Abstract:
An electrochemical biosensor formed by screen printing and method of fabricating such biosensor is disclosed in the present invention. The biosensor can quickly absorb a sample to be measured therein, effectively control volume of the sample fed and “fill-and-position” the sample therein. The biosensor includes an electrode layer (electrode area) comprising two or three electrodes, which are a working electrode, a reference electrode and an auxiliary electrode (tri-electrode) on an insulating substrate. An active reaction layer containing reactant, reaction catalyst, mediator, wetting agent and surfactant is spread on the surface of the electrode layer. A sample inflow area is formed above the electrode area by adding an upper cover on top of a middle insulating layer with a U-shaped opening formed therein. Sample solution with a minute amount about 0.8 to 1 μl can be rapidly introduced into the electrode area and the active reaction layer via the inflow area by siphon or capillary, where the ingredient of the sample can be analysed by measuring reaction between the sample, reaction catalyst and mediator in the reaction layer using electrochemical potentiometric or amperometric method. An upwardly extended closed space formed within the upper cover above the electrode area adjacent to the front of conductive wires can be effectively used to control sample volume and “fill-and-position” the sample.

Description:
BACKGROUND OF THE INVENTION 
     1. Field of the Invention 
     The present invention relates to an electrochemical biosensor formed by screen printing and a method of fabricating such biosensor. 
     2. Description of the Related Art 
     Recently, electrode sensors have been commercially utilized successfully for the fabrication of a variety of clinical measuring products, such as blood sugar, uric acid and cholesterol measuring devices, for their easy and low cost production processes and wide application of cheaper portable measuring devices. Taking the biggest and the most widely used blood sugar measuring device on market as an example, the leading manufacturers include Roche, Abbott, Bayer and Therasense and all of which fabricate blood sugar sensors by electrochemistry. The first generation of such sensors requires higher amount of blood sample (5–10 μl and above) and takes longer (30–60 seconds) to measure a sample. Hence, they are still not considered ideal although the amounts of blood sample and measuring time they require are much less than those conventional colorimetric method does. As technology has improved over the years, the latest generation of sensors only requires 0.3 μl (Freestyle by Therasense) or 1 μl (OneTouch Ultra by Lifescan), and measuring time has also been reduced to 5–10 seconds. Such sensors have become a guide for products of a similar kind and technological development, as well as for further research and development of different electrode structures. 
     U.S. Pat. No. 5,437,999 by Diebold et al in 1995 has disclosed a sensor including opposing working and counter electrode elements spatially displaced by a spacer having a cut-out portion forming a capillary space between the working and counter electrode elements and a vent port in the working or counter electrode where air can be vented. A precise minute amount of a sample can be introduced via the capillary space and brought into contact with electrodes and reagents. Such sensors can be fabricated by photolithography or screen printing but processes of affixing two insulating substrates with an electrode thereon are very complicated and expensive. U.S. Pat. No. 5,779,867 by Shieh in 1998 has also disclosed a glucose sensor generally comprising a sensor electrode, a reference electrode, and a corpuscle separation thin film carrier strip sandwiched therebetween, which can filter erythrocyte, and an opening where a sample can be introduced. The carrier strip can be used to control volume of the sample flowing into the carrier strip and to remove interruption of erythrocyte during reactions. However, the amount of the sample introduced and the speed of filtering cannot be effectively and efficiently controlled. U.S. Pat. No. 6,129,823 by Abbott has proposed an electrode strip in which electrodes are covered with one or more mesh layers. The improvement involves a partial occlusion of the mesh which underlays an aperture within an upper cover above the mesh, and the aperture is formed above or adjacent to a working electrode. The partial occlusion can reduce the total volume of blood required to perform a measurement. Such sensor only requires 2.0–2.5 μl of the sample but applies a mesh to reduce the volume of blood and distribute the sample. U.S. Pat. Nos. 6,299,757 and 6,338,790 by Therasense have also suggested two opposing working and counter electrodes with a highly hydrophilic thin film finely constructed therebetween, which can introduce a sample to a sample chamber. The volume of the sample can be strictly controlled down to 0.3 μl by the water hydrophilic thin film, which