CA2068475C - Biosensor and its manufacture - Google Patents
Biosensor and its manufactureInfo
- Publication number
- CA2068475C CA2068475C CA002068475A CA2068475A CA2068475C CA 2068475 C CA2068475 C CA 2068475C CA 002068475 A CA002068475 A CA 002068475A CA 2068475 A CA2068475 A CA 2068475A CA 2068475 C CA2068475 C CA 2068475C
- Authority
- CA
- Canada
- Prior art keywords
- electrode
- base plate
- counter electrode
- insulating base
- working electrode
- Prior art date
- Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
- Expired - Lifetime
Links
Classifications
-
- C—CHEMISTRY; METALLURGY
- C12—BIOCHEMISTRY; BEER; SPIRITS; WINE; VINEGAR; MICROBIOLOGY; ENZYMOLOGY; MUTATION OR GENETIC ENGINEERING
- C12Q—MEASURING OR TESTING PROCESSES INVOLVING ENZYMES, NUCLEIC ACIDS OR MICROORGANISMS; COMPOSITIONS OR TEST PAPERS THEREFOR; PROCESSES OF PREPARING SUCH COMPOSITIONS; CONDITION-RESPONSIVE CONTROL IN MICROBIOLOGICAL OR ENZYMOLOGICAL PROCESSES
- C12Q1/00—Measuring or testing processes involving enzymes, nucleic acids or microorganisms; Compositions therefor; Processes of preparing such compositions
- C12Q1/001—Enzyme electrodes
Abstract
The present invention is directed to a biosensor provided with an insulating base plate, an electrode system mainly consisting of a working electrode and a counter electrode formed on the insulating base plate, and a reaction layer on the electrode system. The counter electrode is partially a circular arc. The invention also provides a manufacturing method for the biosensor comprising a step to form a base by arranging leads, an electrode system, and an insulating layer on an insulating base plate, and a step to form a reaction layer mainly composed of an enzyme on the electrode system.
Advantages of the present invention include easy forming of the reaction layer and prevention of delamination, whereby measuring accuracy, preservative properties and reliability of the biosensor are improved. As well, the highly efficient biosensors can be mass produced at low costs.
Advantages of the present invention include easy forming of the reaction layer and prevention of delamination, whereby measuring accuracy, preservative properties and reliability of the biosensor are improved. As well, the highly efficient biosensors can be mass produced at low costs.
Description
BIOSENSOR AND ITS MANUFACTURE
The present invention relates to a biosensor designed to quickly quantify with ease and high accuracy a specific component in various kinds of sample liquids, and a method for manufacturing the biosensor.
Various kinds of biosensors utilizing the specific catalytic effect of an enzyme have been developed in the past years, but a more accurate one is desired.
An example of a biosensor of the aforementioned type, namely, a glucose sensor applied in clinical technology will be described below.
Conventionally, blood plasma obtained from the blood of a patient after treatment in a centrifuge has been measured in order to quantify glucose in the blood. This method requires time and labour. Therefore, a sensor that can measure glucose concentration in the blood using whole blood is desirable.
As a simple glucose sensor, one similar to a test paper used for inspection of urea has been provided. This glucose sensor has a stick-like supporting body and a carrier fixed to the supporting body which includes an enzyme reactive only to glucose and colouring matter which will change colour as a result of the enzyme reaction. When blood is dropped onto the carrier of the sensor and the change in the colouring matter after a predetermined time is visually measured with or without optical means, the amount of glucose contained in the blood can be determined. According to the quantifying method using the glucose sensor of this type, however, the result is greatly influenced by colouring matter in the blood, and therefore the measuring accuracy is low.
In the meantime, according to a different method, a specific component of a living sample, e.g. blood or the like, can be quantified with high accuracy without requiring dilution or stirring of the sample liquid. An example of a biosensor using this method is proposed in Japanese Patent Laid-open Publication Tokkaihei 1-212345 (212345/1989), which will be discussed below.
~ - 2 - 2068475 The biosensor has an electrode system formed on an insulating base plate by screen printing or the like, and an enzyme reaction layer which consists of a hydrophilic polymer layer, a oxidoreductase and an electron acceptor is formed on the electrode system.
The biosensor of the above structure operates as follows.
When a sample liquid is dropped on the enzyme reaction layer, the oxidoreductase and electron acceptor are dissolved in the sample liquid, so that the enzyme reaction progresses with a substrate in the sample liquid. As a result, the electron acceptor is reduced. After the enzyme reaction is complete, the reduced electron acceptor is electrochemically oxidized.
The concentration of the substrate in the sample liquid is obtained from an oxidization current obtained at this time.
In the conventional biosensor, the working electrode and counter electrode are formed of material different from that of the insulating layer. Therefore, the reaction layer on the electrode system mainly consisting of the working electrode and counter electrode is formed in conformity with the configuration of the counter electrode. Since the counter electrode of the conventional biosensor is polygonal and mainly square, the reaction layer is apt to be separated from a corner of the counter electrode.
In a preferred embodiment the present invention provides a biosensor for measuring a substrate in a liquid sample, which is provided with an insulating base plate, an electrode system mainly consisting of a working electrode and a counter electrode formed on said base plate, and a reaction layer formed on said electrode system area and comprising an enzyme and an electron acceptor, wherein the reduced electron acceptor caused by means of a reaction between the enzyme and the said substrate in the liquid sample is electrochemically oxidized and the resulting value of oxidation current between the working electrode and the counter electrode is measured to determine the concentration of the substrate, characterized in that the counter electrode is provided with a circular arc portion at the outer edge of the counter electrode against the ~ .i~
direction of sample supply containing the substrate to be measured. More preferably, the counter electrode is provided with an outer edge which is partially broken.
A biosensor according to a further aspect of the present invention is characterized in that the working electrode is provided with an outer edge of circular, elliptical, deformed from the circular shape, or mainly polygonal shape, the vertex of which assumes a curve.
A biosensor according to yet a further aspect of the present invention is characterized in that the distance between the counter electrode and working electrode is held equal.
The present invention further features a manufacturing method of the biosensor, which is comprised of a manufacturing step for forming a base by providing leads, an electrode system consisting mainly of a working electrode and a counter electrode and, then an insulating layer are formed on an insulating base plate, and a forming step whereby a reaction layer mainly composed of an enzyme is formed onto the working electrode. The manufacturing step of the base includes the following three processes: a process to form the leads on the insulating base plate; a process to form the electrode system consisting mainly of the working electrode and counter electrode onto the insulating base plate; and a process to form the insulating layer on the insulating base plate.
The disadvantages inherent in the prior art can be solved by the biosensor of the present invention.
Moreover, the following effects can be achieved by the aforementioned means. In the first place, since the counter electrode is formed circular, elliptical or deformed from the circular shape, formation of the reaction layer becomes easy and delamination thereof can be prevented.
Further, since the working electrode is formed circular, elliptical, deformed from the circular shape, or mainly polygonal with a vertex thereof turned round by a curve, the density distribution of a current between the working and counter electrodes is made uniform. This is further enhanced _ 4 _ 2068475 by keeping the distance of the counter electrode from the working electrode equal. In consequence, the concentration of the substrate can be measured with higher accuracy.
Likewise, if a part of the circular or deformed counter electrode is removed, the lead part of the working electrode is prevented from being shortcircuited with the counter electrode by the presence of a pin hole in the insulating layer, etc.
According to the manufacturing method of the present invention, in the case where the electrode system is made of the working and counter electrodes, that is, the electrode system is a system of two electrodes, the working and counter electrodes can be manufactured in one process although they have been conventionally formed in different processes, thereby reducing the number of manufacturing processes of the biosensor .
These and other objects and features of the present invention will become clear from the following description taken in conjunction with the preferred embodiments thereof with reference to the accompanying drawings throughout which like parts are designated by like reference numerals, and in which:
Fig. 1 is a cross sectional view of a glucose sensor without a cover and a spacer according to one embodiment of the present invention;
Fig. 2 is an exploded perspective view with a reaction layer removed from the glucose sensor of Fig. l;
Fig. 3 is a plan view of a base of a glucose sensor according to a further embodiment of the present invention;
Fig. 4 is a plan view of a base of a glucose sensor according to a still further embodiment of the present invention;
Fig. 5 is an exploded perspective view with a reaction layer removed from the glucose sensor of Fig. 4; and Figs. 6, 7, 8, and 9 are plan views of a base of a glucose sensor according to still different embodiments of the present invention.
~ ~ 5 ~ 2068475 (Embodiment 1) A glucose sensor will be described hereinbelow as one example of a biosensor of the present invention.
Fig. 1 is a cross sectional view of a glucose sensor without a cover and a spacer according to one embodiment of the present invention, and Fig. 2 is an exploded perspective view with a reaction layer removed from the glucose sensor.
Referring to these drawings, silver paste is printed on an insulating base plate 1 of polyethylene terephthalate through screen printing, thereby forming leads 2, 3.
Conductive carbon paste including a resin binder is further printed in contact with the lead 2, thereby forming a working electrode 4. Then, insulating paste is printed to form an insulating layer 6.
The insulating layer 6 covers the outer periphery of the working electrode 4, so that the area of the exposed part of the working electrode 4 is held constant. At the same time, the insulating layer 6 covers the unnecessary part of the leads 2, 3. More specifically, the unnecessary part of the lead 2 all but the part of the lead 2 connecting with a working device. The unnecessary part of the lead 3 is all but the part of the lead 3 connecting with a counter electrode 5 which will be described later and the part connecting ~ith the measuring device. The glucose sensor is connected to the measuring device via end parts of the leads 2, 3. The measuring device impresses a constant voltage to an electrode system which will be described later, reads the value of an oxidization current, while performing a like function.
The counter electrode 5 is formed by printing conductive carbon paste including a resin binder into contact with the lead 3, which is circular in the outer periphery thereof.
A 0.5 wt% solution of carboxymethyl cellulose (referred to as CMC hereinafter) as a hydrophilic polymer in water is spread onto the electrode system consisting of the working electrode 4 and counter electrode 5 and dried, so that a CMC
layer is formed.
