JP5726662B2 - Blood impedance measuring device, artificial dialysis device, and method for operating blood impedance measuring device - Google Patents

Description

The present invention relates to a blood impedance measuring apparatus provided in a blood extracorporeal circuit, an artificial dialysis apparatus including the blood impedance measuring apparatus, and a method for operating an apparatus for measuring the impedance of blood flowing through the extracorporeal circuit.

  It is known that the electrical impedance of blood reflects various characteristics of blood cells (particularly red blood cells) in the blood. For example, the conductivity of intracellular fluid (erythrocyte intracellular fluid electrolyte) and extracellular fluid (plasma electrolyte), cell membrane capacity, hematocrit (ratio of erythrocyte volume in the blood), red blood cell shape, size and It is known to reflect the orientation state and the like. For example, Patent Document 1 discloses a hematocrit measuring device that measures hematocrit by measuring the electrical impedance of blood.

  The hematocrit measuring device described in Patent Document 1 is a device for continuously measuring hematocrit in an extracorporeal circuit such as hemodialysis, and includes a resistance measurement cell disposed on a cylindrical insulating pipe that constitutes a blood flow path. Two types of alternating current sources (high frequency and low frequency current sources) and an alternating current voltmeter are provided. The resistance measuring cell includes a pair of low-frequency current supply electrodes provided at a predetermined distance in the central axis direction of the insulating pipe, and a pair of low-frequency current supply electrodes disposed adjacent to the low-frequency current supply electrodes. A high-frequency current supply electrode and a pair of voltage detection electrodes are provided. Then, when the low frequency current source supplies 50 kHz alternating current to the low frequency current supply electrode, the 50 kHz alternating voltage generated at the voltage detection electrode and the high frequency current source supply 20 MHz alternating current to the high frequency current supply electrode. The low frequency impedance and high frequency impedance of the blood are measured by measuring the 20 MHz AC voltage generated at the voltage detection electrode. Here, it is known that the hematocrit value of blood is proportional to the logarithm of the ratio of the low frequency impedance and high frequency impedance of blood, and using this relationship, the hematocrit value is calculated from the measured low frequency impedance and high frequency impedance. Seeking.

JP-A-63-133062

Medical Electronics and Biotechnology Volume 37 Special Issue (1999) P158

  In all the conventional measuring devices for measuring blood impedance (contact blood impedance measuring device) including the above-mentioned hematocrit measuring device, the current supply electrodes to which an alternating current is supplied are arranged so as to be in direct contact with blood. Has been. As will be described later, when the current supply electrode is arranged so as to be in direct contact with blood, it is possible to measure blood impedance with a high S / N ratio relatively easily. However, since the current supply electrode is in contact with blood, there is an increased risk of infection with viruses and bacteria. In addition, blood has the lowest electrical resistance in a living body, and blood vessels are connected to the heart through arteries or veins. For this reason, in the extracorporeal circuit for blood, it is not desirable to cause a current to flow through the blood while the current supply electrode is in direct contact with the blood, because there is a risk of electric shock due to leakage current. Further, if a current is kept flowing while the current supply electrode is in direct contact, metal may be deposited on the electrode surface or a blood clot may be generated in the blood.

  An object of the present invention is to provide a blood impedance measuring device that can reduce the risk of infection and electric shock, and can reduce the risk of metal depositing on the electrode surface or the formation of blood clots in blood. An artificial dialysis apparatus and a blood impedance measuring method are provided.

According to a first aspect of the present invention, there is provided a blood impedance measuring device for measuring impedance of blood flowing in a blood circuit circulating outside the body,
At least one pair of electrodes provided in the blood circuit;
A signal source that is connected to the at least one pair of electrodes and inputs an input electrical signal that varies with time to the at least one pair of electrodes;
A detection circuit connected to the at least one pair of electrodes and detecting an output electric signal induced in the at least one pair of electrodes when the time-variable electric input signal is inputted to the at least one pair of electrodes;
An arithmetic circuit that calculates blood impedance based on the input electrical signal and the output electrical signal;
The at least one pair of electrodes and an insulator that electrically insulates the blood flowing through the blood circuit, and the time-varying electrical input signal includes a low-frequency alternating current signal and a high-frequency alternating current signal. Including
The arithmetic circuit calculates the capacitance of the insulator from the magnitude of blood impedance measured using a low-frequency alternating current signal, and based on the calculated capacitance of the insulator, blood impedance measuring device that to calculate the impedance of blood is provided.

According to the blood impedance measuring device of the present invention, there is no fear of electric shock because the electrode and the blood flowing through the blood circuit are electrically insulated by the insulator. In addition, since the electrode does not come into direct contact with blood, there is no possibility that metal deposited on the electrode surface enters the blood. In addition, the capacitance of the insulator can be easily determined from the impedance value measured in the low frequency region. Therefore, it is possible to accurately calculate the value of the blood impedance by removing the component due to the capacitance of the insulator.

