JP4751089B2 - Blood rheology measuring device - Google Patents

Description

  The present invention relates to an apparatus for measuring body fluid circulating in a living body, and more particularly to a technique for grasping the state of blood and evaluating health, diagnosing diseases, evaluating drug effects, and the like.

  As one of the examination items for judging the health condition of the human body, blood rheology measurement focusing on blood fluidity has attracted attention. As a means for measuring blood rheology, an apparatus (product name MC-FAN) for measuring the time required for a certain amount of blood collected from a subject to pass through a microchannel (microchannel) has been developed (see Non-Patent Document 1). .) At present, the MC-FAN apparatus is a standard machine in blood rheology measurement.

However, in the measurement using the MC-FAN apparatus, blood must be collected as described above, and the measurement can be performed only by medical institutions, which is extremely inconvenient for the purpose of easily checking the health condition of anyone at any time. large. In addition, blood collection has a large physical and psychological burden on the subject, and since the number of times that can be measured per day is only a few times at most, continuous data cannot be obtained in time series. There is a strong correlation between blood rheology and blood flow velocity in vivo. That is, it is considered that when the viscosity of blood is high, the blood flow velocity is slow, and when the viscosity is low, the blood flow velocity is high. Therefore, it is possible to know blood rheology indirectly by measuring the blood flow velocity in the living body.
In order to calculate the blood rheology index from the blood flow velocity in the blood vessel, it is necessary to measure the blood pressure of the living body using a cuff in addition to the blood flow velocity measurement as described in Patent Document 1. there were. The blood rheology using this blood pressure value and blood flow velocity, that is, the method for calculating the index of blood kinematic viscosity, is based on the concept that the blood flow pressure inside the target artery is approximated by the blood pressure value. (See Patent Document 1).
JP 2003-159250 A Keiji Kikuchi “Measurement of whole blood fluidity using a capillary model” (Food Research Result Information, NO.11, 1999)

  However, the method of calculating blood rheology, that is, the index of blood kinematic viscosity using the blood pressure value and blood flow velocity, approximates the blood pressure inside the artery with the blood pressure value, resulting in measurement errors. There was a problem of being big. Furthermore, from the viewpoint of blood pressure measurement mechanism and complexity, there is also a problem that it is difficult to downsize an apparatus indispensable for measuring blood rheology at a site such as a wrist or a fingertip.

  Accordingly, an object of the present invention is to provide a simple, high-precision and small-sized blood rheology measuring apparatus and measuring method that can be measured at a site such as a wrist or a fingertip and does not require blood pressure measurement.

In order to solve the above-mentioned problems, a blood rheology measurement device according to the present invention includes a means for detecting a pulse wave phase velocity from arterial blood flow information periodically pulsating in synchronization with a heartbeat of a living body, and the pulse wave phase. It is characterized by comprising computing means for computing blood rheological indices from speed.

FIG. 12 is a characteristic diagram showing the effect of the present invention, and is an index of blood rheology by a blood collection method using the blood rheology index value μ and MC-FAN measured by the blood rheology apparatus according to the present invention. The correlation of blood transit time T is shown. The vertical axis indicates the μ value, and the μ value is smaller near the origin of the vertical axis as seen in FIG. 12, and the β value increases as it goes upward. Although details will be described later, a small μ value means that the viscosity of blood is large.
On the other hand, the horizontal axis indicates the whole blood passage time T. The value of T is smaller near the origin of the vertical axis as viewed in FIG. 8, and the value T increases as it goes upward. That is, a small value of the whole blood passage time T means that the blood being measured is a smooth blood, that is, the μ value is a large value. On the other hand, a large value of the whole blood passage time T means that the blood being measured is muddy blood with a high viscosity. That is, a high viscosity means that the μ value is small, and when these relationships are taken into consideration and FIG. 12 is viewed, the μ value and the whole blood passage time T have a significant correlation. Can be considered.
Accordingly, as can be seen from FIG. 12, the blood rheology measuring apparatus according to the present invention can accurately measure blood rheology with the wrist or fingertip without the need for blood pressure measurement. Blood rheology measurement device can be supplied, and as a result, it is possible for anyone other than an expert to easily check the accurate rheology without collecting blood from the subject, and to use it for health status confirmation. .