is the lowest in the field. However, the processes of fabricating such sensors are very complex and extremely costly. ROC (Taiwan) Patent Publication No. 268,095 by Shieh has disclosed the technique of electrode fabrication by screen printing, in which an electrically conductive film and insulating layer are produced by screen printing. A metal layer is formed by electroplating and a circular recess, containing a so-called bio layer, is formed by coating a working and a reference electrodes with insulating paste. Sample of about 10 μl can be dropped to the recess to be measured. Such technique requires a larger amount of sample and processing such sensors introduces numerous electroplating process steps. ROC (Taiwan) Patent No. 124,332 by Apex Biotechnology Corp. has disclosed an inflow area formed above an electrode area. Mesh containing surfactant is spread above the inflow and electrode areas and sample can be brought into the electrode area by capillary or siphon. Such application is similar to that developed by Abbott, which utilizes mesh for the inflow of sample and is thus more costly, is also restricted to the amount of sample required. 
     U.S. Pat. No. 6,258,229 by Winarta et al in 2001 has disclosed a disposable electrode strip, which claims to require less than 1 μl of liquid sample. A piece of gold/polyester or tin oxide/gold polyester film is cut to shape, forming a base layer of sensor. A CO 2  laser is used to score the gold or tin oxide/gold polyester film and the film is scored by the laser creating scoring line such that two electrodes at sample end and three contact points are formed at an electrical contact end. A piece of double-sided tape is cut to size and shape, forming middle layer with a U-shaped channel, which contains an electrode area. A top layer, which is placed and coextensive with the middle layer, has a vent opening, which forms a fluid sample channel between sample inlet and the middle of the vent opening, which enables the fluid channel to restrict the volume of fluid to less than 1 μl. Such design is similar to that disclosed in U.S. Pat. No. 5,120,420 by Nankai et al in 1992, except that electrodes are formed in a different way. The electrode sensor disclosed by Nankai et al is a bi-electrode sensor by screen printing an insulating board. A fluid channel is formed by transversely adhering two spacers on opposing ends of electrodes and a top layer without an opening on top of spacers, which in turn forms a channel transverse to a working electrode. By this way, the volume of sample flowing into the channel cannot be controlled and the sample is likely to float a vent opening, which causes contamination. Another improvement employed by Winarta et al, which applies a middle layer with a U-shaped opening on top of a working electrode and subsequently a top layer with a vent opening over the middle layer, forms a fluid sample channel between sample inlet and the middle of the vent opening. With this structure, sample fluid may float the vent opening when the size of which is too small. On the other hand, sample fluid will be retained at the edge of the vent opening when the size of which is appropriate. However, as the size of sensors is getting smaller, it is likely to touch the vent opening by hand which causes outflow of sample fluid and thus contamination. 
     From the above analysis, it is understood that in order to achieve smaller volume of sample fluid and faster analysis yet avoid any possible contamination, it is necessary to design electrodes which incorporate capillary and siphon. 
     SUMMARY OF THE INVENTION 
     It is an object of the present invention to provide a biosensor, which incorporates the above principle, and disclose an electrode area with rapid sample inflow and less volume with advantages such as simple structure, needing no mesh and no contamination due to outflow of sample fluid. According to the present invention, only 0.5–0.8 μl of sample is required and analysis can be completed in about 5–10 seconds. 
     According to the present invention, the biosensor is formed by screen printing and includes an electrode layer (electrode area) comprising two or three electrodes, which are a working electrode, a reference electrode and an auxiliary electrode (tri-electrode) on an insulating substrate. An active reaction layer containing reactant, reaction catalyst, mediator, wetting agent and surfactant is spread on the surface of the electrode layer. A sample inflow channel above the electrodes between an upper cover and a middle insulating layer is used to introduce sample solution into the electrode area and the active reaction layer by siphon or capillary. Ingredient of the sample can be analyzed by electrochemical potentiometric or amperometric method. Further, the present invention provides an upwardly extended close chamber formed within the upper cover above the electrode area adjacent to the front of conductive wires, which can be effectively used to control sample volume and “fill-and-position” the sample. 
    