  In the impedance measuring apparatus of the present invention, the at least one pair of electrodes may include a pair of electrodes connected to both the signal source and the detection circuit. Alternatively, the at least one pair of electrodes may include a first electrode pair connected to the signal source and not connected to the detection circuit, and a second electrode pair connected to the detection circuit and not connected to the signal source. . In the former case, the pair of electrodes can serve both as an electrode pair for input electrical signals and an electrode pair for output electrical signals, thereby simplifying the structure of the impedance measuring device and reducing the manufacturing cost of the device. it can. In the latter case, an electrode pair for input electrical signals and an electrode pair for output electrical signals are provided independently. For example, when a current signal is used as the input electrical signal and a voltage signal is used as the output electrical signal, the first voltage drop due to blood impedance and a relatively large input signal current is generated in the electrode pair for the output electrical signal. And the impedance of the insulator directly under the electrode for the output electrical signal and the second voltage drop due to the weak current flowing in the voltage detection circuit are measured simultaneously. Here, the second voltage drop can be suppressed smaller than the first voltage drop. In other words, the voltage due to the impedance of the desired blood can be made relatively larger than the voltage due to the impedance of the capacitor of the insulator. Thereby, the influence of the electrostatic capacitance Cs of an insulator can be reduced and the impedance of blood can be measured accurately.

  In the blood impedance measuring apparatus of the present invention, the blood circuit may be constituted by a resin tube, the at least one pair of electrodes may be disposed outside the tube, and the insulator is the blood It may be the resin tube that defines the circuit. In this case, even when the blood circuit is a resin tube that is commonly used, it is possible to measure the impedance of blood in a non-contact manner by providing an electrode outside the tube.

  In the blood impedance measuring device of the present invention, the insulator may be formed as an insulating film that covers the surfaces of the at least one pair of electrodes, and the insulating film provided on the electrodes forms a part of the blood circuit. It may be defined and in contact with the blood. In this case, for example, a metal tube (pipe) whose inner surface is coated with a fluorine resin or the like can be used as the electrode and the insulator according to the present invention. Since the thickness of the insulating film (thickness of the coating) can be reduced relatively easily, the capacitance due to the insulating film can be kept small, and the blood impedance can be measured with high accuracy.

According to a second aspect of the present invention, an artificial dialysis device,
  A blood circuit that circulates blood,
  A dialysate circuit for circulating the dialysate;
  A dialysate supply device connected to the dialysate circuit and circulating the dialysate while supplying the dialysate;
  A blood pump connected to the blood circuit to circulate the blood;
  A dialyzer connected to the blood circuit and the dialysate circuit;
  An artificial dialysis device is provided comprising a blood impedance measuring device according to a first aspect of the present invention connected to a blood circuit.

  According to the second aspect of the present invention, blood impedance can be easily measured during artificial dialysis, and a rapid change in osmotic pressure that may occur during artificial dialysis can be monitored.

  According to a third aspect of the present invention, there is provided a method of operating an apparatus for measuring impedance of blood flowing in a blood circuit,
  The electrical signal input means of the device inputs an electrical input signal including a low frequency electrical input signal to blood flowing through the blood circuit through at least a pair of electrodes insulated by an insulator disposed between the blood. When,
  Detecting an output electric signal induced in the at least one pair of electrodes when the detection means of the device inputs the electric input signal to the at least one pair of electrodes;
  The capacitance calculating means of the device calculates the capacitance of the insulator based on the low frequency electrical input signal and a low frequency output electrical signal induced by the low frequency electrical input signal. When,
  An apparatus operating method for measuring blood impedance, wherein the impedance calculation means of the device comprises calculating the blood impedance based on the capacitance of the insulator, the electrical input signal, and the output electrical signal. Is provided.

  According to the present invention, since the electrode is not in direct contact with an electrolyte such as blood, there is no risk of metal deposition on the electrode surface. In addition, since the electrode is not in direct contact with an electrolyte such as blood, the risk of blood clots occurring in the blood can be reduced.

FIG. 1 is a schematic view of an artificial dialysis apparatus. (Reprinted from http://www.medi-net.or.jp/tcnet/dtx/004.html) FIG. 2a is a schematic diagram of a dialyzer (reproduced from http://www.nininkai.com/kiso.html). FIG. 2b is a schematic diagram (reproduced from http://mobara-cl.com/hemo_dialysis/hemo_dialysis.html) showing the function of the dialyzer. FIG. 3 is a schematic diagram of a blood impedance measuring apparatus having a pair of (two-terminal) electrodes. FIG. 4a is an equivalent circuit diagram of the contact model. FIG. 4b is an equivalent circuit diagram of the non-contact model. FIG. 5 is a model diagram of a living tissue. FIG. 6 is a Cole-Cole plot for hypotonic solution, whole blood and hypertonic solution. FIG. 7 is a schematic view of a contact blood cell. FIG. 8a shows the measurement result of | Z | for the contact-type blood cell. FIG. 8 b shows the measurement result of | Z | for the non-contact blood cell. FIG. 9a is a graph plotting the real part of the measured impedance. FIG. 9b is a graph plotting the imaginary part of the measured impedance. FIG. 10 is a schematic view of a blood impedance measuring apparatus having four terminal electrodes. FIG. 11 is a schematic view of a thin chamber. FIG. 12 is a schematic view of a thin chamber having concentric electrodes. FIG. 13 is a flowchart showing an impedance measurement method according to the present invention.

