Biosensor and method of manufacturing same

0001This invention relates to a biosensor according to the preamble of Claim 1 and also to a method for producing the same.

0002In recent years, a variety of biosensors have been developed using specific catalytic actions of enzymes, and their applications to the clinical and examining field have been attempted. As inspection items and specimens are now increasing, there is a demand for biosensors which permit rapid and accurate measurements.

0003With a sensor for glucose, for instance, since diabetics are drastically increasing in number, the measurement and control of blood-sugar levels by hitherto employed procedures in which the blood is centrifugated to obtain plasma and subjected to the measurement, takes a relatively long time. Thus, a sensor of the type which enables one to measure it with the whole blood ' is now required. In the simplest form, there is known a sensor which includes, similar to a urine test paper, a support and a carrier formed on the support and containing an enzyme reacting only with glucose and a dye which undergoes a change at the time of the enzyme reaction or with a product formed during the enzyme reaction. The blood is added to the carrier, and a change of the dye after a certain period of time is measured visually or optically. However, this system is disadvantageous in that colored matters in the blood greatly interfere with the measurement, resulting in a low accuracy.

0004A biosensor of the type corresponding to the preamble of claim 1 (shown in Fig. 1) has been described in EP-A-0 136 362 as a system in which a specific component in a biological sample is quantitatively determined in high accuracy. This biosensor includes an insulating substrate 7, a measuring electrode 8 and a counter electrode 9 primarily made of platinum, which have, respectively, leads 10, 11, embedded in the substrate 7, and a perforated body 12 which covers exposed portions of these electrodes and carries an oxido- reductase and an electron acceptor. When the liquid sample is dropped on the perforated body, whereupon the oxido-reductase and the electron acceptor in the perforated body dissolve in the sample liquid, thus permitting the enzyme reaction with a substrate in the liquid sample to proceed. As a result, the electron acceptor is reduced. After completion of the enzyme reaction, the reduced electron acceptor is electrochemically oxidized and a concentration of the substrate in the liquid sample is determined from a current for the oxidation.

0005In this arrangement, although the measurement can be performed simply by replacing the perforated body every measurement, additional procedures, such as washing, are undesirable for the electrode system. If parts including the electrodes could be disposably replaced whenever the measurement is effected, the measuring procedure would become very simple but the system would be very expensive in view of the electrode materials such as platinum and the arrangement. Although the platinum electrodes may be formed by a sputtering method or a vacuum evaporation method, they are still expensive in view of the manufacture.

0006The reference J. Agric. Biol. Chem. volume 49 (2), 541 - 543 (1985) describes the use of a glucose oxidase-immobilized carbon paste electrode as a glucose sensor. According to this reference the enzyme is immobilized on the surface of a carbon paste electrode by coating the surface of the electrode with a nitrocellulose film.

0007It is therefore the object of the present invention to provide a biosensor which measures a specific component in a biologial sample simply, rapidly and in high accuracy and, furthermore, can be handled easily and manufactured simply.

0008According to the invention this object is achieved by providing a biosensor according to the preamble of claim 1 with an electrode system made primarily of carbon and covering the surface of at least the measuring electrode with albumin or glucose oxidase.

0009The invention is advantageously developed by the measures mentioned in the subclaims.

0010Fig. 1 is a schematic view of a known sensor for glucose using an enzyme electrode; Fig. 2 is a schematic view of a biosensor according to one embodiment of the invention; Fig. 3 is a longitudinal section of the biosensor shown in Fig. 2; Figs. 4,5, and 6 are, respectively, response characteristics of the biosensor shown in Fig. 2; Fig. 7 is a schematic view of a biosensor according to another embodiment of the invention; and Fig. 8 is a longitudinal section of the biosensor shown in Pig. 7.

Example 1

0011A sensor for glucose is described as one embodiment of a biosensor. Fig. 2 shows one embodiment of a sensor for glucose and is an exploded view for constituent parts. A conductive carbon paste containing a resin binder is printed, by screen printing, on an insulative base 12 made of polyethylene terephthalate in the form of parallel stripes, followed by heating for drying to form an electrode system consisting of a counter electrode 13, a measuring electrode 14 and a reference electrode 15. The electrode system is partially covered, after which an insulative paste mainly composed of a polyester is printed in the same manner as described above while leaving portions 13', 14' and 15' (each 1 mm² in area) of the respective electrodes acting as electrochemical sites, followed by heating to form an insulative layer 16. Subsequently, the exposed portions 13', 14' and 15' are polished and subjected to heat treatment in air at 100°C for 4 hours.

