Method for measurement of viscosity change in blood or blood plasma and sensor therefor

Abstract: Here is disclosed a method for measurement of viscosity change in blood or the like comprising steps of disposing a sensor including of an endothermic or exothermic element in blood or the like, stimulating blood or the like so as to cause a viscosity change therein and detecting the viscosity change by continuously measuring any one of changes occurring respectively in a average temperature ϑ w or a surface temperature ϑ s of the sensor containing therein the endothermic or exothermic element, a differential temperature ϑ w - ϑ ∞ or ϑ s - ϑ ∞ between a temperature ϑ ∞ of blood or the like and ϑ w or ϑ s, a kinematic viscosity ν of blood or the like and a heat transfer coefficient on the sensor surface. Such measurement may be performed by a sensor comprising an electric insulator through which lead wire extends, metallic wire wound around the electric insulator and connected at opposite ends to the portions of the lead wire exposed on the surface of the electric insulator, and a portion of the electric insulator, around which the metallic wire is wound, being coated with suitable material.

Description: BACKGROUND OF THE INVENTION The present invention relates to a method for measurement of viscosity changes, for example, a blood coagulation process, using an electrically conductive sensor as well as a sensor used to perform said method. Generally, determination of viscosity change occurring in blood or blood plasma gives knowledge of actual conditions of blood and, for example, enables a blood type to be easily identified. Furthermore, the determination of viscosity change has been commonly used to diagnose so-called hyperviscosity diseases such as a cerebral infarct and a myocardinal infarct. Especially, blood coagulation is particularly important and determination of the coagulation time has been most commonly used for diagnosis of various diseases such as haemophilia, von Willebrand's disease, Christmas disease and hepatic diseases. Pathologically, it has also been put in practice to know a condition of immune reaction by reacting plasma in blood with antigen or antibody. Typical methods for measurement of blood coagulation time which have conventionally been employed include those relying on measurement of prothrombin time (PT), measurement of activated partial thromboplastin time (APTT), measurement of thrombin time, fibrinogen test and hepaplastin test. Typical methods for examination of immune reaction include those relying on examination or measurement of complement supply reaction, fluorescent reaction and enzyme immunity. In the following, the term "blood" should also contain blood plasma. With the conventional methods for measurement of viscosity change in blood, determination of coagulation or other phenomena have been macrographically made in most cases, even through these methods have utilized stimulators and test reagents which are commercially available and contain therein predetermined stabilized ingredients. Subjective judgment of the operator has necessarily limited the reliability of measurements. Repetition of measuring procedures to improve such reliability of measurement has often resulted in inconsistent measurements. It is also well known to measure blood coagulation time by use of a mechanical apparatus, for example, through determination of prothrombin time by use of spectrophotometer. However, such a method is disadvantageous in that any disturbance on top of liquid to be tested will cause light scattering which, in turn, will cause a measurement error. Not only the spectrophotometer-method but also the other methods of prior art by which the blood coagulation time is measured by use of the mechanical apparatus are inconvenient and complicated in their mechanisms. The inventors disclosed a method for measurement of physical property change occurring in liquid or the like in Japanese Disclosure Gazette No. 1985-152943. The disclosed invention proposes, as a part thereof, the method closely related to the present invention, i.e., the method for detection of any thrombus formed on the inner wall of artificial blood vessel through a change of heat transfer coefficient as sensed by a sensor utilizing metallic wire which is fixedly arranged in the inner wall of said artificial blood vessel. However, such method intends only to detect formation of thrombus on the inner wall of the artificial blood vessel but not to detect any abnormal condition occurring in blood or plasma itself. SUMMARY OF THE INVENTION A primary object of the present invention is to provide a method for measurement of a viscosity change occurring in blood and a sensor usable to perform the method which can be commonly used for various kinds of measurement concerning the viscosity change in blood, without being prone to erroneous measurement. Know in the art are the methods disclosed by Soviet Documents OPISANIYE ISOBRETENIYA K AVTORSKOMU SVIDETELSTVU SU-A-590 665 and SU-A-1 157 456. According to these documents blood coagulation times are measured using thermo-resistor sensors. Due to the first document short voltage impulses are passed. During coagulation the