-The surface of the insulating layer according to the instant embodiment is water-repellent, and therefore the CMC
aqueous solution is repelled by the insulating layer and spreads only over the electrode system of the working and counter electrodes 4, 5. If the counter electrode 5 is formed circular as illustrated in Fig. 2, it is possible to form a uniform CMC layer. Moreover, delamination of the layer which has sometimes occurred after drying when the counter electrode is square can be prevented.
Subsequently, a mixed solution obtained by dissolving glucose oxidase (referred to as GOD hereinafter) as an enzyme and potassium ferricyanide as an electron acceptor in 0.5 wt%
CMC aqueous solution is dropped onto the CMC layer, dried for ten minutes in a warm drier at 50C. As a result, a reaction layer 7 is formed.
A solution of lecithin in toluene as a surface-tension surfactant is spread from a part of the base plate 1 corresponding to a sample feed port 13 onto the reaction layer 7 and dried, and a lecithin layer 8 is formed.
The insulating base plate 1, a spacer 11 and a cover 12 are bonded to each other in the position as indicated by the broken chain line in Fig. 2.
A 3~ glucose standard solution is supplied as a sample liquid from the sample feed port 13 to the glucose sensor formed in the above-described manner.
When the sample liquid is brought in contact with the sample feed port 13, it is also in contact with the lecithin layer 8 and is smoothly guided onto the reaction layer by the lecithin layer.
It is not always necessary to distinguish the sample feed port 13 from an air hole 14. The sample liquid may be supplied from the air hole 14 as the sample feed port 13 functions as an air hole. In this case, if the air hole 14 is formed so that at least a part of the reaction layer comes immediately below the air hole 14, the lecithin layer 8 can be made smaller.
one minute after the sample liquid is fed to the glucose sensor, +0.5 V pulse voltage is impressed between the working electrode and the counter electrode. An oxidization current five seconds later is measured.
When the sample liquid is supplied to the glucose sensor, the reaction layer 7 is dissolved into the sample liquid, and glucose in the sample liquid is oxidized by the GOD. At this time, potassium ferricyanide is reduced to potassium ferrocyanide by the transferred electrons, and the oxidization current based on the concentration of the generated potassium ferrocyanide is allowed to flow because of the impression of the pulse voltage. The current is proportional to the concentration of the substrate, namely, glucose.
When the response characteristic of the glucose sensor of the instant embodiment is measured, favourable linearity is confirmed up to a concentration not smaller than 900 mg/d~
(0.05 mol/~). Moreover, a coefficient of variation (CV value) is as good as approximately 2% when 50 glucose sensors are used.
(Embodiment 2) Fig. 3 is a plan view of a base when the counter and working electrodes are formed circular.
On an insulating base plate I formed of polyethylene terephthalate, there are formed leads 2, 3, a working electrode 4, an insulating layer 6 and a counter electrode 5 in the same manner as in Embodiment 1 above through screen printing. The base shown in Fig. 3 is thus obtained.
A 0.5 wt% CMC aqueous solution is spread and dried on an electrode system of the working and counter electrodes 4, 5, constituting a CMC layer. GOD after being dissolved in a phosphate buffer solution (PH=5.6) is dropped onto the CMC
layer and dried to form a CMC-GOD layer. Moreover, a solution of 0.5 wt% polyvinyl pyrrolidone (referred to as PVP
hereinafter) in ethanol as a hydrophilic polymer is spread and dried to form a PVP layer. Potassium ferricyanide of minute crystals is mixed in a solution of 0.5% lecithin in toluene as a dispersant, which is dropped and dried on the PVP layer ~ 8 - 2068475 thereby to form a potassium ferricyanide-lecithin layer. In this manner, a reaction layer is formed.
Then, a toluene solution in lecithin is spread as a surfactant from a part on the base plate corresponding to a sample feed port onto the reaction layer and dried, thus constituting a lecithin layer.
The insulating base plate is formed into one body with a spacer and a cover in the same manner as in Embodiment 1.
Accordingly, a biosensor according to the second embodiment is obtained.
Similar to Embodiment 1, if an air hole is formed at such a position that at least a part of the reaction layer comes immediately below the air hole, the lecithin layer may be made smaller. It is preferable in this case to supply the sample liquid from the air hole, as in Embodiment 1.
A 3~e sample of whole blood from a man is supplied as a sample liquid to the glucose sensor of the above-described structure. One minute later, +0.5 V pulse voltage is applied between the working electrode and the counter electrode. When an oxidization current five minutes later is obtained, good linearity is observed up to a glucose concentration in the whole blood of not lower than 450 mg/de (0.025 mol/Q).
Moreover, when 30 sensors are used for the same sample, a 2%
or lower coefficient of variation (CV value) is measured.
A comparative glucose sensor having a square working electrode with the same area as in Fig. 2 is manufactured in the same fashion. However, the coefficient of variation to the whole blood is about 4% when 30 sensors are used.
Therefore, if the working electrode is made circular without a vertex, higher accuracy is achieved in measurement.
(Embodiment 3) Fig. 4 is a plan view of a base of a glucose sensor according to a third embodiment of the present invention, and Fig. 5 is an exploded perspective view wherein a reaction layer is removed from the glucose sensor of Fig. 4.
Silver paste is printed by screen printing on an insulating base plate 1 of polyethylene terephthalate, so that leads 2, 3 are formed.
Then, conductive carbon paste including a resin binder is printed to form a working electrode 4 and a counter electrode 5. The working and counter electrodes 4, 5 are kept in contact with the leads 2, 3, respectively.
Insulating paste is then printed to form an insulating layer 6.
The insulating layer 6 covers the outer periphery of the working and counter electrodes 4, 5. The exposed area of the working electrode 4 is maintained constant. Moreover, the unnecessary part of the leads 2, 3, that is, all except for the part connecting with the measuring device is covered with the insulating layer 6. The glucose sensor is connected to the measuring device at the end parts of the leads 2, 3. The measuring device functions to impress a constant voltage to an electrode system which will be described later and to read an oxidization current, etc.
According to the third embodiment, the working electrode 4 and counter electrode 5 are manufactured in one printing process, so that the number of manufacturing processes of the glucose sensor is reduced.
The counter electrode 5 is mainly circular, but a part thereof is notched as shown in Fig. 4. The insulating layer 6 is present in the cut part, and the lead 2 in contact with the working electrode 4 is provided below the insulating layer 6.
Accordingly, even if a pin hole is found on the part of the insulating layer 6 corresponding to the cut part, the counter electrode 5 is prevented from being shortcircuited with the lead 2.
According to the manufacturing method described above, a large quantity of disposable glucose sensors can be manufactured at low cost with good yield.
A 0.25 wt% CMC aqueous solution is spread and dried on the electrode system comprised of the working and counter electrodes 4, 5, so that a CMC layer is formed.
As mentioned before, the insulating layer 6 is present at the cut part of the counter electrode S. The surface of the insulating layer is water-repellent in the instant embodiment.
Therefore, if the area of the cut part is large, the CMC
aqueous solution is repelled by the insulating layer, thereby making it impossible to form a uniform CMC layer on the electrode system. For the above reason, it is desirable to set the distance or length of the cut part of the electrode 5 to be not larger than 2 mm, preferably not larger than 1.6 mm.
A mixed solution obtained by dissolving GOD and potassium ferricyanide in 0.25 wt% CMC aqueous solution is dropped onto the CMC layer, and dried for ten minutes in a warm drier at 50C. A reaction layer is formed.
A toluene solution in lecithin as a surfactant is spread and dried onto the reaction layer from a part of the base plate 1 corresponding to a sample feed port 13, thereby forming a lecithin layer.
The insulating base plate 1 is bonded to a spacer 11 and a cover 12 as indicated by the broken line in Fig. 5.
When the sample liquid is brought in contact with the sample feed port 13, the sample liquid contacts the lecithin layer as well. Accordingly, the sample liquid can be smoothly introduced onto the reaction layer by the lecithin layer.
As in Embodiment 1, the sample feed port 13 is not necessarily distinguished from the air hole 14.
Moreover, if the air hole is formed so that at least a part of the reaction layer is immediately below the air hole, the lecithin layer may be made smaller. In a case as above, similar to Embodiment 1, the sample liquid can preferably be supplied from the air hole.
When the glucose concentration in a glucose standard solution is measured using the glucose sensor manufactured in the above-described manner, a response characteristic with good reproducibility is obtained.
~ - 11 2068475 (Embodiment 4) On a base plate 1 which is insulating and formed of polyethylene terephthalate, leads 2, 3 are formed by printing silver paste through screen printing.
Insulating paste is further printed to form an insulating layer 6.
The insulating layer 6 covers the unnecessary parts of the leads 2, 3. The unnecessary part of the lead 2 is all of the lead except for the part connecting with a working electrode 4 layer (to be described) and the part connecting with a measuring device. On the other hand, the unnecessary part of the lead 3 is all of the lead except for the part connecting with a counter electrode 5 which will be discussed later and the part connecting with the measuring device.
Thereafter, conductive carbon paste including a resin binder is printed thereby forming the working electrode 4 and counter electrode 5. The working electrode 4 is in contact with the lead 2, while the counter electrode 5 is kept in contact with the lead 3. In the manner as above, a base of Fig. 4 is manufactured.
According to this Embodiment 4, both the working electrode 4 and counter electrode 5 are formed in one printing process, so that the number of manufacturing processes for the base can be reduced.
Furthermore, according to Embodiment 4, since the distance between the electrodes 4 and 5 is determined by a screen form plate used in the third step of the manufacturing process for the base, the relative position of the electrodes 4, 5 is not changed even if the printing position is shifted when the insulating layer is formed in the second step.
Accordingly, it is possible to manufacture a mass of disposable glucose sensors at low cost with good yield.
Then, a reaction layer and a lecithin layer are formed in the same manner as in Embodiment 3, and the obtained base plate is integrally formed with a cover and a spacer.
A glucose sensor of the fourth embodiment is thus provided.
.~
~ - 12 - 2068475 When the glucose concentration is measured in the same manner as in Embodiment 3, response with good reproducibility is obtained.
(Embodiment 5) A base shown in Fig. 6 is formed through screen printing in the same manner as in Embodiment 3.