0012Thereafter, a punched holding frame 17 of a synthetic resin such as a polyester is adhered to the insulative layer 16. A perforated body 18 having an enzyme and an electron acceptor is placed in the hole of the frame 17 so that the electrodes 13', 14' and 15' are completely covered. A resin cover 19 having an opening which is smaller than the outer diameter of the perforated body 18 is adhered, thereby obtaining an integrally assembled sensor. The sectional view of this biosensor taken along the the measuring electrode 14 is shown in Fig. 3. The perforated body used above is fabricated as follows: a nylon non-woven fabric is provided as a base; the base is dipped in a solution of 200 mg of glucose oxidase as an oxydo- reductase and 400 mg of potassium ferricyanide as an electron acceptor dissolved in 1 ml of a phosphate buffer solution containing 0.25 wt% of a surface active agent (polyethylene glycol alkylphenyl ether) and having a pH of 5.6 and immersed in ethanol containing 0.25 wt% of a surface active agent for crystallization; and the thus immersed base is dried under reduced pressure to obtain a perforated material.

0013A glucose standard solution provided as a liquid sample was dropped in the perforated body of the thus fabricated sensor. Two minutes after the dropping, a pulse voltage of 700 mV based on the reference electrode was applied, by which the measuring electrode was polarized anodically.

0014In this case, the added glucose reacts with potassium ferricyanide by the action of the glucose oxidase contained in the perforated body 18 to form potassium ferrocyanide. Upon the application of the pulse voltage for the anodic polarization, a current for the oxidation in proportion to a concentration of the formed potassium ferrocyanide is obtained. This current value corresponds to a concentration of the glucose substrate.

0015Fig. 4 shows the relation between a current obtained 10 seconds after application of the voltage and a glucose concentration as one of response characteristics of the sensor, revealing a very good linearity.

Example 2

0016The procedure for fabricating the sensor for glucose as described in Example 1 was repeated except that the thermal treatment after polishing of the carbon electrodes was effected at 100°C, 70°C, 60°C and 50°C, and that the heat treatment was not effected, thereby fabricating a plurality of sensors for each case. These sensors were kept at 30°Cand their variation in response to the glucose standard solution was determined. Fig. 5 shows a variation of a response current relative to an initial response current taken as 100% for the sensors using the electrodes thermally treated at different temperatures. As will be clearly seen from the figure, the variation of the response accompanied by the storage is not so great when the treating temperature is 60°C or higher. However, such a variation is great for 50°C and also for the non- treatment. This is considered to result from the fact that the exposed surface portion of the polished carbon printed electrodes is not stabilized. When the electrode surface was not polished, a response current was about 1/3 of that of polished electrodes. The difference in response current between the polished and non-polished electrodes is considered to be attributed to partial covering of the carbon surface with a resin component contained as a binder in the paste. Polishing can remove the resin binder from the carbon electrode surface and can uniformly smooth the electrode surface. In addition, the heat treatment at temperatures not lower than 60°C, preferably 60 - 170°C, for 1 - 8 hours permits the the exposed electrode portion to become stabilized.

0017According to our experiments, when the heat treatment was effected at a temperature of 70 - 150°C for 4 hours, good results were obtained in that the variation in response current after storage was very small.

0018Good results cannot be obtained when the heat treatment is effected at temperatures not higher than 50°C as discussed above. On the contrary, the heat treatment at temperatures higher than 170°C should rather be avoided because the polyethylene terephthalate substrate of the sensor tends to undergo thermal deterioration and the resin binder in the carbon paste is apt to deteriorate.

Example 3

0019Similar to the procedure described in Example 1, electrodes were formed on an insulative base and, after polishing, were thermally treated at 100°C for 4 hours. Thereafter, an aqueous solution of albumin (50 mg/ml) was dropped over the surface of the electrode portions 13', 14' and 15' and allowed to stand for 5 - 10 minutes, followed by washing with water to remove an excess of the alubmin and drying. By the above procedure, the albumin was adsorbed on the respective electrode surfaces.

0020After formation of the electrode system partially covered with the albumin, sensors for glucose were made in the same manner as in Example 1.