resistance of the thermo-resistor is changing. In the second document, a method is revealed to determine the change of blood serum viscosity. A micro-thermo-resistor is used with high thermal coefficient of resistance in the region of coagulant phase. However, the data and features of said references are vague and generally spoken, so that they only can be taken as a hint that measurement of blood conditions are possible with resistor sensors. According to the present invention, various kinds of measurement concerning the viscosity change in blood may be achieved by a method according to claim 1. Such method for measurement of viscosity change in blood may be effectively performed by using a sensor for measurement of viscosity change on blood comprising a cylindrical electric insulator through which a pair of lead wire extends, a further metallic wire non-inductively wound around said electric insulator, said metallic wire being connected at opposite ends to the portions of said lead wire exposed on the surface of said electric insulators; a portion of said electric insulator around which the metallic wire is wound, is coated, with a suitable insulating material. Generally, a major portion of the time required for blood coagulation is a period elapsing before formation of activated thromboplastin in blood and the reaction slowly goes on. Thereafter, transformation from fibrinogen to fibrin rapidly occurs. Then, blood loses its fluidity and becomes coagulated. During such a process of blood coagulation, a viscosity change occurs in blood and simultaneously various values related to the sensor comprising the endothermic or exothermic element and disposed in blood also correspondingly change. Specifically, for example the average temperature  w of the sensor, the surface temperature  s of the sensor, the differential temperature  w - ∞ or  s -∞ between the temperature of fluid such as blood and  w or  s, respectively, and the heat transfer coefficient α on the surface of the sensor change in their respective numerical values. It is also well known that the changes in these numerical values are in functional relationship with the kinematic viscosity of blood, (see Introduction in "Journal of National Food Engineering Society", January, 1988). Thus, these changes may be determined intermittently or continuously to obtain a kinematic viscosity or an index value related to the kinematic viscosity and thereby to detect a viscosity change of blood. In immunity reactions the reaction will be measured as a blood viscosity change by the sensor in the same manner as has been mentioned above. The sensor may also have an antibody, such as IgG, fixed on the surface thereof so as to react with lactopherin contained in plasma, when a plurality of spheric plastic-latex particles are added to blood, so that an antigen/antibody reaction may occur on surfaces. These spheric particles agglutinate as a result of this reaction. BRIEF DESCRIPTION OF THE DRAWING These and other objects as well as advantages of the present invention will become apparent by the following description of preferred embodiments of the present invention with reference to the accompanying drawings, wherein: Fig. 1 is a schematic diagram illustrating, partially in section, how a sensor constructed in accordance with the present invention is used; Fig. 2 is a perspective view of the sensor constructed as one embodiment of the present invention; Fig. 3 is a detailed view corresponding to Fig. 2, parts being partially broken away; Fig. 4 is a perspective view of the sensor constructed as another embodiment of the present invention; Figs. 5A and 6A are graphic diagrams respectively plotting the prothrombin time of an abnormal plasma and of a normal plasma, in which the ordinate indicates a difference Δ  =  s - ∞ and the abscissa indicates, elapsing time; Figs. 5B and 6B are graphic diagrams respectively plotting the changing rates of Δ  with respect to the elapsing time given in Figs. 5A and 6A, in which the ordinate indicates ∂Δ ∂t and the abscissa indicates elapsing time; Figs. 7A and 8A are graphic diagrams respectively plotting the activated partial thromboplastin time of an abnormal plasma which is made from normal plasma through dilution by physicological saline, and of normal plasma, in which the ordinate indicates a difference Δ =  s - ∞ and the abscissa indicate elapsing time; Figs. 7B and 8B are graphic diagrams respectively plotting the changing rate of said Δ  with respect to the elapsing time given in Figs. 7A and 8A, in which the ordinate indicates ∂Δ ∂t and the abscissa indicates elapsing time; Figs. 9A and 10A are graphic diagrams respectively plotting the thrombin time of human normal plasma samples having fibrin content of 1% and 20%, in which the ordinate indicates a difference Δ  =  s - ∞ and the abscissa indicates elapsing time; Figs. 9B and 10B are graphic diagrams respectively plotting the changing rate of said Δ  with respect to the elapsing time given in Figs. 9A and 10A, in which the ordinate indicates ∂Δ ∂t and the abscissa indicates elapsing time; Fig. 11 is a graphic diagram roughly plotting the changes of  s and  w with respect to the prothrombin time (PT) of normal blood; and - Fig. 