion-selective electrode:
(1) mixing poly(vinyl chloride) carboxylated (PVC-COOH) 33%, bis(2-ethylhexyl) sebacate (DOS) 66%, and nonactin 1%, and then adding tetrahydroofuran (THF) 0.375 ml. Finally, mixing it using an ultrasound device;
(2) dropping 2 microliter of above-mentioned ammonium ion-selective on the tin dioxide sensing window; and
(3) putting it in a dark room for 12 to 24 hours to immobilize the ammonium ion-selective electrode.
(D) The condition of forming a potentiometric urea sensing film:
(1) diluting PVA-SbQ 120 mg/100 microliter (pH 7.0, phosphate solution 5 mmol/L), and mixing it with enzyme solution 10 mg/100 microliter (pH 7.0, phosphate solution 5 mmol/L). The ratio of volume is 1:1;
(2) dropping 1 microliter solution on the tin dioxide sensing window; and photo polymerization it via 4 Watt, 365 nm UV light for 20 minutes; and
(3) putting it in a dark room for 12 to 24 hours to immobilize the urea sensing film.
(E) The condition of forming a potentiometric creatinine sensing film:
(1) diluting PVA-SbQ 50 mg/100 microliter (pH 7.0, phosphate solution 5 mmol/L), and mixing it with enzyme solution 0.2 mg/ml (pH 7.0, phosphate solution 5 mmol/L). The ratio of volume is 1:1;
(2) dropping 1.0 microliter solution on the creatinine sensing window; and photo polymerization it via 4 Watt, 365 nm UV light for 20 minutes; and
(3) putting it in a dark room for 12 to 24 hours to immobilize the creatinine sensing film.
Obviously many modifications and variations are possible in light of the above teachings. It is therefore to be understood that within the scope of the appended claims the present invention can be practiced otherwise than as specifically described herein. Although specific embodiments have been illustrated and described herein, it is obvious to those skilled in the art that many modifications of the present invention may be made without departing from what is intended to be limited solely by the appended claims.
Number | Date | Country | Kind |
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097129173 | Aug 2008 | TW | national |