  As an example of a blood impedance measuring apparatus according to the present invention, a non-contact blood impedance measuring apparatus 100 attached to a hemodialysis artificial dialysis apparatus 200 shown in FIG. 1 will be described as an example. Here, the artificial dialyzer 200 includes a dialysate tank 201 for storing a concentrated dialysate (dialyte concentrate), a dialysate supply device 202, a monitoring device 203, a dialyzer 204, a crit line chamber 205, a blood pump. 206 is mainly provided. Furthermore, the artificial dialysis apparatus 200 includes a dialysate circuit 207 and a blood circuit 208, both of which are constituted by resin tubes.

  The dialysate tank 201 and the dialysate supply device 202 are connected by a tube, and the concentrated dialysate in the dialysate tank 201 is supplied to the dialysate supply device 202. The dialysate supply device 202 is supplied with deionized water (for example, tap water deionized through an ion exchange resin) made from tap water. The dialysate supply device 202 adjusts the concentrated dialysate to an appropriate concentration, and then sends the dialysate whose concentration has been adjusted (hereinafter simply referred to as dialysate) to the monitoring device 203 connected via a tube. The dialysate sent to the monitoring device 203 flows through a dialysate circuit 207 that returns to the monitoring device 203 through the dialyzer 204. The monitoring device 203 monitors the flow rate, temperature, and the like of the dialysate flowing in the dialysate circuit 207, and monitors biological information such as venous pressure of the dialysis patient. Further, as will be described later, blood hematocrit value and / or oxygen saturation, etc., measured by a sensor (not shown) attached to the crit line chamber 205 can be monitored.

  Further, blood drawn from the dialysis patient through the body flows through a blood circuit 208 that returns to the body through the blood pump 206 and the dialyzer 204. In order to prevent the blood drawn outside the body from coagulating in the blood circuit 208, an anticoagulant such as heparin is injected in the middle of the blood circuit 208.

  As shown in FIG. 2a, the dialyzer 204 includes a blood channel 204a formed of a plurality of hollow fiber-like semipermeable membranes provided in the central portion, and a dialysate flow provided outside the blood channel 204a. It has a double-structure flow path composed of a path 204b. The blood channel 204a and the dialysate channel 204b are separated by a semipermeable membrane that defines the blood channel 204a. Further, the blood flowing through the blood flow path 204a and the dialysate flowing through the dialysate flow path 204b flow in opposite directions. The semipermeable membrane that defines the blood flow path 204a has very small pores that allow uremic toxins such as urea, creatinine, and uric acid, and electrolytes such as sodium, potassium, and phosphorus to pass therethrough. Do not allow erythrocytes, leukocytes, proteins, bacteria, viruses, etc. to penetrate. Then, as shown in FIG. 2b, waste products such as uremic toxins in the blood are discharged into the dialysate through the semipermeable membrane, and the electrolyte deficient in the body is compensated from the dialysate side. In addition, since red blood cells, white blood cells, proteins, etc. in the blood cannot permeate the semipermeable membrane, they are not discharged from the blood side to the dialysate side.

  As described above, the crit line chamber 205 is provided upstream of the dialyzer 204 in the blood circuit 208. A thin plate-like transparent translucent portion 205a having a blood flow path formed therein is formed at the central portion of the crit line chamber 205. A sensor that optically measures the hematocrit value of blood and / or the oxygen saturation level of blood can be attached to the translucent portion 205a.

  Further, a non-contact blood impedance measuring apparatus 100 is disposed at a position adjacent to the blood circuit 208 on the upstream side of the crit line chamber 205. As shown in FIG. 3, the blood impedance measuring apparatus 100 includes two annular electrodes 101A and 101B and electrodes 101A and 101B provided so as to be wound around the outside of a resin tube constituting the blood circuit 208. A current generating circuit 102 for passing an alternating current I between the electrodes 101A and 101B, a voltage detecting circuit 103 for detecting an alternating voltage V generated between the electrodes 101A and 101B by the alternating current I, and an alternating current I, the AC voltage V, and the arithmetic circuit 104 for obtaining the impedance Z from these phase differences θ. In blood impedance device 100 shown in FIG. 3, an electrode for inputting alternating current I and an electrode for detecting alternating voltage V are not provided independently, and a pair of electrodes 101A, 101B. Has both functions.

  As described later, by measuring the frequency characteristics of the impedance of blood, water movement inside and outside erythrocytes, membrane permeability (ion permeability) of erythrocyte cell membrane, erythrocyte shape, size and orientation state, etc. / Physical property information can be obtained.

  In particular, since blood is circulated outside the body when performing artificial dialysis, patients undergoing artificial dialysis may be exposed to sudden changes in osmotic pressure. In such cases, it may be necessary to discontinue enforcement due to hypotension or other malfunctions. A sudden change in osmotic pressure first causes a change in the shape of red blood cells in the blood circulating throughout the whole body before the osmotic pressure balance over the whole body is observed, and also changes the frequency characteristics of the impedance of blood. Therefore, if the frequency characteristic of the blood impedance can be monitored and a sudden change in osmotic pressure can be detected, it is possible to take measures such as adjusting the osmotic pressure before the patient causes a systemic shock.