0021A serum sample containing about 90 mg/dl of glucose was dropped in position of the 10 glucose sensors fabricated above. After 2 minutes, a pulse voltage of 700 mV was applied for measurement in the same manner as in Example 1. Good reproducibility was attained as shown in A in Fig. 7. On the other hand, glucose sensors were fabricated in the same manner as described above but using electrodes not subjected to adsorption with albumin. These sensors were subjected to measurement in the same manner as set forth above. As shown in B in Pig. 7, the variation of the response current is greater than that of A. A and B are both indicative of the response current for 10 glucose sensors fabricated in the same manner, but the difference in reproducibility depending on the adsorption treatment is considered to ascribe to the difference in adsorbability of adsorbates, such as proteins, in the serum sample on the electrodes. As is seen from A, the electrodes adsorbing sufficiently with alubmin can prevent the response current from scattering.

0022Instead of alubmin, an aqueous solution of glucose oxidase (100 mg/ml) was used for the treatment in the same manner as described above, with the result that the response characteristic of a high reproducibility was obtained.

0023If at least measuring electrode among the electrodes is subjected to the adsorption treatment, similar results are obtained.

Example 4

0024In the same manner as in Example 3, an albumin-coated electrode system was formed on an insulative base. Fig. 7 shows an exploded view of a sensor prior to assembling. A liquid-retaining layer 23 made of a perforated rayon non-woven fabric is placed, while controlling the height by means of two resin plates 22 serving as a spacer, in order to cover an electrode system therewith. Then, a filtration layer 21 made of a polycarbonate film and having a pore size of 1 um is mounted on the layer 23 and fixed with a holding frame 17. An enzyme and electron acceptor-bearing perforated body 18 and a sample addition layer 20 made of a cellulose non-woven fabric, both in the form of a disk, are placed in the hole of the holding frame 17. A resin cover having an opening which has a diameter smaller than the outer diameters of the perforated disk body 18 and the sample addition layer is adhered, thereby obtaining an integral combination. A sectional view of the thus integrally combined biosensor taken along the measuring electrode 1 is shown in Fig. 8.

0025In the same manner as in Example 1, glucose oxidase and potassium ferricyanide were incorporated in the perforated disk body 18. Glucose sensors using this perforated body 18 were fabricated. The blood (whole blood) was added to each sensor, whereupon it was rapidly spread over and through the entire surface of the sample addition layer 20. While the enzyme and potassium ferricyanide in the perforated body 18 were being dissolved in and permitted to react with the blood, red cells were filtered by means of the filtration layer 21. The resultant filtrate alone was absorbed in the liquid retaining layer 23, allowing a reaction solution to be collected on the electrode portions 13', 14' and 15' in an amount enough to cause the electrode reaction to proceed. In this manner, the glucose in the blood reacted in the same manner as in Example 1 and a concentration of the glucose could be detected through the electrode system.

0026It will be noted that the technique of integral assembling of a biosensor of the invention is not limited to those shown in the examples with respect to the shape and combination of the frame, the cover and the like.

0027The types of materials for the liquid-retaining layer, the sample addition layer and the filtration layer are not limited to those shown in the examples, but any materials which meet the purposes of the invention may be used.

0028On the other hand, reference has been made to, in the above examples, the three-electrode system, but the measurement would be possible using a two-electrode system consisting of a counter electrode and a measuring electrode.

0029As for the electron acceptor incorporated in the perforated body 18, the potassium ferricyanide used in the examples is convenient because the reaction proceeds stably. On the other hand, p-benzoquinone is suitable for high-speed measurement because of the high reaction rate. Alternatively, 2,6-dichlorophenol indophenol, methylene blue, phenazine methosulfate, potassium beta- naphthoquinone-4-sulfonate and the like may also be used.

0030The sensors described in the examples may be applied not only to glucose, but also to systems relating to oxydo-reductases, for example, as an alcohol sensor, a cholesterol sensor or the like. Glucose oxidase has been used as an oxydo- reductase, but other enzymes such as, for example, alcohol oxidase, xanthine oxidase, cholesterol oxidase and the like may also be used.

0031The biosensor of the invention can be used to measure a specific component in various biological liquid samples rapidly, accurately and simply and has very high utility in clinical examinations.

LIST OF REFERENCE NUMERALS IN THE DRAWINGS:

0032<ul id="ul0001" list-style="none"><li>1 support</li><li>2 reagent layer</li><li>3 developing layer</li><li>4 waterproof layer</li><li>5, 21 filtration layers</li><li>6 small hole</li><li>7 insulative substrate</li><li>8,14,14' measuring electrodes</li><li>9,13,13' counter electrodes</li><li>10, 11 leads</li><li>15,15' reference electrodes</li><li>16 insulative layer</li><li>17 holding frame</li><li>18 perforated body</li><li>19 resin cover</li><li>20 sample-accepting layer</li><li>22 resin plate</li><li>23 liquid-retaining layer</li></ul>