    
     
       DESCRIPTIONS OF THE DRAWINGS 
         FIG. 1  is an exploded view illustrating the structure of an electrode sensor by screen printing according to the present invention with a slot; 
         FIG. 2  is an exploded view illustrating the structure of an electrode sensor by screen printing according to the present invention with a T-shaped slot; 
         FIG. 3  is a longitudinal, cross-sectional view of the electrode sensor by screen printing according to the present invention; 
         FIG. 4  is an exploded view illustrating the structure of an upper cover with an upwardly extended closed chamber formed therein according to the present invention; 
         FIG. 5  is a longitudinal, cross-sectional view of the structure of the upper cover with an upwardly extended closed chamber formed therein according to the present invention; 
         FIG. 6  is a longitudinal, cross-sectional view of the structure of an electrode sensor with an upwardly extended closed chamber formed therein; and 
         FIG. 7  shows the influence of whole blood volume on measurements. 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     Sensor 
     According to the present invention, the structure of a tri-electrode biosensor  10  by screen printing illustrated in  FIG. 1 . Conductive wires  12  made of electrically conductive gel such as silver and gold, are formed on an insulating base plate or substrate  11  which is made of polyvinylcholorde (PVC), polyester (PE), polyether, polycarbonate, or the like, by screen printing. Electrode strips are then formed on top of the conductive wires  12  by printing another layer of electrically conductive material s such as carbon, gold, and platinum. Electrodes containing a working electrode  13 , a reference electrode  14  and an auxiliary electrode  15  (no auxiliary electrode in a bi-electrode sensor) are disposed at one end above the layer of conductive wires  12 . The corresponding contact ports  13 ′,  14 ′ and  15 ′ at the other end with respect to the electrodes  13 ,  14 ,  15  can be connected to a measuring device (not shown) and a device activation line  16 ′ can be automatically recognized by the measuring device. A non-electrically conductive or an insulating middle layer  17  which acts as an insulating dielectric layer as well as provides spacing is disposed above the  17   a  formed therein, is disposed above the insulating substrate insulating base plate  11  containing electrodes  13 ,  14 ,  15  by adhesion or screen printing. The insulating middle layer  17  has a Slot  17   a  designates a sample inflow channel. An upwardly extended closed chamber  18   a  with volume of about 2 μl, is formed within an upper cover  18  above and in communication with slot  17   a  at the rear end of slot  17   a . An active reaction layer  20  containing substances of reactant, reaction catalyst (such as enzyme), mediator (such as dimethyl ferrocene, tetrathoiofulvalene), wetting agent (cellulose, hydroxyethyl cellulose, carboxymethyl cellulose, polyvinyl alcohol, polyvinyl, pyrrolidone and gelatine, etc.), and surfactant (Tween 20, triton X-100, surfynol, mega 8, etc.) is spread on the surface of the electrodes  13 ,  14 ,  15 , which defines an electrode reaction area where reactions take place. When the upper cover  18  is adhered to the middle layer  17 , the slot  17   a  defines a capillary inflow channel, which allows the sample such as blood to be rapidly introduced into and fill the electrode reaction area by capillary action upon contact with the front tip of the capillary inflow channel. Reactions induced by reaction catalyst can subsequently take place between reactant and mediator, in which electric current can be generated and measured by the measuring device. The inflow channel can provide the electrodes with rapid fill-in time (less than 1 second) and a minute amount of sample (less than 1 μl). 
     The structure of another electrochemical tri-electrode sensor  10  according to the present invention is illustrated in  FIG. 2 . Conductive wires  12  of electrically conductive materials such as silver, silver chloride, and gold, are formed on an insulating substrate  11 , by screen printing. Electrodes of electrically conductive materials such as carbon, carbon, and platinum, comprising a working electrode  13 , a reference electrode  14  and an auxiliary electrode  15  are printed on the conductive wires  12 . The corresponding electrodes  13 ′,  14 ′ and  15 ′ with respect to the electrodes  13 ,  14 ,  15  are contact ports to a measuring device (not shown in figure), whereas a device activation line U 16  can be automatically recognized by the measuring device. A middle layer  17  of insulating material with a T-shaped slot  17   a  formed there, is formed on top of the insulating substrate  11  containing electrodes by adhesion or coating a layer of insulating paste by screen printing. An upper cover  18  containing an upwardly extending closed chamber  18   a  with volume of about 2 μl is formed on top of the middle layer  17  and the closed chamber  18   a  is positioned above the intersection of the T-shaped slot  17   a.  A sample inflow channeled is formed between the layer  17  and the upper cover  18  while  17   b  and  17   c  form air vents on opposite sides of the sensor  10 . Sample such as blood can be rapidly introduced into and field an electrode reaction are  20  by capillary upon contact with the front tip of capillary inflow channel. Similar to  FIG. 1 , the sample is configured not to go beyond chamber  18   a  along the inflow channel. In addition, same venting effect can be achieved by removing either air vent  17   b  or  17   c.    