  In recent years, it has become clear that the membrane permeability of the cell membrane of erythrocytes changes due to diabetes. And it is becoming possible to estimate the blood glucose level from the change in membrane permeability of the cell membrane of erythrocytes. As described above, information on the membrane permeability of the cell membrane of erythrocytes can be obtained by measuring the frequency characteristics of the impedance of blood. By measuring the frequency characteristics of the impedance of blood and extracting information about the membrane permeability of the cell membrane of red blood cells, if the change in blood glucose level can be monitored during dialysis, the amount of blood glucose during dialysis can be appropriately determined. It becomes possible to control. Since it is said that 30% of dialysis patients have diabetic nephropathy, a blood glucose level monitor during dialysis can provide useful information for doctors.

  Next, blood impedance measurement using the blood impedance measuring apparatus 100 will be described. As described above, the electrodes 101A and 101B are arranged at a predetermined distance in the extending direction of the resin tube constituting the blood circuit 208. The electrodes 101A and 101B are electrically connected to the current generation circuit 102 and are also electrically connected to the voltage detection circuit 103. In addition, the electrodes 101A and 101B are not in direct contact with blood and are separated by a resin tube. Here, the equivalent circuit of blood is the extracellular fluid resistance Re corresponding to the cell interstitial fluid, the cell membrane capacity Cm corresponding to the cell membrane constituted by the phospholipid membrane, and the intracellular fluid resistance Ri corresponding to the intracellular fluid. Using elements, it can be represented as in FIG. This corresponds to a model of a living tissue as shown in FIG. Here, in the living tissue, a relatively low frequency alternating current flows through the extracellular fluid as shown by the solid line in FIG. This is because the cell membrane acts as a capacitor having a cell membrane capacitance Cm, and therefore it is difficult for a DC current and a relatively low frequency AC current to pass through. On the other hand, since a relatively high frequency alternating current can easily pass through the cell membrane, it not only flows through the extracellular fluid, but also passes inside the cell as shown by the dotted line in FIG. To flow.

  Here, when the electrodes 101A and 101B are in direct contact with blood, the measured impedance is the impedance Z at both ends of the equivalent circuit shown in FIG. 4a. ₁ It corresponds to. That is, the impedance to be measured is the blood impedance itself. On the other hand, when the electrodes 101A and 101B and blood are separated by a resin tube and do not contact each other as in the present embodiment, the measured impedance is the both ends of the equivalent circuit shown in FIG. 4b. Impedance Z ₂ It corresponds to. Here, the equivalent circuit shown in FIG. 4b is obtained by connecting a capacitance Cs corresponding to a resin tube constituting the blood circuit 108 in series to the blood equivalent circuit shown in FIG. 4a. In the following description, the impedance Z is appropriately determined. ₁ Equivalent circuit impedance Z of the contact model ₁ Called impedance Z ₂ Is equivalent circuit impedance Z of non-contact model ₂ Call it.

  Here, the equivalent circuit impedance Z of the contact model ₁ And the equivalent circuit impedance Z of the non-contact model ₂ Can be expressed as Equation 1 and Equation 2, respectively. Here, as can be seen by comparing Equation 1 and Equation 2, the equivalent circuit impedance Z of the contact model ₁ And the equivalent circuit impedance Z of the non-contact model ₂ Real part of Z ₁ } And Real {Z ₂ } Can be expressed by the same formula (see Equation 3).

  As described above, when a contact-type blood impedance measuring apparatus is used, the impedance to be measured is approximately equal to the impedance of blood. However, when the non-contact type blood impedance measuring apparatus 100 is used as in the present embodiment, the component of the second term of Equation 2, that is, the blood circuit, is used in order to obtain the blood impedance from the measured impedance. The influence of the component of the capacitance Cs caused by the resin tube constituting the 108 must be removed. Here, according to Equation 2, in the limit of ω → 0, the first term → Re, while the second term → ∞. From this, in the low frequency region, the impedance Z ₂ Size (| Z ₂ |) Shows that the contribution of the second term is dominant. That is, the impedance Z measured in the low frequency region ₂ Size (| Z ₂ Since |) is considered to be substantially equal to the contribution (1 / ωCs) due to the capacitance Cs, it is possible to estimate the capacitance Cs of the resin tube.

  And the impedance Z in the low frequency region ₂ If the electrostatic capacity Cs of the resin tube can be estimated by measuring the size of, the magnitude of the component of the second term of Equation 2 can be obtained, and the influence can be removed. In this way, the impedance Z measured by the non-contact blood impedance measuring apparatus 100 as in the present embodiment. ₂ From this, the impedance of blood can be determined.