(Embodiment 1)
FIG. 1 is a block diagram for explaining an embodiment according to the present invention. FIG. 1 is a block diagram for detecting a pulse wave phase velocity from an ultrasonic Doppler signal based on a blood flow velocity.
The sensor unit is composed of two sensor units. Both of these two sensor units are ultrasonic sensors, and are composed of an ultrasonic sensor 1 and an ultrasonic sensor 4. These two ultrasonic sensors have the function of transmitting ultrasonic waves and receiving reflected ultrasonic waves with Doppler signals generated at the two parts at two different parts separated by a specific distance of the same artery. I have.
The ultrasonic sensor 1 includes a transmission element 2 and a reception element 3, and the ultrasonic sensor 4 includes a transmission element 5 and a reception element 6, and the transmission element 2 and the transmission element 5 are connected to an ultrasonic oscillation circuit 7. The transmitting element 2 and the transmitting element 5 convert electrical vibration generated in the ultrasonic oscillation circuit 7 into mechanical ultrasonic waves and transmit ultrasonic waves into the living body.
The ultrasonic wave transmitted from the transmitting element 2 into the living body is reflected along with the Doppler signal to the blood flow in the first part of the target artery. The reflected ultrasonic wave accompanied with the Doppler signal is converted into an electric signal by the receiving element 3 and then input to the first reception detection circuit 8. The first reception detection circuit 8 detects the Doppler signal in the first part. Moreover, the ultrasonic wave transmitted from the transmitting element 5 into the body is reflected with a Doppler signal to the blood flow in the second part of the target artery. The reflected ultrasonic wave accompanied with the Doppler signal is converted into an electric signal by the receiving element 6 and then input to the second receiving detection circuit 9. The second reception detection circuit 9 detects the Doppler signal in the second part.