     Insulating Substrate 
     Insulating substrate can be made of a variety of materials such as polymer, plastics, and ceramics. Materials should be chosen according to the requirement and application of electrode materials. For example, soft material should be chosen for invasive type sensors to reduce pain and avoid hurting tissues. For such sensors, insulating polymer materials such as polycarbonate, polyester, polyethylene terephthalate (PET), polyvinylchloride (PVC), polyether, polyamide, polyurethane, polyamide, etc., can be adapted. On the other hand, rigid materials which are not easy to be ruptured or bent, such as ceramics including silica or aluminum dioxide, can be adapted. With regard to measurement outside a human body, width of the insulating substrate is generally between 3 and 15 μm and more precisely between 5 and 10 μm. Thickness is between about 50 and 800 μm and more precisely between 200 and 400 μm. Length of the insulating substrate depends on different factors and may be between about 1 and 8 cm and more precisely between 2 and 5 cm. 
     Layer of Electrically Conductive Wires and Electrodes 
     As illustrated in  FIG. 1 , a layer of electrically conductive wires  12  made of electrically conductive materials such as silver, gold and platinum, is formed by screen printing, which is for connecting electrodes and a measuring device. Materials with high electrical conductivity and low resistance can reduce impedance of the electrodes and therefore increase signals of detected current. Electrically conductive material such as carbon paste can be printed on top of the wires  12  and a device activation line  16  can be automatically recognized by the measuring device. Apart from a reference electrode  14 , wires  12  are completely coated. The exposed surface of silver wire in electrode  14  can be processed electrochemically to form a reference electrode of silver chloride, or processed electrochemically to form a reference electrode of silver chloride, or printed by silver/silver chloride ink. In the latter case, silver chloride processing is not necessary. 
     Middle Insulating Layer 
     Middle insulating layer  17  can be formed by printing or adhering dielectric material above electrodes, which in turn covers the carbon surface not required to be exposed and provides a reaction region with fixed area. 
     Reaction Reagents Area 
     Reaction reagents are spread on top of electrodes, which include reaction catalyst, buffer solution, binder, mediator, surfactant, etc. For example, when glucose is measured, the catalyst can be glucose oxidase or dehydrogenase. The ingredient of binder contains polymer or wetting agent including cellulose, polyvinyl alcohol, gelatine, surfactant, etc., such as Tween-20, Triton X-100, Surfynol, and Mega 8, which can dissolve and disperse sample and reagents and provide hydrophile and dispersion for capillary inflow channel. Therefore, the reaction reagent layer can provide both reaction and capillary, which not only fills sample in electrodes for analysis of reactions , but also provides electric current generated by reactions in electrodes for quantitative analysis of the sample. Preferred mediator, depending on requirement of different measurements, should have redox potential between −100 and +500 mV. Fore example, ferrocene such as dimethylferrocene, tetrathiafulvalene and derivative or complex of both can be applied. A lower potential can avoid interfering materials in the sample, while higher electron conducting efficiency can provide stronger electric current signals. Buffer solution can maintain pH within a fixed range, generally between 4 and 9 and preferably between 5 and 8. Useable buffer solutions include phosphoric salt, acetate salt, citrate salt, etc., and concentration can rage between 10 and 1000 mmole/1 and preferably between 30 and 1000 mmole/1. 
     Capillary Inflow Layer 
     Capillary inflow channel is formed by adding a middle layer  17  and an upper cover  18  on the top of electrodes  13 ,  14 ,  15 ,  17   a  represents a sample capillary channel and  17   b  and  17   c,  which can exist independently, are air vents on opposite site of a sensor shaped design). The volume of the inflow channel can be adjusted by varying thickness of the middle layer  17  and width of channel  17   a.  The thickness of the inflow channel is generally between 20 and 400 μm and preferably between 50 and 200 μm. The length of the hollow inflow channel is between 2 and 8 mm and the width of which is between 0.5 and 5 mm and preferably between 1 and 2 mm. The volume of the hollow inflow channel is between 0.05 and 16 μl and a volume between about 0.5 and 4 μl is required when actual measurement is performed. The time between a sample being in contact with the edge of the inflow channel and filling-in the inflow channel is less than 1 second. 