  In summary, the method for measuring the impedance of blood using the non-contact blood impedance measuring apparatus 100 of the present embodiment is as follows (see FIG. 13). First, at least one pair of electrodes (electrodes 101A and 101B) is arranged in the blood circuit 108 so as to face the flow path in the blood circuit 108 via an insulator (S1). Next, an electrical input signal including a low-frequency electrical input signal (for example, an alternating current signal) is input to at least one pair of electrodes (S2). Here, the low frequency electric input signal is preferably a signal in a frequency region of about 10 kHz to 200 kHz, and more preferably a signal in a frequency region of about 50 kHz to 100 kHz. Here, an electrical input signal including a low frequency electrical input signal in such a frequency region and a higher frequency electrical input signal (for example, about several hundred kHz to 100 MHz) is input to the electrodes. Next, an electrical output signal induced in at least one pair of electrodes by the electrical input signal is detected (S3). The electrical output signal detected here includes a low-frequency electrical output signal corresponding to the above-described low-frequency electrical input signal and a high-frequency electrical output signal corresponding to the high-frequency electrical input signal. Next, the capacitance of the insulator is calculated based on the low frequency electrical input signal and the low frequency output electrical signal induced by the low frequency electrical input signal (S4). As described above, the contribution of the capacitance of the insulator is dominant in the magnitude of the impedance calculated from the low-frequency electrical input signal and the output electrical signal induced thereby. Using this, the capacitance of the insulator is calculated from the low-frequency electrical input signal and the output electrical signal induced thereby. Then, the impedance of the blood is calculated based on the capacitance of the insulator, the electrical input signal (in the high frequency region), and the output electrical signal (in the high frequency region) (S5). In this way, the blood impedance can be calculated by removing the influence of the capacitance of the insulator.

  Next, a method for extracting physiological information from the measured blood impedance will be described. Here, a so-called Cole-Cole plot can be created by obtaining an admittance that is the reciprocal of the measured blood impedance and plotting the admittance at each frequency on a complex plane. FIG. 6 shows an example of a Cole-Cole plot created for hypotonic solution, whole blood, and hypertonic solution. In the Cole-Cole plot, it is known that the admittance locus has an arc shape as shown in FIG. Here, the value on the horizontal axis at the left end of the arc indicates 1 / Re, and the value on the right end indicates (1 / Re) + (1 / Ri). In FIG. 6, the hypotonic solution is a liquid obtained by mixing blood with a saline solution having a concentration lower than that of physiological saline so that the osmotic pressure is lower than that of cells (for example, red blood cells). It is a liquid obtained by mixing blood with a higher concentration of saline than physiological saline so that the osmotic pressure is higher than that of cells (for example, red blood cells). Since the extracellular fluid resistance Re decreases in the order of hypotonic solution, whole blood, and hypertonic solution, the reciprocal 1 / Re value increases in this order. Correspondingly, in the Cole-Cole plot shown in FIG. 6, the hypotonic solution, whole blood, and hypertonic solution are arranged in this order from the left side.

  In the hypertonic solution, the water in the red blood cells is deprived by the extracellular fluid and the red blood cells contract. At this time, assuming that the hematocrit is constant, the intracellular fluid resistance Ri of the red blood cells increases. On the other hand, in the hypotonic solution, the red blood cells take the water of the extracellular fluid and the red blood cells expand. At this time, the intracellular fluid resistance Ri of the red blood cells is reduced. In this way, it is possible to estimate whether the red blood cells are expanding or contracting based on the magnitude of the intracellular fluid resistance Ri. This is useful in determining the presence or absence of cell edema.

  As described above, by creating the above-described Cole-Cole plot based on the measured blood impedance, it is possible to extract various information related to the extracellular fluid resistance Re and the intracellular fluid resistance Ri.

  Further, the frequency dependence of the real part and the imaginary part of the measured blood impedance is plotted, respectively, and the extracellular fluid resistance Re, the intracellular fluid resistance Ri in Equation 1 and
The extracellular fluid resistance Re, the intracellular fluid resistance Ri, and the cell membrane capacitance Cm may be obtained by performing fitting by a non-linear least square method using the function of the imaginary part and the real part of Equation 1 with the cell membrane capacity Cm as a parameter.

  In the above discussion, the impedance of blood is modeled using three parameters (extracellular fluid resistance Re, intracellular fluid resistance Ri, and cell membrane capacitance Cm). May be modeled. For example, a so-called two-time constant model in which a second intracellular fluid resistance Ri2 and a cell membrane capacitance Cm2 are added in parallel to the equivalent circuit may be used. For example, Non-Patent Document 1 shows that as the osmotic pressure increases, the two center dispersion frequencies obtained by f = 1 / (2πRiCm) move away and one decreases. As described above, this reflects that the cells contract and Ri increases. Also, for example, a model that characterizes blood impedance by parameters such as hematocrit, the lengths of the short and long axes of red blood cells, plasma conductivity, intracellular fluid conductivity, cell membrane dielectric constant, and cell thickness is adopted. Also good. Also in this case, the above parameters may be obtained by fitting a frequency plot of the measured blood impedance by a non-linear least square method. At this time, for example, when the hematocrit value of blood is optically measured by a sensor arranged in the crit line chamber, the measured value is used as the hematocrit value, and other parameters are obtained by the nonlinear least square method. May be.