The Doppler signal of the first part and the Doppler signal of the second part detected by the first reception detection circuit 8 and the second reception detection circuit 9 have a blood flow velocity with a periodic time change synchronized with the biological heartbeat. Contains signal components. A device for separating and extracting a blood flow velocity from these Doppler signals is a blood flow velocity calculation processing device. That is, the arithmetic processing device for extracting the blood flow velocity signal component of the first part from the Doppler signal of the first part is the first blood flow velocity arithmetic processing apparatus 10, and the second processing unit calculates the second blood flow velocity calculation processing apparatus 10 from the second part. The second blood flow velocity calculation processing device 11 is an arithmetic processing device for extracting the blood flow velocity signal component of the part. Two types of blood flow velocity signals output from these two types of blood flow velocity calculation processing devices are input to the rheology calculation processing unit 12 according to the present invention. The rheology calculation processing unit 12 includes a pulse wave phase velocity calculation processing unit 13 and a rheology index calculation processing unit 14. The rheology index detected by the rheology calculation processing device 14 is output by the output device 15.
(Embodiment 2)
FIG. 2 is a block diagram for explaining a second embodiment according to the present invention. FIG. 2 is a block diagram for detecting the pulse wave phase velocity from the volume pulse wave. The sensor unit is composed of two sensor units. These two sensor parts are both optical sensors, and are composed of an optical sensor 16 and an optical sensor 19. These two optical sensors have a function of irradiating light to two different parts separated by a specific distance of the same artery and receiving attenuated light caused by volume pulse wave changes at the two parts.
The optical sensor 16 is a sensor for measuring the volume pulse wave at the first site of the target artery, and is composed of a light emitting element 17 and a light receiving element 18. The optical sensor 19 is a sensor for measuring a volume pulse wave in the second part of the artery, and is composed of an issuing element 20 and a light receiving element 21. The light emitting element 17 and the light emitting element 20 are connected to the light emitting circuit 22. The light emitting element 17 and the light emitting element 20 irradiate the living body with light. This irradiated light quantity attenuates in proportion to the volume pulse wave in the artery.
The light emitted from the light emitting element 17 is received by the light receiving element 18 as attenuated light proportional to the volume pulse wave at the first site in the artery and converted into an electrical signal. This electrical signal is input to the first light receiving detection circuit 23, and the volume pulse wave signal of the first part is output from the first light receiving detection circuit 23. Similarly, the light emitted from the light emitting element 20 is received by the light receiving element 21 as attenuated light proportional to the volume pulse wave in the second part of the artery and converted into an electrical signal. This electrical signal is input to the second light receiving detection circuit 24, and the volume pulse wave signal of the second part is output from the second light receiving detection circuit 24.
These two types of plethysmogram signals are input to the rheology calculation processing unit 12 according to the present invention, as in FIG. The rheology calculation processing unit 12 includes a pulse wave phase velocity calculation processing unit 13 and a rheology index calculation processing unit 14. The rheology index detected by the rheology calculation processing unit 14 is output by the output device 15.
(Embodiment 3)
FIG. 3 is a block diagram for explaining a third embodiment according to the present invention. FIG. 3 is a block diagram for detecting the pulse wave phase velocity from the movement displacement of the living body surface propagated from the artery that moves in synchronism with the arterial blood flow. The sensor unit is composed of two sensor units. These two sensor parts are both displacement detection sensors, and are composed of a displacement detection element 25 and a displacement detection element 26. These two displacement detection elements have a function of detecting a biological surface displacement propagating from a beating artery at two different biological surface sites separated by a specific distance.
The displacement detection element 25 is an element for measuring the displacement at the first part on the surface of the living body, and the displacement detection element 26 is an element for measuring the displacement at the second part on the surface of the living body. These two displacement detection elements are connected to a first drive detection circuit 27 and a second drive detection circuit 28, respectively. The displacement signal at the first part on the surface of the living body is output from the first drive detection circuit 27, and the displacement signal at the second part is output from the second drive detection circuit 28, and the rheology calculation processing according to the present invention is the same as in FIGS. Input to the unit 12. The rheology calculation processing unit 12 includes a pulse wave phase velocity calculation processing unit 13 and a rheology index calculation processing unit 14. The rheology index detected by the rheology calculation processing device 14 is output by the output device 15.
The theoretical background of the arithmetic processing performed by the rheology index arithmetic processing device 12 according to the present invention will be described below. FIG. 4 is a schematic diagram of an artery that pulsates in synchronism with a heartbeat, and blood flow has a blood flow velocity distribution 31 in the Z-axis direction and the radial direction, which is the axial direction of the artery, due to the pressure distribution 30 in the artery 29. Arise. Needless to say, the pressure distribution 30 correlates with the blood pressure value. Further, since the arterial wall 32 has elasticity, the arterial wall 32 undergoes vibrational displacement in the Z-axis direction and the radial direction. This vibration displacement is a pulse wave displacement 33 in the figure. Further, the pulse wave displacement 33 propagates through the artery wall 32 as a wave in the Z-axis direction along with the pulsation. This wave is a pulse wave 34. Further, the pulse wave displacement 33 propagates through the living body internal tissue 35 and induces a living body surface displacement 37 on the living body surface 36. This biological surface displacement 37 propagates as a wave on the biological surface in parallel with the pulse wave 34 of the artery. This wave is a biological surface wave 38.
In the above model, the pressure distribution, the blood flow velocity distribution, and the pulse wave displacement can be analytically determined using the Navier-Stokes equation in fluid dynamics and the dynamic equation of motion of the arterial wall. That is, assuming that the angular frequency of the heartbeat is ω and the phase velocity of the pulse wave 34 is C, the pressure distribution 30 of the artery at any time t, any Z-axis position coordinate z and any radial position coordinate r is P, The axial velocity component of the blood flow velocity distribution 31 is V, the arterial radial displacement ξ and the Z axial displacement η of the arterial pulse wave 34 are blood density, blood kinematic viscosity, arterial inner diameter, arterial wall thickness, arterial young. The rate and arterial Poisson's ratio are determined by the following equations (1) to (4), with ρ, ν, R, h, E, and σ, respectively.