     The chamber  18   a,  in the upper cover  8  can be round, rectangular or of other geometry shape and the desired size can be between 0.5 and 4 mm. The location of an opening of the chamber  18   a  is above a rear end of the inflow channel and behind a working electrode. Blood sample can be filled in the reaction area, which the flowing of the sample is then stopped by the opening of the chamber. The spacing layer  17  and the upper cover  18  can be made of transparent opaque insulating materials such as plastics or polymers including PVC, Mylar, etc. Chamber  18   a  may be transparent for better inspection visually of sample flowing in and protection of sensor. The upper cover can be formed by 2 steps. The first step is to form opening  18   a  in the upper cover, as shown in  FIG. 1  and the second step is to apply another thin plate  19  (as shown in  FIGS. 4 and 5 ).  FIGS. 3 and 6  show the sensor illustrated in  FIG. 1  in longitudinal, cross-sectional view, which contains the thin plate  19 . 
     Filling Detecting Device 
     Filling detecting device is designed to detect if a sample is filled above three electrodes. For a tri-electrode type sensor, if working electrode is disposed at the outer edge of inflow channel, filling detection can be arranged by using working electrode and auxiliary electrode and by monitoring electric current, potential and impedance. Impedance between working and reference electrodes is infinite by potentionmetry when no sample is present and decreases significantly when sample is filled inside the inflow channel area, by which parameter of electrochemical analysis is activated when sample is filled. For a bi-electrode type sensor, similar method can be applied. In order to apply electrodes for filling detection, distribution of electrodes should be the same as direction of sample flow. That is, working electrode needs to be in contact with sample ahead of auxiliary electrode and subsequently compete filling of sample can be determined. Similarly auxiliary electrode can be arranged to be in contact with sample ahead of working electrode, and vice versa. 
     Electrochemical Analysis 
     When electrodes are assembled, sensors can be cut by die cutting or punching. Sample analysis can be performed by connecting the sensor to a palm electrochemical device. Analysis can be performed by varied methods, such as chronoamperometry (0–0.6 V), which measures stationary current, or total charge within fixed time at constant potential. The total amount of charge, which is integral of electric current and time, and stationary current are proportional to the concentration of sample. Measuring device can also incorporate filling detection in the sensor, where parameter of electrochemical analysis can be activated when the measuring device detects a signal of filling, which in turn can increase accuracy of measurement. Especially when the overall measuring time is less than 10 seconds, a tiny error in time may result in large difference. 
     The present invention will now be applied by way of taking blood sugar as examples. It is intended to demonstrate the preferred embodiments rather than to limit the scope of the present invention. 
     EXAMPLE 1 
     Fabrication of Glucose Sensor by Screen Printing 
     A layer of electrically conductive silver paste is formed on a polyporpylene synthetic substrate  11 . by 300 mesh screen printing, which is dried and heated for 30 minutes at 50° C., and three electrodes (working electrode  13 , reference electrode  14  and auxiliary electrode  15 ) are printed by carbon paste thereon. The substrate  11  is again heated for 15 minutes at 90° C. and printed by insulating gel, which is subsequently dried and hardened under ultraviolet light to form an insulating layer with an inflow reaction are  17   c ,  17   b  and  17   c  (for sensors with air vents). Reaction reagents of 2–6 μl, containing 0.5–3 units of glucose oxidase, 0.1–1% of polyvinyl alcohol, pH 4.0–9.0 and 10 14 mM potassium phosphate as buffer solution, 10–100 mN potassium chloride, 0.05–0.5% of dimethylferrocene, 0.005%–0.2% tween−20, 0.005%–0.2% of sufynol and 011%–1.0% of carboxymethyl cellulose are spread on the recessed inflow channel area  17   a.  The substrate is dried at 45° C. for one hour and an upper cover  18  with an opening  18   a  formed therein is adhered on top of the substrate  11 . A transparent upper cover  19  is pressed above the substrate  11  and sensors can be cut by die cutting from the substrate  11 . 
     EXAMPLE 2 
     Standard Glucose Solution and Whole Blood Measurement 
     Standard potassium phosphate buffer solution (pH 7.4) is disposed containing glucose with a concentration of 0–400 mg/dl. The sample solution is measured by an electrochemical device (CHInstrument Co. 650A) in conjunction with a sensor according to Example 1 under a measuring potential of 100 mV for 8 seconds. The volume of sample solution is 3 μl for every measurement and the volume of sample solution introduced into the sensor for every measurement is less than 3 μl. The measuring results are listed in Table 1: 
     