  Next, in order to confirm the usefulness of the non-contact type blood impedance measuring apparatus 100, a contact type blood cell and a non-contact type blood cell, which will be described later, were prepared and the following comparative experiment was performed. FIG. 7 shows a contact blood cell 300 used in this experiment. The contact-type blood cell 300 includes a substantially rectangular parallelepiped resin (polycarbonate) cell 301, a blood channel 302 formed inside the cell 301, and a pair of electrodes 303 provided in the blood channel 302. . The blood flow path 302 includes an inlet 302a and an outlet 302b provided in the most upstream part and the most downstream part, a first blood reservoir 302c and a second blood reservoir 302d communicating with the inlet 302a and the outlet 302b, respectively. And a communication channel 302e that communicates with the second blood reservoirs 302c and 302d. The first and second blood reservoirs 302c and 302d are formed as cylindrical spaces having a diameter of about 20 mm. The communication flow path 302e is formed as a tubular flow path having an inner diameter of 6 mm and a length of 20 mm. The pair of electrodes 303 is a circular platinum electrode having a diameter of about 14 mm, and one electrode is provided at the bottom of each of the first and second blood reservoirs 302c and 302d.

  The blood flow path 302 inside the contact-type blood cell 300 described above was filled with porcine blood (whole blood), and an impedance analyzer (HP Technologies manufactured by Agilent Technologies, Inc., HP4294A) was connected to the electrode 303 to measure the impedance of whole blood. . The measured frequency range is 5 kHz to 100 MHz. Furthermore, when the blood flow path 302 is filled with a liquid (referred to as a hypotonic solution) obtained by mixing 0.6% saline with porcine blood, and 1.2% saline is mixed with porcine blood. The same impedance measurement was performed for the case where it was filled with a liquid (called hypertonic solution). As described above, it is considered that the red blood cells are expanded in the hypotonic solution and the red blood cells are contracted in the hypertonic solution as compared with the whole blood.

  Further, for comparison, in place of the contact-type blood cell 300, a non-contact-type blood cell (corresponding to the electrodes 101A and 101B and the resin tube blood circuit 208 in FIG. 3) is replaced with whole blood, hypotonic solution, and hypertonic solution. The same impedance measurement was performed for the case where it was satisfied.

  FIG. 8 is a graph plotting the absolute value (hereinafter, referred to as | Z |) of the impedance measured at each frequency. FIG. 8a shows the measurement results for the contact-type blood cell 300, and FIG. 8b shows the measurement results for the non-contact-type blood cell. Here, according to FIG. 8b, it can be seen that | Z | increases exponentially in a low frequency region of 0.1 MHz or less. As described above, this corresponds to the fact that the component of the second term on the right-hand side of Equation 2 increases dramatically, particularly in the low frequency region. That is, since the resin constituting the blood circuit 208 is arranged between the electrodes 101A and 101B and the measurement object such as whole blood, the resin portion has a low frequency region due to the capacitance of the resin portion. It is considered that the measured | Z | is increased. In other words, | Z | measured in the low frequency region is dominated by the component (the component of the second term on the right side of Equation 2) due to the capacitance of the resin portion described above. The capacitance of the resin portion can be obtained from the measured value of | Z | And the impedance of blood can be extracted by removing the component resulting from the electrostatic capacitance of the resin part. The capacitance of the resin portion can be obtained in advance by calibrating with a calibration aqueous solution with known electrical characteristics.

  Thus, in the measurement using the non-contact type blood cell, the contribution of the electrostatic capacity of the resin part constituting the blood circuit 208 is removed, and the real part and imaginary part of the measured impedance are plotted in FIG. 9a. , B. In addition, the result measured using the contact-type blood cell 300 as a comparison is also shown, respectively. Referring to FIG. 9a, the magnitude of the value of the real part (resistance component) of the impedance is slightly different when the contact type blood cell 300 and the non-contact type blood cell are used because the shape of the cell is different. However, the magnitude relationship of the resistance component in whole blood, hypotonic solution, and hypertonic solution, and the frequency dependence of the resistance component are almost the same regardless of which cell is used. Similarly, in FIG. 9b, the magnitude of the value of the imaginary part (reactance component) of the impedance is the same when the contact-type blood cell 300 and the non-contact-type blood cell are used. Therefore, although the values are deviated, the magnitude relationship of the reactance components in the whole blood, the hypotonic solution, and the hypertonic solution, and the frequency dependence of the reactance components are almost the same regardless of which cell is used. This difference in absolute value (value itself) is considered to be eliminated if a solution having a known conductivity, such as an aqueous potassium chloride solution, is measured in advance in each measurement cell, and each measurement cell is calibrated with the data. In particular, the frequency dependence of the reactance component shows that a peak is formed around a certain frequency. Further, it can be seen that the value of the peak frequency is almost the same when the contact blood cell 300 and the non-contact blood cell are used. The position of the peak frequency is shifted according to the difference between whole blood, hypotonic solution, and hypertonic solution. In addition, as described above, the extracellular fluid resistance Re, the intracellular fluid resistance Ri, and the cell membrane capacity Cm are used as parameters to perform fitting by the non-linear least square method to obtain the extracellular fluid resistance Re, the intracellular fluid resistance Ri, and the cell membrane capacity Cm. May be.

  As described above, it was found from the comparison experiment using the contact type blood cell 300 and the non-contact type cell that the impedance measurement of blood without contact is useful.