Further, the change amount S of the volume pulse wave can be defined as the change amount of the cross-sectional area of the artery 29 shown in FIG. 4, but since the pulsatile displacement amount ξ is usually sufficiently smaller than the artery radius R, the following ( 5) Can be approximated by equation

  Here, Po and Pm in the equation (1) are a constant pressure term and an amplitude term of the arterial pressure distribution that periodically change in synchronization with the heartbeat. Jo in the equations (1) and (2) is a 0th order Bessel function. In the equations (2) to (5), the dimensionless functions F and Φ are expressed by the following equations using dimensionless parameters γ defined by the artery wall density ρo, blood density ρ, artery inner diameter R, and artery wall thickness h. Given in.

Where the dimensionless parameter γ is

In the equation (7), Γ is a gamma function. Further, the variable α defining the equations (6) and (7) is defined by the following equation using the kinematic viscosity ν of blood, the heartbeat angular frequency, and the artery inner diameter R.

From the above analysis results, the blood flow velocity V, the volume pulse wave S, and the pulse wave displacements ξ and η are all proportional to the amplitude Pm of the arterial pressure and at the same time determined by the heartbeat angular frequency and the pulse wave phase velocity C. Wave component, ie

It turns out that it has. Similarly, in the biological surface wave 38 in FIG. 4, it is difficult to give a detailed theoretical expression related to the amplitude component, but it is proportional to the amplitude Pm of the arterial pressure and the traveling wave component Ψ (t, z).

  The pulse wave phase velocity C is determined from a mechanical boundary condition in which the pulsating viscous flow and the vibration of the artery are satisfied in the artery wall. That is, the pulse wave phase velocity C is

Determined by Here, the function Φ is a dimensionless function defined in (6) and (7). This pulse wave phase velocity C has a characteristic that is determined independently of the amplitude term Pm of the intra-arterial pressure. Further, since the dimensionless function Φ is a function of the parameter α including the kinematic viscosity ν of blood defined by the equation (9), the pulse wave phase velocity C can be regarded as a function of the kinematic viscosity ν of blood. Prove. In the equation (11), the pulse wave phase velocity Co when the kinematic viscosity ν of blood is 0 is

Therefore, the normalized phase velocity Cm is defined by the following equation.

FIG. 5 is a characteristic diagram showing the relationship between the normalized phase velocity defined by equation (13) and α defined by equation (9). The vertical axis represents the normalized phase velocity Cm, and the horizontal axis represents α. It is found that the normalized phase velocity Cm has a monovalent function relationship with α, and the function shape is a characteristic curve 39. This characteristic curve 39 shows typical values of σ and γ in living tissue.
σ = 0.5 γ = 0.1-0.5
It turned out that there was no big change, and it was decided by the change of the value of α. In the measurement range in which the intra-arterial radius R can be regarded as substantially constant, this α is independent of the intra-arterial pressure P, as defined by the equation (9), and the heartbeat angular frequency ω, blood kinematic viscosity ν and It is determined by the radius R within the artery. That is, if the value of α is calculated using the characteristic curve 39 shown in FIG. 5 and the pulse wave phase velocity C or the normalized phase velocity Cm, the kinematic viscosity ν of blood can be detected without measuring blood pressure. . That is, the fundamental part of the blood rheology measurement method according to the present invention is to measure the pulse wave phase velocity, which is a physical quantity unrelated to the intra-arterial pressure, from the traveling wave component Ψ (t, z) given by the equation (10). There is a fundamental part of the calculation processing to calculate α by calculating the pulse wave phase velocity. Incidentally, the value of Co given by equation (12) varies depending on race, gender, age, and arterial site to be measured, but it is assumed in advance for a subject who uses the blood rheology device according to the present invention. There is no problem if a method of setting a plurality of numerical values in the arithmetic processing unit and setting them according to the target person is adopted. The above is the theoretical background of the arithmetic processing according to the present invention.
As described above, if the value of α is determined using the characteristic curve 39 shown in FIG. 5, the kinematic viscosity ν of blood can be determined. That is, the pulse wave phase velocity C or the normalized pulse wave phase velocity Cm is nothing but an index of blood rheology.
Hereinafter, a calculation processing method relating to a specific blood rheology index according to the present invention will be described.