       
         
               
             
               
               
               
             
           
               
                 TABLE 1 
               
             
             
               
                   
               
               
                 Results of standard glucose measurements 
               
             
          
           
               
                   
                 Glucose Concentration (mg/dl) 
                 Charge (μ coulomb) 
               
               
                   
                   
               
               
                   
                  0 
                 0.690 
               
               
                   
                  25 
                 1.532 
               
               
                   
                  50 
                 2.952 
               
               
                   
                 100 
                 5.248 
               
               
                   
                 200 
                 7.400 
               
               
                   
                 400 
                 9.577 
               
               
                   
                   
               
             
          
         
       
     
     Whole blood sample can also be measured by sensors according to the present invention. Table 2 shows results of by measuring fresh vain whole blood sample with glucose additive with a measuring potential of 100 mV and volume of 2 μl. 
     
       
         
               
             
               
               
               
             
               
               
               
             
           
               
                 TABLE 2 
               
             
             
               
                   
               
               
                 Results of whole blood measurements with varied glucose addition 
               
             
          
           
               
                   
                 Glucose Concentration (mg/dl) 
                 Charge (μ coulomb) 
               
               
                   
                   
               
             
          
           
               
                   
                  80 
                 1.556 
               
               
                   
                 105 
                 2.636 
               
               
                   
                 130 
                 3.440 
               
               
                   
                 180 
                 5.946 
               
               
                   
                 280 
                 9.707 
               
               
                   
                 380 
                 11.733 
               
               
                   
                 480 
                 12.464 
               
               
                   
                 580 
                 13.945 
               
               
                   
                   
               
             
          
         
       
     
     EXAMPLE 3 
     Measurements of Blood Sugar with Varied Volume of Whole Blood 
     Electrode sensors according to Example 1 are employed, which provide whole blood samples with different volume required in the present invention. Vein whole blood samples are mixed with standard glucose solution, which in turn form solutions with a concentration of 300 mg/dl. 
     The method of measurements is to provide whole blood samples with different volume and supply samples by siphon under conditions set out in Example 2. As shown in  FIG. 7 , when the volume of a sample is insufficient (e.g., of less than 0.5 1), the concentration of glucose is low. Conversely, when the volume of a sample is above 0.8 1, the measured glucose concentration si near that in the sample solution, and the whole amount of the sample cannot be introduced into the sensor. That is, the more the volume of a sample is supplied, the more volume of the sample will be redundant, since inflow reaction channel is saturated with the sample and cannot accommodate more solution. The front edge of sample does not go beyond the intersection between  18   a  and the inflow channel, which is the evidence that the volume of sample solution can be effectively controlled and restricted.