  The blood impedance measuring device of the present invention is not limited to the blood impedance measuring device 100 described above. For example, the blood impedance measuring apparatus 100 described above has a resin tube and a pair of annular electrodes 101A and 101B provided to wrap around the resin tube. However, the present invention is not necessarily limited to such a configuration. The configuration is not limited. For example, like the blood impedance measuring apparatus 100A shown in FIG. 10, two pairs of annular electrodes 101C, 101D, 101E, and 101F may be provided instead of the pair of annular electrodes 101A and 101B. . Here, the electrodes 101C and 101D make a pair, and the electrodes 101E and 101F make a pair. A current generation circuit 102 is connected to one pair of electrodes 101C and 101D, and a voltage detection circuit 103 is connected to the other pair of electrodes 101E and 101F. Thus, in blood impedance measuring apparatus 100A shown in FIG. 10, an input signal electrode for applying an alternating current and an output signal electrode for picking up an induced alternating voltage are independently provided. Is provided.

  Here, when an alternating current is passed through the target blood cell (corresponding to the electrodes 101C to 101F and the resin tube blood circuit 208 in FIG. 10), the voltage is caused by the impedance of the resin capacitor immediately below the electrodes 101C and 101D. Will occur. Furthermore, the current passing through this capacitor flows into the blood and generates a voltage due to the impedance of the blood. Here, the voltage due to the impedance of the resin capacitor can be subtracted by estimating its capacitance as in Equation 2, but if the voltage due to the blood impedance is weaker than the impedance of the resin capacitor There is a possibility that the measurement error will increase, and it is conceivable to reduce the measurement accuracy and S / N. Therefore, electrodes 101E and 101F for detecting output signals are provided inside the electrodes 101C and 101D separately from the electrodes 101C and 101D, and a voltage is detected through these electrodes. The electrodes 101E and 101F simultaneously measure the voltage generated by the impedance of the blood immediately below the resin and the voltage drop due to the impedance of the capacitor of the resin and the input current flowing through the voltage detection circuit. In general, the current flowing in the voltage detection circuit can often be made weak. Therefore, in the configuration of FIG. 10, in order to measure the voltage drop due to the blood impedance and the relatively large input signal current, the impedance of the resin directly under the output signal (detection) electrode, and the voltage drop due to the weak current, The voltage due to the impedance of the desired blood can be made relatively larger than the voltage due to the impedance of the resin capacitor. In other words, the influence of Cs in FIG. 4b can be reduced, and it becomes possible to approximate to FIG. 4a. This method of improving voltage detection is known as the four-electrode method. The influence of the insulating resin may be measured in advance with a known calibration electrolyte (potassium chloride solution or sodium chloride solution). Note that the electrodes do not necessarily have to be two pairs, and a plurality of input signal electrode pairs and a plurality of output signal electrode pairs can be used as necessary.

  As shown in FIG. 11, blood impedance measuring apparatus 100 is provided between two connecting portions 401 connected to blood circuit 208, and between two connecting portions 401, and a thin disc-shaped chamber portion 402, chamber portion 402 A thin chamber 400 having a pair of electrodes 403 provided on a circular surface (front surface and back surface) may be provided. In this case, since the chamber portion 402 has a thin shape, when the electrode 403 is connected to the current generation circuit 102 and the voltage detection circuit 103, the connection can be easily performed using a clip or the like. .

  Note that the thin chamber does not necessarily have a disc shape, and may have an arbitrary shape. For example, it may have a thin plate shape like a thin chamber 400A shown in FIG. Further, the electrode attached to the thin chamber is not necessarily provided so as to sandwich the thin chamber. For example, a thin chamber 400A shown in FIG. 12 may have a pair of electrodes 403A and 403B arranged concentrically instead of the pair of circular electrodes 403. The pair of concentric electrodes 403A and 403B are both connected to the current generation circuit 102 and the voltage detection circuit 103, but are connected to the current generation circuit 102 as in the blood impedance measurement device 100A described above. The electrode pair and the electrode pair connected to the voltage detection circuit 103 may be provided independently.

  Alternatively, the blood impedance measurement device does not necessarily have to be inserted in series with the blood circuit 208, and a side path is formed in the blood circuit 208, and the blood impedance measurement device may be provided in the side path. In this case, even if a malfunction occurs in the blood circuit 208 and hemolysis or a blood clot occurs and the blood impedance measuring device cannot be used, the blood purification treatment for the patient can be continued.

  In the above description, the electrode of the blood impedance measuring device is provided on the surface of the resin tube constituting the blood circuit, but the present invention is not necessarily limited to such a configuration. For example, a thin insulating film may be coated on the inner surface of a metal tube that also serves as an electrode. As the thin insulating film, for example, a fluorine resin can be used. Here, the value of the capacitance Cs included in the second term on the right-hand side of Equation 2 is such that as the area of the electrode is increased, the thickness of the insulator (or insulating film) between the electrode and blood is increased. The thinner it is, the smaller it can be. As described above, in order to extract the impedance of blood, it is necessary to set the size of the capacitance Cs to be as small as possible in advance, considering that the component due to the capacitance Cs must be removed. Is desirable. In that respect, when a thin insulating film is coated on the inner surface of a metal tube, the thickness of the insulating film can be made very thin, and the capacitance Cs can be kept small. it can.