  The normalized phase velocity Cm defined by the equation (13) is converted as follows.

FIG. 6 is a characteristic diagram showing the relationship between β and α defined by equation (14), where the vertical axis is α and the horizontal axis is β. By transforming the normalized phase velocity Cm into a new variable β using the equation (14), α, which is an indicator of blood rheology, can be approximately approximated by a cubic curve of β as in the characteristic curve 40 in the figure. Things turn out. That is, when the normalized phase velocity Cm is calculated from the measured pulse wave phase velocity C and the β value is calculated using the equation (14), the α value is determined as a polynomial of β shown below. is there.

Here, the coefficient an of the polynomial and the degree m of the polynomial in the equation (15) are numerical values obtained experimentally. Furthermore, these numerical values are numerical values that vary depending on the race, sex, age, medical history, etc. of the subject who uses the blood rheology apparatus according to the present invention. However, if a method in which a plurality of numerical values are built in the arithmetic processing unit and set according to the target person is adopted, it is merely a design matter without any problem.
The blood obtained by determining the α value using the β value given by the equation (14) and the equation (15), and dividing by the square root of the heartbeat angular frequency ω, thereby eliminating the influence of the α change due to the heart rate change. The rheological index μ is introduced. That is,

Thus, it is possible to detect blood rheology with higher accuracy, that is, kinematic viscosity ν of blood. The above is the method for determining the blood rheology index according to the present invention.

Next, the structure of the sensor unit of the blood rheology apparatus according to the present invention, the method for determining the pulse wave phase velocity, and the algorithm of the rheology calculation processing unit shown in FIG. 1 will be described. FIG. 7 illustrates a sensor structure for detecting a pulse wave phase velocity from reflected ultrasound accompanied by a Doppler shift with respect to arterial blood flow in the blood rheology measurement device according to the present invention by transmitting and receiving ultrasound from the living body surface. FIG. In the figure, two pairs of sensor parts, that is, sensor part A 41 and sensor part B 42 are formed on the same substrate 43. In the sensor part A41, a transmitting element A44 and a receiving element A45 are formed. In addition, a transmitting element B46 and a receiving element B47 are formed in the sensor unit B42. These sensors are mounted on the biological surface 48.
The incident ultrasonic wave A50 passing through the living body internal tissue 49 from the transmitting element A44 is reflected to the arterial blood flow A53 in the site A52 of the artery 51, and is electrically converted by the receiving element A45 as reflected ultrasonic wave A54 with Doppler shift. The Further, the incident ultrasonic wave B55 passing through the living body internal tissue 49 from the transmitting element B46 is reflected to the arterial blood flow B57 in the arterial part B56, and is electrically converted by the receiving element B47 as the reflected ultrasonic wave B58 with Doppler shift. Is done. At this time, the distance between the artery part A52 and the part B56 is the distance 59 between the measurement parts, which is equal to the distance between the sensor part A41 and the sensor part B42 formed on the substrate 43.
Further, the distance 59 between the measurement sites is L. In addition, transmission and reception of ultrasonic waves performed by the sensor unit A41 and the sensor unit B42 are performed simultaneously. FIG. 7 is a schematic diagram for measuring the blood flow velocity of an artery using ultrasonic waves. A sensor for irradiating light inside a living body and detecting a pulse wave phase velocity from the attenuation value. The structure is conceptually exactly the same. That is, in the case of volume pulse wave measurement, in the symbol of FIG. 7, the transmitting element is the light receiving element, the receiving element is the light receiving element, and the ultrasonic wave is replaced with light, and the reflected ultrasonic wave with the Doppler shift with respect to the blood flow is the volume pulse wave. It will replace the attenuated light due to the change.