12 is a graphic diagram roughly plotting the change in the heat transfer coefficient with respect to the changes of  s and  w as illustrated in Fig. 11. DETAILED DESCRIPTION OF THE INVENTION First, the method according to the present invention will be discussed in reference with Fig. 1 and Figs. 5 through 10. It has already been described that the variable values  s,  w,  s -  ∞ ,  w -  ∞ , α and ν change as blood viscosity changes. Now, a relationship between the differential value  s -  ∞ ( s : sensor surface temperature;  ∞ : blood temperature) and the kinematic viscosity of blood will be considered by way of example. Concerning the relationship established between the steady state heat transfer coefficient α and the difference  s -  ∞ , the steady state heat transfer coefficient α is given by a following equation: α = Q A (s - ∞) where Q measured in Watts represents the heat generated in the sensor probe and transferred to the surrounding liquid or heat flux, and A represents the surface area of the sensor, measured in m². Accordingly, if both "Q" and "A" in equation (1) are known, the heat transfer coefficient α can be obtained from said differential temperature  s -  ∞. It should be understood that, when the sensor is cylindrical, the sensor surface area "A" can be calculated so far as the diameter "d" and the length "ℓ" of this cylindrical sensor are known, because A ≒ π˙d˙ℓ in such case. Then, the relationship between the differential temperature  s -  ∞ and the kinematic viscosity ν will be considered. The sensor is disposed, for example, in stationary distilled water of which the physical property values is already known, and then applied with constant current, (DC), which may have various values while a differential temperature  s -  ∞ between said distilled water and the (heated) sensor is measured. This procedure allows a relationship to be established among the Nusselt number Nu corresponding to the dimensionless number of the heat transfer coefficient, the Prandtl number Pr corresponding to the dimensionless number of the kinematic viscosity and the Grashof number Gr corresponding to the dimensionless number of the temperature difference, i.e., an equation generally representing free convection heat transfer around said sensor in the form, for example, of Nu = Co Gr C1 Pr C2 where Co, C1 and C2 represent constants. Nu, Gr and Pr can be expressed by following equations: Nu = α L/λ Gr = L 3 gβ ( s -  ∞)/ν 2 Pr = ν /a where "L" represents typical length (m), λ represents thermal conductivity (W/mK), "g" represents gravitational acceleration (m/s²), β represents coefficient of volumetric expansion (1/K), represents kinematic viscosity (m²/s), and "a" represents thermal diffusivity (m²/s). Accordingly, from the equations (2) through (5) as set forth above, the kinematic viscosity can be expressed by a following equation: ν 2C1-C2 = C  g C1 AL 3C1-1 Q -1 λ β C1 a -C2 ( s -  ∞) C1+1 When platinum wire adapted to be applied with current "I" and thereby to be heated is employed as the sensor, Q = RI 2 where "R" represents the electric resistance (Ω) of the platinum wire used as the sensor and "I" represents the value "A" of direct electric current applied to the sensor. In the above-mentioned equation (6), g, "A" and "L" represent constants. Further when the fluid include a large quantity of water or the composition of the fluid remain relatively unchanged, it can be assumed that λ, β and α respectively change in ranges sufficiently smaller that the renge in which changes, so that the kinematic viscosity ν can be ultimately expressed by a following equation as a function exclusively of the differential temprature  s -  ∞ and the heat flux "Q": ν 2C1-C2 = C 3 Q -1 ( s -  ∞) C1 + 1 where C₃ represents a constant. Using blood as the fluid (F) and applying the sensor (S) with current so as to maintain the heat flux "Q" constant, the differential temperature  s -  ∞ between blood and the surface of the sensor may be measured to obtain a kinematic viscosity ν and thereby to detect a change in blood viscosity. The above-mentioned blood temperature  ∞ may be measured by use of the resistance thermometer comprising platinum and said  s may be measured by employing the invention disclosed in the US Patent 4,611,928 of the inventors. According to said US-Patent 4,611,928 the relationship between the sensor surface temperature  s and the sensor average temperature  w is expressed by;  s =  w - Ao·I 2 (1 + α w  w) where α w : temperature coefficient of electric resistance I : value of current applied to the sensor Ao : constant And  w is expressed by  w = (V/i·Ro - 1)/α w where Ro : electric resistance developed in the sensor metallic wire at 0°C V : value of voltage across the sensor Therefore,  w can be calculated from the voltage value "V" and the current value "I" of the sensor, and  s can be calculated from  w. Furthermore, from equation (9) and (10) as set forth above,  s can be expressed by a function of both the voltage value and the current value,  s = f(V, I) Accordingly,  s can be calculated also from the voltage value and the current value. When the blood temperature  ∞ and the sensor average temperature  w is measured by use of the resistance