When a working electrode 4 and a counter electrode 5 are in the shape as shown in Fig. 6, the following effects are realized:
The number of manufacturing processes can be reduced as the electrodes 4, 5 are printed simultaneously;
a uniform reaction layer can be formed on an electrode system of the working and counter electrodes 4, 5 if the electrodes 4, 5 are made of the same material; and a lead 2 which is connected to the working electrode 4 is prevented from being shortcircuited with the counter electrode 5.
The glucose sensor of Embodiment 5 is formed in the same manner in the succeeding steps as in the third embodiment, which shows a response characteristic with good reproducibility in measuring glucose concentration.
(Embodiment 6) The base shown in Fig. 7 is obtained in the same manner as in Embodiment 3 through screen printing.
Since a working electrode 4 and a counter electrode 5 are formed in the shape as indicated in Fig. 7, the following effects are achieved:
The number of manufacturing processes can be reduced as the working electrode 4 and counter electrode 5 are printed at the same time;
a uniform reaction layer is formed on an electrode system comprised of the working and counter electrodes 4, 5;
shortcircuiting between the lead 2 connected to the working electrode 4 and counter electrode 5 can be prevented;
since the distance between the working electrode 4 and counter electrode 5 is held constant, and no vertex is present j '.
,, ~
~ - 13 - 2068475 in the working electrode 4, highly accurate measurement is realized.
Subsequently, the glucose sensor of Embodiment 6 is formed in the same manner as in Embodiment 3. When glucose concentration is measured with the glucose sensor, the glucose sensor assumes a response characteristic with good reproducibility.
(Embodiment 7) A base is formed as shown in Fig. 8 through screen printing in the same manner as in Embodiment 4.
Since a working electrode 4 and a counter electrode 5 are formed in the shape as shown in Fig. 8, a glucose sensor of Embodiment 7 enjoys the following effects:
The number of manufacturing processes can be reduced by printing the working electrode 4 and counter electrode 5 simultaneously;
a uniform reaction layer can be formed on an electrode system comprised of the working and counter electrodes 4 and 5 if the electrodes 4, 5 are made of the same material;
a lead connected to the working electrode 4 can be prevented from being shortcircuited with the counter electrode 5;
the distance from the working electrode 4 to the counter electrode 5 is kept constant, and the working electrode 4 has no vertex, whereby the concentration can be measured with higher accuracy.
In addition to the above, a further effect can be obtained:
Since the working and counter electrodes 4, 5 are formed on an insulating layer 6, a step difference is generated at the portion of the insulating layer held between the electrodes 5 and 4. A sample liquid is fed from a sample feed port, and does not intersect the step difference until the liquid reaches the working electrode. As a result, the reaction layer can be uniformly dissolved into the sample liquid, thereby improving the measuring accuracy.
The glucose sensor is manufactured subsequently in the same manner as in Embodiment 4, which carries a response characteristic with good reproducibility when measuring glucose concentration.
If the electrode system of two electrodes, namely, working electrode 4 and counter electrode S is changed to an electrode system of three electrodes wherein a reference electrode 20 and a lead 21 are added as shown in Fig. 9, the concentration can be measured more accurately.
In the foregoing embodiments, even when the working electrode or counter electrode is changed from circular to elliptical, the same effects are achieved. The electrode may not be always a true circle or a true ellipse, and can be deformed from a circle or an ellipse so long as the above effects are ensured.
Although the foregoing embodiments are all related to a glucose sensor, the present invention is widely applicable to any enzyme-related system, e.g. a sucrose sensor, a fructose sensor, an alcohol sensor, a lactic acid sensor, a cholesterol sensor, an amino acid sensor, etc.
Further, in place of glucose oxidase used as the enzyme in the above embodiments, invertase, mutarotase, fructose dehydrogenase, alcohol oxidase, lactate oxidase, lactate dehydrogenase, cholesterol oxidase, amino acid oxidase, xanthine oxidase or the like may be employed.
Carboxymethyl cellulose and polyvinyl pyrrolidone are used as the hydrophilic polymer in the above embodiments.
However, it is not restricted to these, but polyvinyl alcohol, gelatine and its derivative, acrylic acid and its salt, methacrylic acid and its salt, starch and its derivative, maleic anhydride and its salt, and cellulose derivative may be used. More specifically, there are, for cellulose derivatives, hydroxypropyl cellulose, hydroxyethyl cellulose, methyl cellulose, ethyl cellulose, ethyl hydroxyethyl 3S cellulose, carboxymethyl ethyl cellulose.
According to Embodiment 2, lecithin is used as a dispersant. However, the dispersant is not particularly restricted to lecithin so long as it does not influence the activity of the enzyme. For instance, polyoxyethylene alkyl ether, polyethylene glycol fatty ester, oleic acid, polyoxyethylene glycerin fatty acid, cyclodextrin, etc. may be employed.
Moreover, as an electron acceptor, p-benzoquinone, phenazine methosulfate, ferrocene and the like can be used instead of potassium ferricyanide in the foregoing embodiments.
As is clearly described hereinabove, the biosensor of the present invention is capable of measuring a specific component in various kinds of samples quickly and accurately with ease.
Moreover, the manufacturing method of the present invention enables mass production of biosensors with high preservative properties and reliability at low cost. The manufacturing method is very useful.
Although the present invention has been fully described in connection with the preferred embodiments thereof with reference to the accompanying drawings, it is to be noted that various changes and modifications would be apparent to those skilled in the art. Such changes and modifications are to be understood to be included within the scope of the present invention as defined by the appended claims unless they depart therefrom.
`~
The present invention relates to a biosensor designed to quickly quantify with ease and high accuracy a specific component in various kinds of sample liquids, and a method for manufacturing the biosensor.
Various kinds of biosensors utilizing the specific catalytic effect of an enzyme have been developed in the past years, but a more accurate one is desired.
An example of a biosensor of the aforementioned type, namely, a glucose sensor applied in clinical technology will be described below.
Conventionally, blood plasma obtained from the blood of a patient after treatment in a centrifuge has been measured in order to quantify glucose in the blood. This method requires time and labour. Therefore, a sensor that can measure glucose concentration in the blood using whole blood is desirable.
As a simple glucose sensor, one similar to a test paper used for inspection of urea has been provided. This glucose sensor has a stick-like supporting body and a carrier fixed to the supporting body which includes an enzyme reactive only to glucose and colouring matter which will change colour as a result of the enzyme reaction. When blood is dropped onto the carrier of the sensor and the change in the colouring matter after a predetermined time is visually measured with or without optical means, the amount of glucose contained in the blood can be determined. According to the quantifying method using the glucose sensor of this type, however, the result is greatly influenced by colouring matter in the blood, and therefore the measuring accuracy is low.
In the meantime, according to a different method, a specific component of a living sample, e.g. blood or the like, can be quantified with high accuracy without requiring dilution or stirring of the sample liquid. An example of a biosensor using this method is proposed in Japanese Patent Laid-open Publication Tokkaihei 1-212345 (212345/1989), which will be discussed below.
~ - 2 - 2068475 The biosensor has an electrode system formed on an insulating base plate by screen printing or the like, and an enzyme reaction layer which consists of a hydrophilic polymer layer, a oxidoreductase and an electron acceptor is formed on the electrode system.
The biosensor of the above structure operates as follows.
When a sample liquid is dropped on the enzyme reaction layer, the oxidoreductase and electron acceptor are dissolved in the sample liquid, so that the enzyme reaction progresses with a substrate in the sample liquid. As a result, the electron acceptor is reduced. After the enzyme reaction is complete, the reduced electron acceptor is electrochemically oxidized.
The concentration of the substrate in the sample liquid is obtained from an oxidization current obtained at this time.
In the conventional biosensor, the working electrode and counter electrode are formed of material different from that of the insulating layer. Therefore, the reaction layer on the electrode system mainly consisting of the working electrode and counter electrode is formed in conformity with the configuration of the counter electrode. Since the counter electrode of the conventional biosensor is polygonal and mainly square, the reaction layer is apt to be separated from a corner of the counter electrode.
In a preferred embodiment the present invention provides a biosensor for measuring a substrate in a liquid sample, which is provided with an insulating base plate, an electrode system mainly consisting of a working electrode and a counter electrode formed on said base plate, and a reaction layer formed on said electrode system area and comprising an enzyme and an electron acceptor, wherein the reduced electron acceptor caused by means of a reaction between the enzyme and the said substrate in the liquid sample is electrochemically oxidized and the resulting value of oxidation current between the working electrode and the counter electrode is measured to determine the concentration of the substrate, characterized in that the counter electrode is provided with a circular arc portion at the outer edge of the counter electrode against the ~ .i~
direction of sample supply containing the substrate to be measured. More preferably, the counter electrode is provided with an outer edge which is partially broken.
A biosensor according to a further aspect of the present invention is characterized in that the working electrode is provided with an outer edge of circular, elliptical, deformed from the circular shape, or mainly polygonal shape, the vertex of which assumes a curve.
A biosensor according to yet a further aspect of the present invention is characterized in that the distance between the counter electrode and working electrode is held equal.
The present invention further features a manufacturing method of the biosensor, which is comprised of a manufacturing step for forming a base by providing leads, an electrode system consisting mainly of a working electrode and a counter electrode and, then an insulating layer are formed on an insulating base plate, and a forming step whereby a reaction layer mainly composed of an enzyme is formed onto the working electrode. The manufacturing step of the base includes the following three processes: a process to form the leads on the insulating base plate; a process to form the electrode system consisting mainly of the working electrode and counter electrode onto the insulating base plate; and a process to form the insulating layer on the insulating base plate.
The disadvantages inherent in the prior art can be solved by the biosensor of the present invention.
Moreover, the following effects can be achieved by the aforementioned means. In the first place, since the counter electrode is formed circular, elliptical or deformed from the circular shape, formation of the reaction layer becomes easy and delamination thereof can be prevented.
Further, since the working electrode is formed circular, elliptical, deformed from the circular shape, or mainly polygonal with a vertex thereof turned round by a curve, the density distribution of a current between the working and counter electrodes is made uniform. This is further enhanced _ 4 _ 2068475 by keeping the distance of the counter electrode from the working electrode equal. In consequence, the concentration of the substrate can be measured with higher accuracy.
Likewise, if a part of the circular or deformed counter electrode is removed, the lead part of the working electrode is prevented from being shortcircuited with the counter electrode by the presence of a pin hole in the insulating layer, etc.