  FIG. 8 is a conceptual diagram illustrating a sensor structure for detecting a pulse wave phase velocity from a motion displacement of a living body surface propagated from an artery that moves in synchronism with arterial blood flow in the blood rheology measurement device according to the present invention. In the figure, two pairs of displacement detection sensors, that is, a displacement detection element A 60 and a displacement detection element B 61 are formed on the same substrate 62. These sensors are attached to the living body surface 63. The arterial displacement at the part A65 of the artery 64 in the living body propagates through the living tissue 69 as the propagation wave A66, and similarly the arterial displacement at the part B67 of the artery 64 forms the living body surface wave 70 on the living body surface 63 as the propagation wave B68. This biological surface wave 70 is detected by the displacement detection element A60 and the displacement detection element B61. At this time, the distance between the part A65 and the part B67 is the distance 71 between the measurement parts, which is equal to the distance between the displacement detection element A60 and the displacement detection element B61 formed on the substrate 62. Further, the distance 71 between the measurement sites is L. In addition, the measurement of the biological surface wave by the displacement detection element A60 and the displacement detection element B61 is performed simultaneously.

FIG. 9 shows two types of blood flow velocity waveforms in the part A52 and the part 56 of the artery 51 detected from the reflected ultrasonic signal accompanied by the Doppler shift detected by the sensor structure described in FIG. That is, the blood flow velocity waveform at the region A52 is the blood flow velocity waveform A72, and the blood flow velocity waveform at the region 56 is the blood flow velocity waveform B73. The vertical axis is output intensity, and the horizontal axis is time. Both waveforms show periodic changes in synchronization with the heartbeat signal. In the figure, the blood flow velocity waveform A72 and the blood flow velocity waveform B73 both have N peak values in the time interval ΔT.
FIG. 10 shows a flowchart of calculation processing in the pulse wave phase velocity calculation processing unit 13 of the rheology calculation processing unit 12 according to the present invention. In the pulse wave phase velocity calculation process 13, first, in S <b> 1, first, the peak values of the blood flow velocity waveform A and the blood flow velocity waveform B that are input and their appearance times are measured. The results are summarized in Table 1.

That is, the peak value and the appearance time of the blood flow velocity waveform A are V _A (1) to V _A (N) and τ _A (1) to τ _A (N), and the peak value and the appearance of the blood flow velocity waveform B, respectively. The times are V _B (1) to V _B (N) and τ _B (1) to τ _B (N), respectively. In S2, the heartbeat angular frequency is calculated using the following equation (17) from the time difference between the peak appearance times at site A or site B.

In S3, the average delay time difference Δτ _AB is calculated from the peak appearance times of the parts A and B using the following equation (18).

In S4, the pulse wave phase velocity C is calculated using the following equation (19) from the average time difference Δτ _AB calculated in S3 and the distance L between the measurement sites shown in FIGS.

The above calculation processing is performed in the pulse wave phase velocity calculation processing unit 13.
The above calculation process is to measure the traveling wave component Ψ (t, z) given by the equation (10) between two points different by the distance L, and to measure the traveling wave phase velocity from the delay time difference. Don't be. 9, 10 and Table 1 are diagrams for explaining the calculation of the pulse wave phase velocity from the blood flow velocity waveform based on the ultrasonic Doppler effect. Needless to say, even when calculating the wave phase velocity, the same processing formula is used.
FIG. 11 shows an example of a flowchart of calculation processing in the rheology index calculation processing unit 14 of the rheology calculation processing unit 12 according to the present invention. The rheology index calculation processing unit 14 is provided with a memory unit 74. In this memory unit 74, the value of the pulse wave phase velocity Co when the kinematic viscosity ν of the blood, which varies with the race, sex, age, etc., of the subject using the blood rheology measuring device according to the present invention is 0, And the coefficient information of the polynomial function f (x) for determining the characteristic curve 40 shown in FIG. 6 is built in, and the function of setting numerical values according to the subject is provided. 11, firstly, _{C 0} values and polynomial function f (x) is set at S5, S6. Next, in S7, the normalized phase velocity _Cm is calculated according to the equation (13). S8, converting the normalized phase velocity _{C m} (14) to the β value in accordance with equation at. In S9, a polynomial function value f (β) with the β value as a variable is calculated. At this time, the polynomial function value f (β) is defined by equation (15). Finally, in S10, the polynomial function value f (β), the heartbeat angular frequency, and the equation (16) are used to calculate the μ value that is a rheological index.