thermometer,  s can be calculated also by using a following equation:  s =  ∞ + Co'( w -  ∞) C1' or  s =  ∞ + Co"(  w -  ∞) Co', Co", C1' : specific constants of the sensor In this way, the viscosity change occurring in blood or blood plasma can be determined by measuring the variable values of  s,  w,  s -  ∞ ,  w -  ∞ , α and ν . Now a method for determination of a viscosity change and, therefore, a specific coagulation time of blood from a change of  s -  ∞ , based on a result of the experiment conducted in connection with  s -  ∞, will be explained. For the experiment, an apparatus as shown by Fig.1 was used. Referring to Fig. 1, a sensor "S" is placed in a quantity of fluid "F" such as blood contained in a vessel 10 and, in a measuring system "M", a pair of current lead wire 2a, 2b connected to platinum wire are connected to a current source 5 for energization while another pair of voltage lead wires 2c, 2d are connected to a potentiometer 6 for voltage measurement so that electrical resistance of the sensor can be measured on the basis of four wire method. In the measuring system "M", the DC power source 5, the digital voltmeter 6 and a controller 7 are connected one to another by GP-IB (general purpose interface bus). With the platinum wire 3 being applied with current so as to maintain the heat flux thereof constant, by continuously measuring values of voltage impressed to the platinum wire 3, the sensor average temperature  w, the differential temperature  s -  ∞ between blood and the surface of the sensor, and the differential temperature  w -  ∞ between blood and the sensor is determined from the equations (9) through (13). Additionally there is the difference between the method of measuring the blood temperature  ∞ by one sensor and that by two sensors, which is described as follows; In case of one sensor, Step 1; the sensor is applied with direct current of about 100µ A or 1mA. Step 2; the measurement of voltage values V impressed to the platinum wire of the sensor. Step 3; the blood temperature  ∞ can be calculated from the equation (10) and the voltage value V. Step 4; further the sensor is applied with direct current of more than 1mA i.e., 20mA or 60mA. Step 5; the same as Step 2. Step 6; the sensor average temperature  w can be calculated by the same method of Step 3. Step 7; the differential temperature  s -  ∞ can be calculated from the equation (12). Repeating the measurement and the calculation of Step 5 through 7, the changing value of  s - ∞ with respect to the elapsing time can be obtained. In case of two sensor S₁, S₂, Step 1: the sensor S₁ is applied with direct current of about 100µ A or 1mA Step 2: the measurement of voltage values V₁ impressed to the platinum wire of the sensor S₁. Step 3: the blood temperature  ∞ can be calculated from the equation (10) and the voltage value V₁. Step 4: the sensor S₂ is applied with direct current of more than 1mA, i.e., 20mA or 60mA. Step 5: the measurement of voltage value V₂ impressed to the platinum wire of the sensor S₂. Step 6: the sensor average temperature  w can be calculated from the equation (10) and the voltage value V₂. Step 7: the differential temperature  s -  ∞ can be calculated from the equation (12). Repeating the measurement and the calculation of Step 2 through 3 and Step 5 through 7, the changing value of  s -  ∞ with respect to the elapsing time can be obtained. The sensor as has been described in reference with Figs. 2 through 4 was used in this embodiment. The result of the experiments conducted by the inventors will be set forth below. In the experiments of the inventors, samples of human normal plasma (VNC) were injection-added with various reagents and coagulant. Thereafter they were continuously measured by use of the sensor. EXAMPLE 1 Each 0.1ml of normal whole blood, normal plasma, abnormal whole blood and abnormal plasma was mixed by injection with 0.2ml of clotting factor stimulator such as tissue thromboplastin and calcium chloride. Then the prothrombin time (PT) was measured. Each quantity of liquid to be tested was prepared in a tube (diameter 8mm) provided with the sensor comprising platinum wire which presents electric resistance of 50Ω at a temperature of 0°C, and the sensor was heated by applying thereto direct electric current of about 40mA or about 60mA. Figs. 5A and 5B show the differential temperature  s - ∞ measured as a function of the elapsing time using abnormal plasma as liquid to be tested, with the sensor heating current of about 40mA, and Figs. 6A and 6B are similar to Figs. 5A and 5B except that normal plasma was used as liquid to be tested with the sensor heating current of about 40mA. In the experiment corresponding to Figs. 5A, 5B and Figs. 6A, 6B, a point of inflection, at which the changing rate of the differential temperature  s -  ∞ sharply decreases was determined as a fixed point and thereby the prothrombin time (PT) was measured to be 34.6 sec and 13.0 sec, respectively. 