According to the manufacturing method of the present invention, in the case where the electrode system is made of the working and counter electrodes, that is, the electrode system is a system of two electrodes, the working and counter electrodes can be manufactured in one process although they have been conventionally formed in different processes, thereby reducing the number of manufacturing processes of the biosensor .
These and other objects and features of the present invention will become clear from the following description taken in conjunction with the preferred embodiments thereof with reference to the accompanying drawings throughout which like parts are designated by like reference numerals, and in which:
Fig. 1 is a cross sectional view of a glucose sensor without a cover and a spacer according to one embodiment of the present invention;
Fig. 2 is an exploded perspective view with a reaction layer removed from the glucose sensor of Fig. l;
Fig. 3 is a plan view of a base of a glucose sensor according to a further embodiment of the present invention;
Fig. 4 is a plan view of a base of a glucose sensor according to a still further embodiment of the present invention;
Fig. 5 is an exploded perspective view with a reaction layer removed from the glucose sensor of Fig. 4; and Figs. 6, 7, 8, and 9 are plan views of a base of a glucose sensor according to still different embodiments of the present invention.
~ ~ 5 ~ 2068475 (Embodiment 1) A glucose sensor will be described hereinbelow as one example of a biosensor of the present invention.
Fig. 1 is a cross sectional view of a glucose sensor without a cover and a spacer according to one embodiment of the present invention, and Fig. 2 is an exploded perspective view with a reaction layer removed from the glucose sensor.
Referring to these drawings, silver paste is printed on an insulating base plate 1 of polyethylene terephthalate through screen printing, thereby forming leads 2, 3.
Conductive carbon paste including a resin binder is further printed in contact with the lead 2, thereby forming a working electrode 4. Then, insulating paste is printed to form an insulating layer 6.
The insulating layer 6 covers the outer periphery of the working electrode 4, so that the area of the exposed part of the working electrode 4 is held constant. At the same time, the insulating layer 6 covers the unnecessary part of the leads 2, 3. More specifically, the unnecessary part of the lead 2 all but the part of the lead 2 connecting with a working device. The unnecessary part of the lead 3 is all but the part of the lead 3 connecting with a counter electrode 5 which will be described later and the part connecting ~ith the measuring device. The glucose sensor is connected to the measuring device via end parts of the leads 2, 3. The measuring device impresses a constant voltage to an electrode system which will be described later, reads the value of an oxidization current, while performing a like function.
The counter electrode 5 is formed by printing conductive carbon paste including a resin binder into contact with the lead 3, which is circular in the outer periphery thereof.
A 0.5 wt% solution of carboxymethyl cellulose (referred to as CMC hereinafter) as a hydrophilic polymer in water is spread onto the electrode system consisting of the working electrode 4 and counter electrode 5 and dried, so that a CMC
layer is formed.
-The surface of the insulating layer according to the instant embodiment is water-repellent, and therefore the CMC
aqueous solution is repelled by the insulating layer and spreads only over the electrode system of the working and counter electrodes 4, 5. If the counter electrode 5 is formed circular as illustrated in Fig. 2, it is possible to form a uniform CMC layer. Moreover, delamination of the layer which has sometimes occurred after drying when the counter electrode is square can be prevented.
Subsequently, a mixed solution obtained by dissolving glucose oxidase (referred to as GOD hereinafter) as an enzyme and potassium ferricyanide as an electron acceptor in 0.5 wt%
CMC aqueous solution is dropped onto the CMC layer, dried for ten minutes in a warm drier at 50C. As a result, a reaction layer 7 is formed.
A solution of lecithin in toluene as a surface-tension surfactant is spread from a part of the base plate 1 corresponding to a sample feed port 13 onto the reaction layer 7 and dried, and a lecithin layer 8 is formed.
The insulating base plate 1, a spacer 11 and a cover 12 are bonded to each other in the position as indicated by the broken chain line in Fig. 2.
A 3~ glucose standard solution is supplied as a sample liquid from the sample feed port 13 to the glucose sensor formed in the above-described manner.
When the sample liquid is brought in contact with the sample feed port 13, it is also in contact with the lecithin layer 8 and is smoothly guided onto the reaction layer by the lecithin layer.
It is not always necessary to distinguish the sample feed port 13 from an air hole 14. The sample liquid may be supplied from the air hole 14 as the sample feed port 13 functions as an air hole. In this case, if the air hole 14 is formed so that at least a part of the reaction layer comes immediately below the air hole 14, the lecithin layer 8 can be made smaller.
one minute after the sample liquid is fed to the glucose sensor, +0.5 V pulse voltage is impressed between the working electrode and the counter electrode. An oxidization current five seconds later is measured.
When the sample liquid is supplied to the glucose sensor, the reaction layer 7 is dissolved into the sample liquid, and glucose in the sample liquid is oxidized by the GOD. At this time, potassium ferricyanide is reduced to potassium ferrocyanide by the transferred electrons, and the oxidization current based on the concentration of the generated potassium ferrocyanide is allowed to flow because of the impression of the pulse voltage. The current is proportional to the concentration of the substrate, namely, glucose.
When the response characteristic of the glucose sensor of the instant embodiment is measured, favourable linearity is confirmed up to a concentration not smaller than 900 mg/d~
(0.05 mol/~). Moreover, a coefficient of variation (CV value) is as good as approximately 2% when 50 glucose sensors are used.
(Embodiment 2) Fig. 3 is a plan view of a base when the counter and working electrodes are formed circular.
On an insulating base plate I formed of polyethylene terephthalate, there are formed leads 2, 3, a working electrode 4, an insulating layer 6 and a counter electrode 5 in the same manner as in Embodiment 1 above through screen printing. The base shown in Fig. 3 is thus obtained.
A 0.5 wt% CMC aqueous solution is spread and dried on an electrode system of the working and counter electrodes 4, 5, constituting a CMC layer. GOD after being dissolved in a phosphate buffer solution (PH=5.6) is dropped onto the CMC
layer and dried to form a CMC-GOD layer. Moreover, a solution of 0.5 wt% polyvinyl pyrrolidone (referred to as PVP
hereinafter) in ethanol as a hydrophilic polymer is spread and dried to form a PVP layer. Potassium ferricyanide of minute crystals is mixed in a solution of 0.5% lecithin in toluene as a dispersant, which is dropped and dried on the PVP layer ~ 8 - 2068475 thereby to form a potassium ferricyanide-lecithin layer. In this manner, a reaction layer is formed.
Then, a toluene solution in lecithin is spread as a surfactant from a part on the base plate corresponding to a sample feed port onto the reaction layer and dried, thus constituting a lecithin layer.
The insulating base plate is formed into one body with a spacer and a cover in the same manner as in Embodiment 1.
Accordingly, a biosensor according to the second embodiment is obtained.
Similar to Embodiment 1, if an air hole is formed at such a position that at least a part of the reaction layer comes immediately below the air hole, the lecithin layer may be made smaller. It is preferable in this case to supply the sample liquid from the air hole, as in Embodiment 1.
A 3~e sample of whole blood from a man is supplied as a sample liquid to the glucose sensor of the above-described structure. One minute later, +0.5 V pulse voltage is applied between the working electrode and the counter electrode. When an oxidization current five minutes later is obtained, good linearity is observed up to a glucose concentration in the whole blood of not lower than 450 mg/de (0.025 mol/Q).
Moreover, when 30 sensors are used for the same sample, a 2%
or lower coefficient of variation (CV value) is measured.
A comparative glucose sensor having a square working electrode with the same area as in Fig. 2 is manufactured in the same fashion. However, the coefficient of variation to the whole blood is about 4% when 30 sensors are used.
Therefore, if the working electrode is made circular without a vertex, higher accuracy is achieved in measurement.
(Embodiment 3) Fig. 4 is a plan view of a base of a glucose sensor according to a third embodiment of the present invention, and Fig. 5 is an exploded perspective view wherein a reaction layer is removed from the glucose sensor of Fig. 4.
Silver paste is printed by screen printing on an insulating base plate 1 of polyethylene terephthalate, so that leads 2, 3 are formed.
Then, conductive carbon paste including a resin binder is printed to form a working electrode 4 and a counter electrode 5. The working and counter electrodes 4, 5 are kept in contact with the leads 2, 3, respectively.
Insulating paste is then printed to form an insulating layer 6.
The insulating layer 6 covers the outer periphery of the working and counter electrodes 4, 5. The exposed area of the working electrode 4 is maintained constant. Moreover, the unnecessary part of the leads 2, 3, that is, all except for the part connecting with the measuring device is covered with the insulating layer 6. The glucose sensor is connected to the measuring device at the end parts of the leads 2, 3. The measuring device functions to impress a constant voltage to an electrode system which will be described later and to read an oxidization current, etc.
According to the third embodiment, the working electrode 4 and counter electrode 5 are manufactured in one printing process, so that the number of manufacturing processes of the glucose sensor is reduced.
The counter electrode 5 is mainly circular, but a part thereof is notched as shown in Fig. 4. The insulating layer 6 is present in the cut part, and the lead 2 in contact with the working electrode 4 is provided below the insulating layer 6.
Accordingly, even if a pin hole is found on the part of the insulating layer 6 corresponding to the cut part, the counter electrode 5 is prevented from being shortcircuited with the lead 2.
According to the manufacturing method described above, a large quantity of disposable glucose sensors can be manufactured at low cost with good yield.
A 0.25 wt% CMC aqueous solution is spread and dried on the electrode system comprised of the working and counter electrodes 4, 5, so that a CMC layer is formed.
As mentioned before, the insulating layer 6 is present at the cut part of the counter electrode S. The surface of the insulating layer is water-repellent in the instant embodiment.
Therefore, if the area of the cut part is large, the CMC
aqueous solution is repelled by the insulating layer, thereby making it impossible to form a uniform CMC layer on the electrode system. For the above reason, it is desirable to set the distance or length of the cut part of the electrode 5 to be not larger than 2 mm, preferably not larger than 1.6 mm.
A mixed solution obtained by dissolving GOD and potassium ferricyanide in 0.25 wt% CMC aqueous solution is dropped onto the CMC layer, and dried for ten minutes in a warm drier at 50C. A reaction layer is formed.