  By measuring the pulse wave phase velocity of a living body, the present invention can detect not only blood rheology for the purpose of medical care and health maintenance / promotion, but also the activity state of the living body (human body) and the blood flow in each part of the living body. It can also be used in measurement to know the correlation between states.

1 is a block diagram for explaining a first embodiment of a blood rheology apparatus configuration according to the present invention. The block diagram explaining the 2nd Example of the blood rheology apparatus structure based on this invention The block diagram explaining the 2nd Example of the blood rheology apparatus structure based on this invention Schematic diagram of an artery that beats in sync with the heartbeat Characteristic diagram showing the relationship between the normalized phase velocity and α according to the present invention Characteristic diagram showing the relationship between β and α relating to the blood rheology index according to the present invention Sensor structure diagram for detecting pulse wave phase velocity from reflected ultrasound for arterial blood flow according to the present invention Sensor structure diagram for detecting pulse wave phase velocity from biological surface wave according to the present invention Blood flow velocity waveform charts at two different sites separated by a specific distance in the same artery The flowchart figure of the arithmetic processing in the pulse wave phase velocity arithmetic processing part which concerns on this invention The flowchart figure of the arithmetic processing in the rheology parameter | index arithmetic processing part which concerns on this invention Correlation diagram between μ value which is blood rheology index measured by blood rheology apparatus according to the present invention and whole blood passage time T of blood rheology index by blood collection method

Claims (5)

Blood flow velocity blood periodically pulsating in synchronism with the heartbeat of the living body is a velocity flow, or the volume pulse is an amount of flow of the blood, the measurement for measuring the living body outside as arterial blood flow information in the living A rheology calculation processing unit that processes the arterial blood flow information measured in the measurement unit,
The measurement unit includes a first sensor unit and a second sensor unit, and the arterial blood flow information measured in the first sensor unit and the second sensor unit is the same type of information,
The rheology calculation processing unit calculates a pulse wave phase velocity calculation unit based on the arterial blood flow information obtained from the first sensor unit and the second sensor unit; An index calculation processing unit that calculates an index of blood rheology from the pulse wave phase velocity detected in the pulse wave phase speed calculation processing unit;
The pulse wave phase velocity calculation processing unit calculates a heartbeat angular frequency (from a time difference between a peak value of the signal measured by the first sensor unit and a peak value of the signal measured by the second sensor unit. ω), pulse wave phase velocity (C),
The index calculation processing unit has a calculation processing function for converting the pulse wave phase velocity (C) into a normalized pulse wave phase velocity (Cm), and the normalized pulse wave phase velocity (Cm) is an index of blood rheology. Blood rheology measurement device characterized by being output as The pulse wave phase velocity (C) the blood rheology according to claim 1, characterized in that the second sensor portion and the first sensor portion is calculated from the volume pulse wave measured by using an optical measuring device. The pulse wave phase velocity (C) is according to claim 1, characterized in that the second sensor portion and the first sensor portion is calculated from the living body surface wave measured using a displacement detecting sensor Blood rheology measuring device. The blood rheology measuring apparatus according to any one of claims 1 to 3 , wherein the first sensor part and the second sensor part are formed on the same substrate. The pulse wave phase velocity calculation processing unit is configured to calculate a heart rate angle associated with a heart rate from a peak value of the signal measured by the first sensor unit and a peak value of the signal measured by the second sensor unit. A frequency (ω) and a pulse wave phase velocity (C) are calculated, and the index calculation processing unit converts the pulse wave phase velocity (C) into a normalized pulse wave phase velocity (Cm), and the normalized phase Speed (Cm)

A first arithmetic processing function for variable conversion, a second arithmetic processing function for calculating a polynomial function value (f (β)) of degree 1 or more with the β value as a variable, and the polynomial function value (f ( β)) and heart rate angular frequency (ω)

And with having a third arithmetic processing function of the variable transformation, blood rheology measuring apparatus according to any one of claims 1 to 4, characterized in that outputs the μ value as an index of blood rheology.

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