10 % 100% Pt, VNC 34.6 sec 13.0 sec I = 40mA (Fig. 5) (Fig. 6) In the Table 1, 10% means that a quantity of normal plasma was diluted 10 times by physiological saline so as to prepare a sample of plasma like as collected from a patient at imminent exitus (i.e., almost dead) ; 100% means, that the sample comprises only normal plasma. EXAMPLE 2 Each 0.2ml of normal whole blood, normal plasma, abnormal whole blood and abnormal plasma was mixed by injection with 0.1ml of clotting factor stimulator comprising kaolin, phospholipid, cerite, silicic acid, elaidic acid or the like. Then the activated partial thromboplastin time (APTT) was measured. Each quantity of liquid to be tested was prepared in a tube (diameter 8mm) provided with metallic wire serving as the sensor and the sensor was heated by applying thereto current of about 40mA. Figs. 7A and 7B show the differential temperature  s -  ∞ measured as a function of the elapsing time using abnormal plasma as liquid to be tested, with the sensor heating current of about 40mA, and Figs. 8A and 8B are similar to Figs. 7A and 7B except that normal plasma was used as liquid to be tested with the sensor heating current of about 40mA. In the experiments respectively corresponding to Figs. 7A, 7B and Figs. 8A, 8B, the activated partial thromboplastin times (APTT) were determined to be 176.0 sec and 32.1 sec, respectively, in the same manner as in EXAMPLE 1. 10 % 100 % APPT, VNC i = 40mA 176.0 sec (Fig. 7) 32.1 sec (Fig. 8) The meaning of 10% and 100% is same as in Table 1. EXAMPLE 3 A series of plasma samples was prepared by adjusting their fibrin contents stepwise from 1% to 20% of that normally found in healthy human plasma through dilution thereof by physiological saline, and the thrombin time was determined on these sample after mixed by injection 0.2ml of clotting factor stimulator to 0.1ml of the respective samples and applying the sensor with current of about 60mA. Figs. 9A and 9B show the differential temperature  s -  ∞ as a function of the elapsing time in the sample of which the fibrin content is 1% of that normally found in healthy human plasma, and Figs. 10A and 10B show the result obtained on the sample of which the fibrin content is 20% of that normally found in healthy human plasma. In these experiments of which the results are shown by Figs. 9A, 9B and Figs. 10A, 10B, the thrombin times were also determined in the same manner as in Example 1 to be 151.3sec and 3.4sec, respectively. Similarly, the thrombin times were determined, on the samples of which the respective fibrin contents are 2% and 10% of that normally found in healthy human plasma, to be 54.3sec and 7.7sec, respectively. As apparent from three Examples as have been described above, a viscosity change occurring in blood or blood plasma can be detected by determining a change in the value  s -  ∞ . Determination of a change in the kinematic viscosity ν permits a change in the viscosity to be detected, by finding a point of inflection in the same manner as in the case of  s -  ∞ , since the kinematic viscosity ν can be expressed as a function only of the heat flux "Q" and  s -  ∞ by the equation (8). The heat transfer coefficient α and  s -  ∞ are reciprocal to each other, since these are in the mutual relationship as expressed by the equation (1). Accordingly, a relationship between a variation of the heat transfer coefficient α and the time "t" can be plotted as a graphic diagram which is reciprocal to the above-mentioned Figs. 5A, 6A, 7A, 8A, 9A and 10A, and a point of inflection may be determined in the same manner as in the case of  s -  ∞ to detect the viscosity change occurring in blood or the like, as seen in Fig. 12. For example, concerning  w -  ∞ , transferring  ∞ from the right side to the left side in the equation (12) results in a following equation:  s -  ∞ = Co'( w - ∞) C1' It will be appreciated from this equation (14) that  w -  ∞ is a function of  s -  ∞ and, just as in the case of  s -  ∞ , a point of inflection may be determined along the curve of  w -  ∞ to detect a viscosity change occurring in blood or the like.  s or  ∞ changes generally as shown by Fig. 11 as the viscosity changes, when the clotting factor stimulator is added by injection to the sample of healthy human blood and then the prothrombin time (PT) is determined. Thus, it is also possible to detect the viscosity change of blood or blood plasma by finding a point of inflection along a curve plotted by the value of  s or  ∞ varying as the viscosity changes. It should be understood that Fig. 12 shows how the heat transfer coefficient α changes with respect to variation of  s and  w shown in Fig. 11 and the coagulation time (Tc) can be determined by finding a point of inflection along a curve plotted by the heat transfer coefficient α . In this manner, the viscosity change can be easily determined also from respective variations of  s,  w,  w -  ∞, α and ν . As will be apparent from the foregoing description, the method of the present invention enables the viscosity change being important to know actual condition of blood or blood plasma to be easily and exactly detected without interposition of any human subjective judgment. With a consequence, the other conditions of blood such as the blood type can be also easily identified by the method of the invention, because the manner of coagulation depends on a particular blood type. Additionally, as has been discribed in reference with Fig. 7 and 9, even a patient being