A toluene solution in lecithin as a surfactant is spread and dried onto the reaction layer from a part of the base plate 1 corresponding to a sample feed port 13, thereby forming a lecithin layer.
The insulating base plate 1 is bonded to a spacer 11 and a cover 12 as indicated by the broken line in Fig. 5.
When the sample liquid is brought in contact with the sample feed port 13, the sample liquid contacts the lecithin layer as well. Accordingly, the sample liquid can be smoothly introduced onto the reaction layer by the lecithin layer.
As in Embodiment 1, the sample feed port 13 is not necessarily distinguished from the air hole 14.
Moreover, if the air hole is formed so that at least a part of the reaction layer is immediately below the air hole, the lecithin layer may be made smaller. In a case as above, similar to Embodiment 1, the sample liquid can preferably be supplied from the air hole.
When the glucose concentration in a glucose standard solution is measured using the glucose sensor manufactured in the above-described manner, a response characteristic with good reproducibility is obtained.
~ - 11 2068475 (Embodiment 4) On a base plate 1 which is insulating and formed of polyethylene terephthalate, leads 2, 3 are formed by printing silver paste through screen printing.
Insulating paste is further printed to form an insulating layer 6.
The insulating layer 6 covers the unnecessary parts of the leads 2, 3. The unnecessary part of the lead 2 is all of the lead except for the part connecting with a working electrode 4 layer (to be described) and the part connecting with a measuring device. On the other hand, the unnecessary part of the lead 3 is all of the lead except for the part connecting with a counter electrode 5 which will be discussed later and the part connecting with the measuring device.
Thereafter, conductive carbon paste including a resin binder is printed thereby forming the working electrode 4 and counter electrode 5. The working electrode 4 is in contact with the lead 2, while the counter electrode 5 is kept in contact with the lead 3. In the manner as above, a base of Fig. 4 is manufactured.
According to this Embodiment 4, both the working electrode 4 and counter electrode 5 are formed in one printing process, so that the number of manufacturing processes for the base can be reduced.
Furthermore, according to Embodiment 4, since the distance between the electrodes 4 and 5 is determined by a screen form plate used in the third step of the manufacturing process for the base, the relative position of the electrodes 4, 5 is not changed even if the printing position is shifted when the insulating layer is formed in the second step.
Accordingly, it is possible to manufacture a mass of disposable glucose sensors at low cost with good yield.
Then, a reaction layer and a lecithin layer are formed in the same manner as in Embodiment 3, and the obtained base plate is integrally formed with a cover and a spacer.
A glucose sensor of the fourth embodiment is thus provided.
.~
~ - 12 - 2068475 When the glucose concentration is measured in the same manner as in Embodiment 3, response with good reproducibility is obtained.
(Embodiment 5) A base shown in Fig. 6 is formed through screen printing in the same manner as in Embodiment 3.
When a working electrode 4 and a counter electrode 5 are in the shape as shown in Fig. 6, the following effects are realized:
The number of manufacturing processes can be reduced as the electrodes 4, 5 are printed simultaneously;
a uniform reaction layer can be formed on an electrode system of the working and counter electrodes 4, 5 if the electrodes 4, 5 are made of the same material; and a lead 2 which is connected to the working electrode 4 is prevented from being shortcircuited with the counter electrode 5.
The glucose sensor of Embodiment 5 is formed in the same manner in the succeeding steps as in the third embodiment, which shows a response characteristic with good reproducibility in measuring glucose concentration.
(Embodiment 6) The base shown in Fig. 7 is obtained in the same manner as in Embodiment 3 through screen printing.
Since a working electrode 4 and a counter electrode 5 are formed in the shape as indicated in Fig. 7, the following effects are achieved:
The number of manufacturing processes can be reduced as the working electrode 4 and counter electrode 5 are printed at the same time;
a uniform reaction layer is formed on an electrode system comprised of the working and counter electrodes 4, 5;
shortcircuiting between the lead 2 connected to the working electrode 4 and counter electrode 5 can be prevented;
since the distance between the working electrode 4 and counter electrode 5 is held constant, and no vertex is present j '.
,, ~
~ - 13 - 2068475 in the working electrode 4, highly accurate measurement is realized.
Subsequently, the glucose sensor of Embodiment 6 is formed in the same manner as in Embodiment 3. When glucose concentration is measured with the glucose sensor, the glucose sensor assumes a response characteristic with good reproducibility.
(Embodiment 7) A base is formed as shown in Fig. 8 through screen printing in the same manner as in Embodiment 4.
Since a working electrode 4 and a counter electrode 5 are formed in the shape as shown in Fig. 8, a glucose sensor of Embodiment 7 enjoys the following effects:
The number of manufacturing processes can be reduced by printing the working electrode 4 and counter electrode 5 simultaneously;
a uniform reaction layer can be formed on an electrode system comprised of the working and counter electrodes 4 and 5 if the electrodes 4, 5 are made of the same material;
a lead connected to the working electrode 4 can be prevented from being shortcircuited with the counter electrode 5;
the distance from the working electrode 4 to the counter electrode 5 is kept constant, and the working electrode 4 has no vertex, whereby the concentration can be measured with higher accuracy.
In addition to the above, a further effect can be obtained:
Since the working and counter electrodes 4, 5 are formed on an insulating layer 6, a step difference is generated at the portion of the insulating layer held between the electrodes 5 and 4. A sample liquid is fed from a sample feed port, and does not intersect the step difference until the liquid reaches the working electrode. As a result, the reaction layer can be uniformly dissolved into the sample liquid, thereby improving the measuring accuracy.
The glucose sensor is manufactured subsequently in the same manner as in Embodiment 4, which carries a response characteristic with good reproducibility when measuring glucose concentration.
If the electrode system of two electrodes, namely, working electrode 4 and counter electrode S is changed to an electrode system of three electrodes wherein a reference electrode 20 and a lead 21 are added as shown in Fig. 9, the concentration can be measured more accurately.
In the foregoing embodiments, even when the working electrode or counter electrode is changed from circular to elliptical, the same effects are achieved. The electrode may not be always a true circle or a true ellipse, and can be deformed from a circle or an ellipse so long as the above effects are ensured.
Although the foregoing embodiments are all related to a glucose sensor, the present invention is widely applicable to any enzyme-related system, e.g. a sucrose sensor, a fructose sensor, an alcohol sensor, a lactic acid sensor, a cholesterol sensor, an amino acid sensor, etc.
Further, in place of glucose oxidase used as the enzyme in the above embodiments, invertase, mutarotase, fructose dehydrogenase, alcohol oxidase, lactate oxidase, lactate dehydrogenase, cholesterol oxidase, amino acid oxidase, xanthine oxidase or the like may be employed.
Carboxymethyl cellulose and polyvinyl pyrrolidone are used as the hydrophilic polymer in the above embodiments.
However, it is not restricted to these, but polyvinyl alcohol, gelatine and its derivative, acrylic acid and its salt, methacrylic acid and its salt, starch and its derivative, maleic anhydride and its salt, and cellulose derivative may be used. More specifically, there are, for cellulose derivatives, hydroxypropyl cellulose, hydroxyethyl cellulose, methyl cellulose, ethyl cellulose, ethyl hydroxyethyl 3S cellulose, carboxymethyl ethyl cellulose.
According to Embodiment 2, lecithin is used as a dispersant. However, the dispersant is not particularly restricted to lecithin so long as it does not influence the activity of the enzyme. For instance, polyoxyethylene alkyl ether, polyethylene glycol fatty ester, oleic acid, polyoxyethylene glycerin fatty acid, cyclodextrin, etc. may be employed.
Moreover, as an electron acceptor, p-benzoquinone, phenazine methosulfate, ferrocene and the like can be used instead of potassium ferricyanide in the foregoing embodiments.
As is clearly described hereinabove, the biosensor of the present invention is capable of measuring a specific component in various kinds of samples quickly and accurately with ease.
Moreover, the manufacturing method of the present invention enables mass production of biosensors with high preservative properties and reliability at low cost. The manufacturing method is very useful.
Although the present invention has been fully described in connection with the preferred embodiments thereof with reference to the accompanying drawings, it is to be noted that various changes and modifications would be apparent to those skilled in the art. Such changes and modifications are to be understood to be included within the scope of the present invention as defined by the appended claims unless they depart therefrom.
`~
Claims (14)
1. A biosensor for measuring a substrate in a liquid sample, which is provided with an insulating base plate, an electrode system mainly consisting of a working electrode and a counter electrode formed on said base plate, and a reaction layer formed on said electrode system area and comprising an enzyme and an electron acceptor, wherein the reduced electron acceptor caused by means of a reaction between the enzyme and the said substrate in the liquid sample is electrochemically oxidized and the resulting value of oxidation current between the working electrode and the counter electrode is measured to determine the concentration of the substrate, characterized in that the counter electrode is provided with a circular arc portion at the outer edge of the counter electrode against the direction of sample supply containing the substrate to be measured.
2. A biosensor according to claim 1, wherein said counter electrode is provided with an outer edge of mainly circular, elliptical or the like shape, which outer edge is partially broken at any other portions not directed to the direction of sample supply containing the substrate to be measured.
3. A biosensor according to claim 1, wherein said counter electrode is provided with an outer edge of circular, elliptical or the like shape.
4. A biosensor according to claim 1, wherein said working electrode is provided with an outer edge of circular, elliptical or the like shape.
5. A biosensor according to claim 1, wherein said working electrode is provided with an outer edge of mainly polygonal, the vertex of which is turned round by a curve.
6. A biosensor according to claim 1, wherein the distance between the working electrode and counter electrode is made equal.
7. A biosensor according to claim 1, wherein said electrode system is formed of material mainly composed of carbon.
8. A biosensor according to claim 1, wherein said reaction layer includes an enzyme, a hydrophilic polymer and an electron acceptor.