  In the above description, the blood impedance measuring apparatus is disposed upstream of the crit line chamber 205. However, the present invention is not limited to such an arrangement, and is disposed at an arbitrary position of the blood circuit 208. be able to. For example, in order to monitor the state of the body, a portion immediately after blood collection (called a blood shunt in the case of artificial dialysis) is preferable, but the portion where the catheter is inserted is fixed stably so as not to bleed. Since it is necessary to continue paying, it is not preferable to add various devices beside it. From such a point of view, the blood is stably circulating in the blood circuit, and even if the impedance measuring device is arranged, it does not interfere with other devices (for example, as in the above embodiment). It is desirable to place it on the front side of the dialyzer. The blood impedance measuring device of the present invention may be provided in an extracorporeal circulating blood circuit drawn out of the human body (or animal), and is not necessarily provided in an artificial dialysis device. For example, the blood impedance measuring device of the invention can be provided in an oxygenator.

  In the above description, the blood impedance measuring device generates an alternating current having a different frequency through the current generation circuit, and detects an alternating voltage generated due to the alternating current, thereby detecting the blood impedance at each frequency. However, the present invention is not limited to this. The impedance may be calculated by applying an alternating voltage between the electrodes and measuring the alternating current flowing at that time with a current detection resistor arranged in series with the electrodes. In addition, the electrical signal input to the electrode is not necessarily a sine wave, and may be an electrical signal that varies with time. For example, the electrical impedance of the blood component may be obtained by applying a pulsed voltage between the electrodes, measuring the current flowing at that time with a current detection resistor in series with the electrodes, and Fourier transforming the time waveform of the current. This method is called TDR (Time Domain Reflectometry) method.

  By attaching the blood impedance measuring apparatus of the present invention to an artificial dialysis apparatus and measuring the impedance of blood during artificial dialysis, it is possible to monitor whether fluctuations in osmotic pressure occur in the blood. As a result, it is possible to take measures such as adjusting the osmotic pressure before the dialysis patient causes systemic shock due to a rapid fluctuation of the osmotic pressure.

  100 Blood impedance measuring device
  200 Artificial dialyzer
  204 dialyzer
  208 Blood circuit

Claims (7)

A blood impedance measuring device for measuring impedance of blood flowing through a blood circuit circulating outside the body,
At least a pair of electrodes provided in the blood circuit;
A signal source that is connected to the at least one pair of electrodes and inputs an input electrical signal that varies with time to the at least one pair of electrodes;
A detection circuit connected to the at least one pair of electrodes and detecting an output electric signal induced in the at least one pair of electrodes when the time-variable electric input signal is inputted to the at least one pair of electrodes;
An arithmetic circuit that calculates blood impedance based on the input electrical signal and the output electrical signal;
Comprising at least a pair of electrodes and an insulator for electrically insulating blood flowing through the blood circuit ;
The temporally varying electrical input signal includes a low frequency alternating current signal and a high frequency alternating current signal,
The arithmetic circuit calculates the capacitance of the insulator from the magnitude of blood impedance measured using a low-frequency alternating current signal, and based on the calculated capacitance of the insulator, blood impedance measuring device that to calculate the impedance of blood.   The blood impedance measuring device according to claim 1, wherein the at least one pair of electrodes includes a pair of electrodes connected to both the signal source and the detection circuit.   The at least one pair of electrodes includes a first electrode pair connected to the signal source and not connected to the detection circuit, and a second electrode pair connected to the detection circuit and not connected to the signal source. The blood impedance measuring apparatus as described. The blood circuit is composed of a resin tube,
The at least one pair of electrodes is disposed outside the tube;
The blood impedance measuring apparatus according to any one of claims 1 to 3, wherein the insulator is the resin tube that defines the blood circuit. The insulator is formed as an insulating film covering the surface of the electrode,
The blood impedance measuring device according to any one of claims 1 to 3, wherein the insulating film provided on the at least one pair of electrodes defines a part of the blood circuit and is in contact with the blood. An artificial dialysis machine,
A blood circuit that circulates blood,
A dialysate circuit for circulating the dialysate;
A dialysate supply device connected to the dialysate circuit and circulating the dialysate while supplying the dialysate;
A blood pump connected to the blood circuit to circulate the blood;
A dialyzer connected to the blood circuit and the dialysate circuit;
An artificial dialysis apparatus comprising the blood impedance measuring apparatus according to any one of claims 1 to 5 connected to a blood circuit. A method of operating a device for measuring the impedance of blood flowing through a blood circuit, comprising:
The electrical signal input means of the device includes a low-frequency alternating current signal and a high-frequency alternating current signal through the blood flowing through the blood circuit through at least a pair of electrodes insulated by an insulator disposed between the blood. Inputting an electrical input signal;
Detecting an output electric signal induced in the at least one pair of electrodes when the detection means of the device inputs the electric input signal to the at least one pair of electrodes;
The capacitance calculating means of the device calculates the capacitance of the insulator based on the low-frequency alternating current signal and a low-frequency output electrical signal induced by the low-frequency alternating current signal. When,
An apparatus operating method for measuring blood impedance, wherein the impedance calculation means of the device comprises calculating the blood impedance based on the capacitance of the insulator, the electrical input signal, and the output electrical signal. .

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