in such serious condition that the blood coagulation time tends to be prolonged can be diagnosed almost independently of any influence by the inevitable low viscosity. The method of the invention further enables even a small variation occurring in blood to be easily detected, thus minimizing an error possibly occurring in diagnostic measurement and allowing the blood diagnosis to be performed over a wide range. The method of the invention effectively simplifies the procedure of measurement and, therefore, measurement may be repeated to improve the accuracy. Moreover, the method of the invention simplifies the construction of the equipment, particularly, the detection sensor used for measurement and correspondingly reduces the cost with respect to the prior art. While the viscosity change was caused by stimulating with addition of clotting factor stimulator in the previously mentioned Examples 1 through 3 of the method according to the present invention, the method of the invention is not limited to such manner of stimulation. Antigen or antibody may be brought into contact with or added into blood or blood plasma to stimulate the latter. For example, IgG may be used as the antibody and fixed onto the sensor surface, so as to react with lactopherin so that quantity of lactopherin can be measured. Stimulation may be also achieved by suitable physical means, for example, by use of high frequency vibration acting on blood or blood plasma or by heating. Now the sensor of the present invention will be described in reference with Figs. 2 through 4 which illustrate one embodiment of the sensor constructed according to the present invention. Referring to Figs. 2 and 3, 1 designates a cylindrical electric insulator, 2a and 2b or 2c and 2d designates a pair of lead wire extending through said electric insulator 1, and 3 designates platinum wire noninductively wound around the electric insulator 1 so as to form a measuring section "S'". The pair of lead wire 2a and 2b or 2c and 2d extend from a rear end 1" through the electric insulator 1 slightly beyond a front end 1' of said electric insulator 1 and then U-turn again to extend through said electric insulator 1 beyond said rear end 1". Opposite ends of the platinum wire 3 are electrically connected to two pairs of lead wire at their respective U-turn sections by means of spot welding or the like. The measuring section "S'" around which the platinum wire 3 is wound is coated with glass layer 4. Fig. 4 illustrates a variant of the sensor constructed in accordance with the present invention. In this embodiment, the platinum wire 3 is noninductively wound around the electric insulator 1 which is, in turn, coated with synthetic resin layer 11. The sensor of the present invention is made in a manner as will be described below. A ceramic hollow rod having a length of 50mm and a diameter of 1.4mm is used as the electric insulator 1 and the platinum wire having a diameter of 13 µm is noninductively wound around said ceramic hollow rod so as to define at its front end the measuring section "S'" extending axially along the ceramic hollow rod over a length of 3mm. Glass pipe is fitted around this assembly and heated to be deposited thereon or said assembly is immersed in liquid glass. Alternatively, said assembly may be immersed in a resinous monomer dispersed system, then the platinum wire may be heated by energization to promote thermal polymerization around the outer surface of the platinum wire and thereby to form a resinous polymer layer around the platinum wire. Here is provided thereby the sensor being capable of analyzing even a small quantity of sample with a high sensibility. The sensor of this invention may be used in accordance with the procedure as has been described in reference with Fig. 1 to detect the viscosity change in blood or blood plasma. Specifically, blood is used as the fluid "F", the sensor "S" is applied with current so as to maintain the heat flux "Q" constant and thereby a variation of the differential temperature  s -  ∞ between blood and the surface of sensor can be determined. Moreover, on the grounds that the sensor average temperature  w is unchangeable, it can be assumed that the heat flux "Q" is constant so as to keep the direct electric current "I" constant. Based on this variation of  s - ∞ , it is possible to determine the viscosity change occurring in blood or blood plasma. Furthermore, the variable values such as  w,  s,  w -  ∞ , α and ν can be calculated from the previously mentioned equations and these numerical values also enable the viscosity change in blood or blood plasma to be determined. Experimental use for detection of the viscosity change in blood or blood plasma indicated that the sensor of this invention actually provides the effect just as has previously been described in connection with the method of the invention for detection of the blood viscosity change. Finally, the sensor of this invention is advantageous in that even a small quantity of sample can be analyzed with a high sensibility and that adhesion of blood ingredients onto the sensor can be sufficiently reduced to facilitate cleaning thereof because the sensor surface is coated with glass or synthetic resin.