9. A method for manufacturing a biosensor for measuring a substrate in a liquid sample, which is provided with an insulating base plate, an electrode system mainly-consisting of a working electrode and a counter electrode formed on said base plate, and a reaction layer formed on said electrode system area and comprising an enzyme and an electron acceptor, wherein the reduced electron acceptor caused by means of a reaction between the enzyme and the said substrate in the liquid sample is electrochemically oxidized and the resulting value of oxidation current between the working electrode and the counter electrode is measured to determine the concentration of the substrate, wherein the counter electrode is provided with a circular arc portion at the outer edge of the counter electrode against the direction of sample supply containing the substrate to be measured, which comprises:
a step for forming a base by providing leads, an electrode system and an insulating layer on the insulating base plate;
a step for forming a reaction layer mainly composed of the enzyme and the electron acceptor on said electrode system, wherein the following three steps are included in said base forming step:
a step to form the leads onto the insulating base plate;
a step to form the electrode system mainly consisting of a working electrode and a counter electrode onto said insulating base plate; and a step to form the insulating layer onto said insulating base plate.
a step for forming a base by providing leads, an electrode system and an insulating layer on the insulating base plate;
a step for forming a reaction layer mainly composed of the enzyme and the electron acceptor on said electrode system, wherein the following three steps are included in said base forming step:
a step to form the leads onto the insulating base plate;
a step to form the electrode system mainly consisting of a working electrode and a counter electrode onto said insulating base plate; and a step to form the insulating layer onto said insulating base plate.
10. A manufacturing method according to claim 9, wherein said three steps are included in said base forming step in the following order:
a step to form the leads onto the insulating base plate;
a step to form the electrode system mainly consisting of a working electrode and a counter electrode onto said insulating base plate; and a step to form the insulating layer onto said insulating base plate.
a step to form the leads onto the insulating base plate;
a step to form the electrode system mainly consisting of a working electrode and a counter electrode onto said insulating base plate; and a step to form the insulating layer onto said insulating base plate.
11. A manufacturing method according to claim 9, wherein said three steps are included in said base forming step in the following order:
a step to form the leads onto the insulating base plate;
a step to form the insulating layer onto said insulating base plate; and a step to form the electrode system mainly consisting of a working electrode and a counter electrode onto said insulating base plate.
a step to form the leads onto the insulating base plate;
a step to form the insulating layer onto said insulating base plate; and a step to form the electrode system mainly consisting of a working electrode and a counter electrode onto said insulating base plate.
12. A manufacturing method according to claim 9, wherein said base is formed through screen printing.
13. A manufacturing method according to claim 9, wherein said counter electrode is provided with an outer edge of mainly circular, elliptical or the like shape, which outer edge is partially broken at any other portion not directed to the direction on sample supply containing the substrate to be measured.
14. A manufacturing method according to claim 9, wherein said reaction layer includes a hydrophilic polymer.
Applications Claiming Priority (2)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| JP04008219A JP3084877B2 (en) | 1992-01-21 | 1992-01-21 | Manufacturing method of glucose sensor |
| JPP04-008219 | 1992-01-21 |
Publications (2)
| Publication Number | Publication Date |
|---|---|
| CA2068475A1 CA2068475A1 (en) | 1993-07-22 |
| CA2068475C true CA2068475C (en) | 1996-02-20 |
Family
ID=11687099
Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CA002068475A Expired - Lifetime CA2068475C (en) | 1992-01-21 | 1992-05-12 | Biosensor and its manufacture |
Country Status (3)
| Country | Link |
|---|---|
| US (1) | US5512159A (en) |
| JP (1) | JP3084877B2 (en) |
| CA (1) | CA2068475C (en) |
Families Citing this family (150)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| US5658443A (en) * | 1993-07-23 | 1997-08-19 | Matsushita Electric Industrial Co., Ltd. | Biosensor and method for producing the same |
| US5651869A (en) * | 1995-02-28 | 1997-07-29 | Matsushita Electric Industrial Co., Ltd. | Biosensor |
| AU5561796A (en) * | 1995-04-21 | 1996-11-07 | Mine Safety Appliances Company | Electrochemical gas sensor for the detection of nitrogen dio xide |
| US6413410B1 (en) * | 1996-06-19 | 2002-07-02 | Lifescan, Inc. | Electrochemical cell |
| AUPP238898A0 (en) * | 1998-03-12 | 1998-04-09 | Usf Filtration And Separations Group Inc. | Heated electrochemical cell |
| US6863801B2 (en) * | 1995-11-16 | 2005-03-08 | Lifescan, Inc. | Electrochemical cell |
| AUPN661995A0 (en) | 1995-11-16 | 1995-12-07 | Memtec America Corporation | Electrochemical cell 2 |
| US5922188A (en) * | 1996-03-12 | 1999-07-13 | Matsushita Electric Industrial Co., Ltd. | Biosensor and method for quantitating biochemical substrate using the same |
| JP3460183B2 (en) * | 1996-12-24 | 2003-10-27 | 松下電器産業株式会社 | Biosensor |
| JP3487396B2 (en) * | 1997-01-31 | 2004-01-19 | 松下電器産業株式会社 | Biosensor and manufacturing method thereof |
| WO1998035225A1 (en) | 1997-02-06 | 1998-08-13 | E. Heller & Company | Small volume in vitro analyte sensor |
| US5759364A (en) * | 1997-05-02 | 1998-06-02 | Bayer Corporation | Electrochemical biosensor |
| US5798031A (en) * | 1997-05-12 | 1998-08-25 | Bayer Corporation | Electrochemical biosensor |
| EP1009851A1 (en) * | 1997-09-05 | 2000-06-21 | Abbott Laboratories | Electrochemical sensor having equalized electrode areas |
| US6071391A (en) | 1997-09-12 | 2000-06-06 | Nok Corporation | Enzyme electrode structure |
| US6117290A (en) * | 1997-09-26 | 2000-09-12 | Pepex Biomedical, Llc | System and method for measuring a bioanalyte such as lactate |
| US6736957B1 (en) * | 1997-10-16 | 2004-05-18 | Abbott Laboratories | Biosensor electrode mediators for regeneration of cofactors and process for using |
| US6036924A (en) | 1997-12-04 | 2000-03-14 | Hewlett-Packard Company | Cassette of lancet cartridges for sampling blood |
| US6103033A (en) | 1998-03-04 | 2000-08-15 | Therasense, Inc. | Process for producing an electrochemical biosensor |
| US6475360B1 (en) | 1998-03-12 | 2002-11-05 | Lifescan, Inc. | Heated electrochemical cell |
| US6878251B2 (en) * | 1998-03-12 | 2005-04-12 | Lifescan, Inc. | Heated electrochemical cell |
| US6391005B1 (en) | 1998-03-30 | 2002-05-21 | Agilent Technologies, Inc. | Apparatus and method for penetration with shaft having a sensor for sensing penetration depth |
| US8974386B2 (en) | 1998-04-30 | 2015-03-10 | Abbott Diabetes Care Inc. | Analyte monitoring device and methods of use |
| US8465425B2 (en) | 1998-04-30 | 2013-06-18 | Abbott Diabetes Care Inc. | Analyte monitoring device and methods of use |
| US8688188B2 (en) | 1998-04-30 | 2014-04-01 | Abbott Diabetes Care Inc. | Analyte monitoring device and methods of use |
| US6175752B1 (en) | 1998-04-30 | 2001-01-16 | Therasense, Inc. | Analyte monitoring device and methods of use |
| US9066695B2 (en) | 1998-04-30 | 2015-06-30 | Abbott Diabetes Care Inc. | Analyte monitoring device and methods of use |
| US8480580B2 (en) | 1998-04-30 | 2013-07-09 | Abbott Diabetes Care Inc. | Analyte monitoring device and methods of use |
| US8346337B2 (en) | 1998-04-30 | 2013-01-01 | Abbott Diabetes Care Inc. | Analyte monitoring device and methods of use |
| US6591125B1 (en) | 2000-06-27 | 2003-07-08 | Therasense, Inc. | Small volume in vitro analyte sensor with diffusible or non-leachable redox mediator |
| US6338790B1 (en) | 1998-10-08 | 2002-01-15 | Therasense, Inc. | Small volume in vitro analyte sensor with diffusible or non-leachable redox mediator |
| US6582583B1 (en) * | 1998-11-30 | 2003-06-24 | The United States Of America As Represented By The Department Of Health And Human Services | Amperometric biomimetic enzyme sensors based on modified cyclodextrin as electrocatalysts |
| US6841052B2 (en) | 1999-08-02 | 2005-01-11 | Bayer Corporation | Electrochemical-sensor design |
| CA2305922C (en) † | 1999-08-02 | 2005-09-20 | Bayer Corporation | Improved electrochemical sensor design |
| US6645359B1 (en) * | 2000-10-06 | 2003-11-11 | Roche Diagnostics Corporation | Biosensor |
| AU1631101A (en) * | 1999-10-07 | 2001-05-10 | Pepex Biomedical, Llc | Sensor for measuring a bioanalyte such as lactate |
| US6464849B1 (en) | 1999-10-07 | 2002-10-15 | Pepex Biomedical, L.L.C. | Sensor for measuring a bioanalyte such as lactate |
| CN100363739C (en) * | 1999-11-15 | 2008-01-23 | 松下电器产业株式会社 | Biosensor, method of forming thin-film electrode, and method and apparatus for quantitative determination |
| DE60029127T2 (en) * | 1999-11-16 | 2007-05-31 | Matsushita Electric Industrial Co., Ltd., Kadoma | BIOSENSOR |
| US6623620B2 (en) * | 1999-11-22 | 2003-09-23 | Hathaway Brown School | Method for detecting or monitoring sulfur dioxide with an electrochemical sensor |
| JP3985417B2 (en) * | 2000-03-08 | 2007-10-03 | 松下電器産業株式会社 | Biosensor and manufacturing method thereof |
| US8641644B2 (en) | 2000-11-21 | 2014-02-04 | Sanofi-Aventis Deutschland Gmbh | Blood testing apparatus having a rotatable cartridge with multiple lancing elements and testing means |
| US6560471B1 (en) | 2001-01-02 | 2003-05-06 | Therasense, Inc. | Analyte monitoring device and methods of use |
| JP4120400B2 (en) * | 2001-04-16 | 2008-07-16 | 松下電器産業株式会社 | Biosensor |
| US7473398B2 (en) * | 2001-05-25 | 2009-01-06 | Roche Diagnostics Operations, Inc. | Biosensor |
| US9795747B2 (en) | 2010-06-02 | 2017-10-24 | Sanofi-Aventis Deutschland Gmbh | Methods and apparatus for lancet actuation |
| US7981056B2 (en) | 2002-04-19 | 2011-07-19 | Pelikan Technologies, Inc. | Methods and apparatus for lancet actuation |
| US9427532B2 (en) | 2001-06-12 | 2016-08-30 | Sanofi-Aventis Deutschland Gmbh | Tissue penetration device |
| EP1404235A4 (en) | 2001-06-12 | 2008-08-20 | Pelikan Technologies Inc | Method and apparatus for lancet launching device integrated onto a blood-sampling cartridge |
| US7025774B2 (en) | 2001-06-12 | 2006-04-11 | Pelikan Technologies, Inc. | Tissue penetration device |
| US9226699B2 (en) | 2002-04-19 | 2016-01-05 | Sanofi-Aventis Deutschland Gmbh | Body fluid sampling module with a continuous compression tissue interface surface |
| US7682318B2 (en) | 2001-06-12 | 2010-03-23 | Pelikan Technologies, Inc. | Blood sampling apparatus and method |
| EP1404233B1 (en) | 2001-06-12 | 2009-12-02 | Pelikan Technologies Inc. | Self optimizing lancing device with adaptation means to temporal variations in cutaneous properties |
| US8337419B2 (en) | 2002-04-19 | 2012-12-25 | Sanofi-Aventis Deutschland Gmbh | Tissue penetration device |
| ATE497731T1 (en) | 2001-06-12 | 2011-02-15 | Pelikan Technologies Inc | DEVICE FOR INCREASING THE SUCCESS RATE OF BLOOD YIELD OBTAINED BY A FINGER PICK |
| JP4149911B2 (en) | 2001-06-12 | 2008-09-17 | ペリカン テクノロジーズ インコーポレイテッド | Electric lancet actuator |
| CA2429360C (en) * | 2001-10-10 | 2012-01-24 | Lifescan, Inc. | Electrochemical cell |
| US6863800B2 (en) * | 2002-02-01 | 2005-03-08 | Abbott Laboratories | Electrochemical biosensor strip for analysis of liquid samples |
| US9795334B2 (en) | 2002-04-19 | 2017-10-24 | Sanofi-Aventis Deutschland Gmbh | Method and apparatus for penetrating tissue |
| US7291117B2 (en) | 2002-04-19 | 2007-11-06 | Pelikan Technologies, Inc. | Method and apparatus for penetrating tissue |
| US7232451B2 (en) | 2002-04-19 | 2007-06-19 | Pelikan Technologies, Inc. | Method and apparatus for penetrating tissue |
| US7976476B2 (en) | 2002-04-19 | 2011-07-12 | Pelikan Technologies, Inc. | Device and method for variable speed lancet |
| US7331931B2 (en) | 2002-04-19 | 2008-02-19 | Pelikan Technologies, Inc. | Method and apparatus for penetrating tissue |
| US8360992B2 (en) | 2002-04-19 | 2013-01-29 | Sanofi-Aventis Deutschland Gmbh | Method and apparatus for penetrating tissue |
| US7297122B2 (en) | 2002-04-19 | 2007-11-20 | Pelikan Technologies, Inc. | Method and apparatus for penetrating tissue |
| US8221334B2 (en) | 2002-04-19 | 2012-07-17 | Sanofi-Aventis Deutschland Gmbh | Method and apparatus for penetrating tissue |
| US9314194B2 (en) | 2002-04-19 | 2016-04-19 | Sanofi-Aventis Deutschland Gmbh | Tissue penetration device |
| US7491178B2 (en) | 2002-04-19 | 2009-02-17 | Pelikan Technologies, Inc. | Method and apparatus for penetrating tissue |
| US8372016B2 (en) | 2002-04-19 | 2013-02-12 | Sanofi-Aventis Deutschland Gmbh | Method and apparatus for body fluid sampling and analyte sensing |
| US7717863B2 (en) | 2002-04-19 | 2010-05-18 | Pelikan Technologies, Inc. | Method and apparatus for penetrating tissue |
| US7175642B2 (en) | 2002-04-19 | 2007-02-13 | Pelikan Technologies, Inc. | Methods and apparatus for lancet actuation |
| US8784335B2 (en) | 2002-04-19 | 2014-07-22 | Sanofi-Aventis Deutschland Gmbh | Body fluid sampling device with a capacitive sensor |
| US7547287B2 (en) | 2002-04-19 | 2009-06-16 | Pelikan Technologies, Inc. | Method and apparatus for penetrating tissue |
| US7901362B2 (en) | 2002-04-19 | 2011-03-08 | Pelikan Technologies, Inc. | Method and apparatus for penetrating tissue |
| US7674232B2 (en) | 2002-04-19 | 2010-03-09 | Pelikan Technologies, Inc. | Method and apparatus for penetrating tissue |
| US9248267B2 (en) | 2002-04-19 | 2016-02-02 | Sanofi-Aventis Deustchland Gmbh | Tissue penetration device |
| US7909778B2 (en) | 2002-04-19 | 2011-03-22 | Pelikan Technologies, Inc. | Method and apparatus for penetrating tissue |
| US8702624B2 (en) | 2006-09-29 | 2014-04-22 | Sanofi-Aventis Deutschland Gmbh | Analyte measurement device with a single shot actuator |
| US7229458B2 (en) | 2002-04-19 | 2007-06-12 | Pelikan Technologies, Inc. | Method and apparatus for penetrating tissue |
| US8267870B2 (en) | 2002-04-19 | 2012-09-18 | Sanofi-Aventis Deutschland Gmbh | Method and apparatus for body fluid sampling with hybrid actuation |
| US7892183B2 (en) | 2002-04-19 | 2011-02-22 | Pelikan Technologies, Inc. | Method and apparatus for body fluid sampling and analyte sensing |
| US7226461B2 (en) | 2002-04-19 | 2007-06-05 | Pelikan Technologies, Inc. | Method and apparatus for a multi-use body fluid sampling device with sterility barrier release |
| US7371247B2 (en) | 2002-04-19 | 2008-05-13 | Pelikan Technologies, Inc | Method and apparatus for penetrating tissue |
| US8579831B2 (en) | 2002-04-19 | 2013-11-12 | Sanofi-Aventis Deutschland Gmbh | Method and apparatus for penetrating tissue |
| US7648468B2 (en) | 2002-04-19 | 2010-01-19 | Pelikon Technologies, Inc. | Method and apparatus for penetrating tissue |
| US7381184B2 (en) | 2002-11-05 | 2008-06-03 | Abbott Diabetes Care Inc. | Sensor inserter assembly |
| US7265881B2 (en) * | 2002-12-20 | 2007-09-04 | Hewlett-Packard Development Company, L.P. | Method and apparatus for measuring assembly and alignment errors in sensor assemblies |
| CN100415877C (en) | 2002-12-24 | 2008-09-03 | 池田食研株式会社 | Coenzyme-binding glucose dehydrogenase |
| US8574895B2 (en) | 2002-12-30 | 2013-11-05 | Sanofi-Aventis Deutschland Gmbh | Method and apparatus using optical techniques to measure analyte levels |
| EP1578262A4 (en) | 2002-12-31 | 2007-12-05 | Therasense Inc | Continuous glucose monitoring system and methods of use |
| GB0303305D0 (en) * | 2003-02-12 | 2003-03-19 | Secr Defence | Apparatus for collecting particles |
| ES2347248T3 (en) | 2003-05-30 | 2010-10-27 | Pelikan Technologies Inc. | PROCEDURE AND APPLIANCE FOR FLUID INJECTION. |
| US7850621B2 (en) | 2003-06-06 | 2010-12-14 | Pelikan Technologies, Inc. | Method and apparatus for body fluid sampling and analyte sensing |
| WO2006001797A1 (en) | 2004-06-14 | 2006-01-05 | Pelikan Technologies, Inc. | Low pain penetrating |
| EP1671096A4 (en) | 2003-09-29 | 2009-09-16 | Pelikan Technologies Inc | Method and apparatus for an improved sample capture device |
| US20050067277A1 (en) * | 2003-09-30 | 2005-03-31 | Pierce Robin D. | Low volume electrochemical biosensor |
| WO2005037095A1 (en) | 2003-10-14 | 2005-04-28 | Pelikan Technologies, Inc. | Method and apparatus for a variable user interface |
| USD902408S1 (en) | 2003-11-05 | 2020-11-17 | Abbott Diabetes Care Inc. | Analyte sensor control unit |
| US7822454B1 (en) | 2005-01-03 | 2010-10-26 | Pelikan Technologies, Inc. | Fluid sampling device with improved analyte detecting member configuration |
| US8668656B2 (en) | 2003-12-31 | 2014-03-11 | Sanofi-Aventis Deutschland Gmbh | Method and apparatus for improving fluidic flow and sample capture |
| CA2554060C (en) | 2004-02-06 | 2013-04-16 | Bayer Healthcare Llc | Electrochemical biosensor |
| WO2005078118A1 (en) | 2004-02-06 | 2005-08-25 | Bayer Healthcare Llc | Oxidizable species as an internal reference for biosensors and method of use |
| CA2556331A1 (en) | 2004-02-17 | 2005-09-29 | Therasense, Inc. | Method and system for providing data communication in continuous glucose monitoring and management system |
| EP1751546A2 (en) | 2004-05-20 | 2007-02-14 | Albatros Technologies GmbH & Co. KG | Printable hydrogel for biosensors |