Claims: A method of measurement of viscosity changes in blood or blood plasma, using an electrically conductive sensor (S) which is electrically heatable by passing an electric current therethrough and having means for determining the voltage impressed on said sensor by said electric current, comprising the following steps: (1) immersing said sensor in said blood (F) or blood plasma; (2) applying an electric current to said sensor (S); (3) measuring the current applied to the sensor and the voltage (V) impressed on said sensor by said electric current, and calculating therefrom a sensor temperature of at least one of the average sensor temperature and the surface temperature of the sensor; (4) measuring the temperature Θ ∞ of said blood; (5) calculating the difference between the sensor temperature and said blood temperature Θ w - Θ ∞ and (6) detecting changes in the said difference, whereby the changes in the viscosity of the said blood are correspondingly detected. A method for measuring significant changes of viscosity ν according to claim 1 characterized in that said sensor (S) contains a heat source heatable by an electric current I o flowing through an electrical resistor forming a part of said sensor, thereby maintaining an average sensor temperature Θ w and a surface temperature Θ s of the sensor, respectively, expressed by the following equation: Θ s = Θ w - A o · I 2 (1 + α w · Θ w ) where α w is the temperature coefficient of the electric resistance and A o a constant, and Θ w is expressed by the following equation: where R o is the electrical resistance developped in the sensor, V is the voltage across the sensor, so, as Θ s may be expressed as a function of V and I o , and ν as a function f of Δ Θ = (Θ s - Θ ∞ ), a change of Δ Θ measured over the time t is also to be defined as a change of the viscosity ν of the sample of blood or blood plasma, thereby detecting changes in the viscosity by detecting changes in Δ Θ at an elapsed time t. A method for measuring a change of ν according to claim 2, characterized in that for measuring the prothrombin time (PT) or the activated partial thromboplastin time (APTT) a clotting factor stimulator is added to the samples and the elapsed time t till the change rate of the differential temperature Δ Θ is determined as a fixed point and thereby PT or APTT is measured. Use of the method for measuring significant change of ν according to claim 2 for measuring immunity reactions in samples of blood or blood plasma. Use of the method for measuring significant change of ν according to claim 2 to identify the type of blood. A sensor for measuring significant viscosity changes according to claim 2 characterized in a cylindrical electric insulator (1) through which a pair of lead wires (2a, 2b; 2c, 2d) extends, a further metallic wire (3) non inductively is wound around said electric insulator and being connected at opposite ends to the portions of said lead wire exposed on the surface of said electric insulator (1), and in that the portion of said electric insulator (1) around which the metallic wire (3) is wound, is coated with a suitable insulating material (11). A sensor according to claim 6, characterized in that the portion of the electric insulator around which the metallic wire is wound is coated with glass. A sensor according to claim 6, characterized in that the portion of the electric insulator around which the metallic wire is wound is coated with synthetic resin.

Inventors: HORI TOMOSHIGE, ITOH KENSUKE, SHIINOKI YAHUHIKO KITASHINAGAW, HORI, TOMOSHIGE, ITOH, KENSUKE, SHIINOKI, YAHUHIKO KITASHINAGAWA APARTMENT 203
Applicants: SNOW BRAND MILK PROD CO LTD, SNOW BRAND MILK PRODUCTS CO., LTD.
Application year: 1989
Main classification: G01N33
IPC codes: G01N33, G01N27