| WO2005120365A1 (en) | 2004-06-03 | 2005-12-22 | Pelikan Technologies, Inc. | Method and apparatus for a fluid sampling device |
| US9775553B2 (en) | 2004-06-03 | 2017-10-03 | Sanofi-Aventis Deutschland Gmbh | Method and apparatus for a fluid sampling device |
| US20060025666A1 (en) * | 2004-06-10 | 2006-02-02 | Kinesense, Inc. | "Bulls-eye" surface electromyographic electrode assembly |
| US8652831B2 (en) | 2004-12-30 | 2014-02-18 | Sanofi-Aventis Deutschland Gmbh | Method and apparatus for analyte measurement test time |
| EP2365074A1 (en) | 2005-03-25 | 2011-09-14 | Ikeda Food Research Co. Ltd. | Coenzyme-linked glucose dehydrogenase and polynucleotide encoding the same |
| JP4501793B2 (en) * | 2005-06-24 | 2010-07-14 | パナソニック株式会社 | Biosensor |
| AU2006272909B2 (en) | 2005-07-20 | 2013-02-07 | Bayer Healthcare Llc | Gated amperometry |
| AU2006297572B2 (en) | 2005-09-30 | 2012-11-15 | Ascensia Diabetes Care Holdings Ag | Gated Voltammetry |
| US7766829B2 (en) | 2005-11-04 | 2010-08-03 | Abbott Diabetes Care Inc. | Method and system for providing basal profile modification in analyte monitoring and management systems |
| KR100777776B1 (en) * | 2006-03-22 | 2007-11-21 | 주식회사 올메디쿠스 | Working electrode structure of biosensor for reducing measurement error |
| US7620438B2 (en) | 2006-03-31 | 2009-11-17 | Abbott Diabetes Care Inc. | Method and system for powering an electronic device |
| US8226891B2 (en) | 2006-03-31 | 2012-07-24 | Abbott Diabetes Care Inc. | Analyte monitoring devices and methods therefor |
| MX2008014251A (en) | 2006-05-08 | 2008-11-26 | Bayer Healthcare Llc | Electrochemical test sensor with reduced sample volume. |
| US8492130B2 (en) * | 2006-06-29 | 2013-07-23 | Ikeda Food Research Co., Ltd. | FAD-conjugated glucose dehydrogenase gene |
| US7993512B2 (en) | 2006-07-11 | 2011-08-09 | Bayer Healthcare, Llc | Electrochemical test sensor |
| EP2679150B1 (en) | 2006-10-24 | 2020-07-22 | Ascensia Diabetes Care Holdings AG | Transient decay amperometry |
| US8123686B2 (en) | 2007-03-01 | 2012-02-28 | Abbott Diabetes Care Inc. | Method and apparatus for providing rolling data in communication systems |
| WO2009032760A2 (en) | 2007-08-30 | 2009-03-12 | Pepex Biomedical Llc | Electrochmical sensor and method for manufacturing |
| WO2009051901A2 (en) * | 2007-08-30 | 2009-04-23 | Pepex Biomedical, Llc | Electrochemical sensor and method for manufacturing |
| WO2009056299A1 (en) * | 2007-10-31 | 2009-05-07 | Roche Diagnostics Gmbh | Electrical patterns for biosensor and method of making |
| JP5583594B2 (en) * | 2007-12-10 | 2014-09-03 | バイエル・ヘルスケア・エルエルシー | Method of depositing reagent substance in test sensor |
| WO2009076302A1 (en) | 2007-12-10 | 2009-06-18 | Bayer Healthcare Llc | Control markers for auto-detection of control solution and methods of use |
| US8431011B2 (en) * | 2008-01-31 | 2013-04-30 | Abbott Diabetes Care Inc. | Method for automatically and rapidly distinguishing between control and sample solutions in a biosensor strip |
| WO2009126900A1 (en) | 2008-04-11 | 2009-10-15 | Pelikan Technologies, Inc. | Method and apparatus for analyte detecting device |
| US8343427B2 (en) * | 2008-05-26 | 2013-01-01 | Panasonic Corporation | Biosensor |
| JP5405916B2 (en) * | 2008-06-24 | 2014-02-05 | パナソニック株式会社 | Biosensor, method for manufacturing the same, and detection system including the same |
| US9445755B2 (en) | 2008-11-14 | 2016-09-20 | Pepex Biomedical, Llc | Electrochemical sensor module |
| WO2010056876A2 (en) | 2008-11-14 | 2010-05-20 | Pepex Biomedical, Llc | Manufacturing electrochemical sensor module |
| US8951377B2 (en) | 2008-11-14 | 2015-02-10 | Pepex Biomedical, Inc. | Manufacturing electrochemical sensor module |
| US9375169B2 (en) | 2009-01-30 | 2016-06-28 | Sanofi-Aventis Deutschland Gmbh | Cam drive for managing disposable penetrating member actions with a single motor and motor and control system |
| US20100213057A1 (en) | 2009-02-26 | 2010-08-26 | Benjamin Feldman | Self-Powered Analyte Sensor |
| EP3357568A1 (en) | 2009-04-14 | 2018-08-08 | Biocartis NV | Hifu induced cavitation with reduced power threshold |
| ES2822105T3 (en) | 2009-04-15 | 2021-04-29 | Biocartis Nv | Optical detection system to monitor the rtPCR reaction |
| RU2522350C2 (en) | 2009-04-15 | 2014-07-10 | Биокартис Са | Protection of bioanalytical chambers for samples |
| ES2540118T3 (en) | 2009-05-06 | 2015-07-08 | Biocartis Nv | Device for cutting a sample holder |
| US9314195B2 (en) | 2009-08-31 | 2016-04-19 | Abbott Diabetes Care Inc. | Analyte signal processing device and methods |
| US8965476B2 (en) | 2010-04-16 | 2015-02-24 | Sanofi-Aventis Deutschland Gmbh | Tissue penetration device |
| US20120122197A1 (en) * | 2010-11-12 | 2012-05-17 | Abner David Jospeh | Inkjet reagent deposition for biosensor manufacturing |
| EP2698629B1 (en) * | 2011-04-12 | 2016-12-28 | Panasonic Healthcare Holdings Co., Ltd. | Biosensor and measuring device using same |
| WO2012159040A2 (en) | 2011-05-19 | 2012-11-22 | Pepex Biomedical Inc. | Fluid management and patient monitoring system |
| WO2012162151A2 (en) | 2011-05-20 | 2012-11-29 | Pepex Biomedical, Inc. | Manufacturing electrochemical sensor modules |
| EP2925229A4 (en) | 2012-12-03 | 2017-01-25 | Pepex Biomedical, Inc. | Sensor module and method of using a sensor module |
| WO2015016012A1 (en) * | 2013-08-01 | 2015-02-05 | 有限会社清水坂ガラス館 | Earring, and fitting implement for same |
| WO2015187959A1 (en) | 2014-06-04 | 2015-12-10 | Pepex Biomedical, Inc. | Electrochemical sensors and methods for making electrochemical sensors using advanced printing technology |
| EP3156790A1 (en) | 2015-10-15 | 2017-04-19 | ARKRAY, Inc. | Biosensor and manufacturing method of biosensor |
| JP6754259B2 (en) * | 2015-10-15 | 2020-09-09 | アークレイ株式会社 | Biosensor and its manufacturing method |
Family Cites Families (8)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| US4571292A (en) * | 1982-08-12 | 1986-02-18 | Case Western Reserve University | Apparatus for electrochemical measurements |
| WO1986007632A1 (en) * | 1985-06-21 | 1986-12-31 | Matsushita Electric Industrial Co., Ltd. | Biosensor and method of manufacturing same |
| US4851088A (en) * | 1987-03-05 | 1989-07-25 | Honeywell Inc. | Electrochemical detection of carbon dioxide |
| JPH07122624B2 (en) * | 1987-07-06 | 1995-12-25 | ダイキン工業株式会社 | Biosensor |
| US4948490A (en) * | 1988-02-19 | 1990-08-14 | Honeywell Inc. | Tetraalkylammonium ion solid electrolytes |
| JP2502656B2 (en) * | 1988-02-19 | 1996-05-29 | 松下電器産業株式会社 | Biosensor manufacturing method |
| DE68924026T3 (en) * | 1988-03-31 | 2008-01-10 | Matsushita Electric Industrial Co., Ltd., Kadoma | BIOSENSOR AND ITS MANUFACTURE. |
| GB8817421D0 (en) * | 1988-07-21 | 1988-08-24 | Medisense Inc | Bioelectrochemical electrodes |
-
1992
- 1992-01-21 JP JP04008219A patent/JP3084877B2/en not_active Expired - Lifetime
- 1992-05-12 CA CA002068475A patent/CA2068475C/en not_active Expired - Lifetime
-
1994
- 1994-08-19 US US08/293,131 patent/US5512159A/en not_active Expired - Lifetime
Also Published As
| Publication number | Publication date |
|---|---|
| JP3084877B2 (en) | 2000-09-04 |
| JPH05196595A (en) | 1993-08-06 |
| US5512159A (en) | 1996-04-30 |
| CA2068475A1 (en) | 1993-07-22 |
Similar Documents
| Publication | Publication Date | Title |
|---|---|---|
| CA2068475C (en) | Biosensor and its manufacture | |
| US6656702B1 (en) | Biosensor containing glucose dehydrogenase | |
| US5264103A (en) | Biosensor and a method for measuring a concentration of a substrate in a sample | |
| US6077408A (en) | Biosensor and method of manufacturing the same | |
| USRE36991E (en) | Biosensor and method for producing the same | |
| US20010050227A1 (en) | Biosensor | |
| JPH1142098A (en) | Quantitative determination of substrate | |
| JPS63317757A (en) | Glucose sensor | |
| JPH02310457A (en) | Biosensor | |
| JP2960265B2 (en) | Biosensor and measurement method using the same | |
| JP3024394B2 (en) | Biosensor and measurement method using the same | |
| JP3267933B2 (en) | Substrate quantification method | |
| JP3437016B2 (en) | Biosensor and method of quantifying substrate using the same | |
| JP3611268B2 (en) | Biosensor and manufacturing method thereof | |
| JP3070818B2 (en) | Biosensor and manufacturing method thereof | |
| JP2702818B2 (en) | Biosensor and manufacturing method thereof | |
| JPH0652249B2 (en) | Biosensor | |
| JP3370414B2 (en) | Manufacturing method of biosensor | |
| JP3297623B2 (en) | Biosensor | |
| JPH0814562B2 (en) | Biosensor | |
| JPH10227755A (en) | Biosensor | |
| JP3278109B2 (en) | Standard solution for lipid determination | |
| JPH0652248B2 (en) | Biosensor | |
| JP2678087B2 (en) | Biosensor | |
| JP3487064B2 (en) | Substrate quantification method |
Legal Events
| Date | Code | Title | Description |
|---|---|---|---|
| EEER | Examination request | ||
| MKEX | Expiry |