JP3860830B1 - Control device for gas supply mechanism for ventilator and control method of gas supply mechanism using control device - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide a ventilator control method capable of increasing a margin of stability in control of a gas supply mechanism of a ventilator.
The flow rate F of the support gas is detected, and the flow rate F of the support gas is estimated by an observer 54. Next, a flow rate deviation ΔF between the detected flow rate F and the estimated flow rate ^ F is obtained, information on the patient's spontaneous breathing pressure Pmus is obtained, and based on this information, a target pressure Pin for controlling the gas supply mechanism 20 is obtained. Calculate. A margin for the stability limit of the entire system 14 can be increased by calculating the target pressure Pin based on the flow rate deviation ΔF and appropriately selecting parameters by the transfer function for adjustment. As a result, even if the actual entire system fluctuates, it is possible to make it difficult for the runaway to occur. Further, the response of the assist can be improved as compared with the conventional PAV method.
[Selection] Figure 1

Description

The present invention relates to a control device for a gas supply mechanism used in a ventilator and a control method for the gas supply mechanism using the control device .

Proportional assist ventilation (proportional assisted ventilation, Proportional as a control method of the ventilator's gas supply mechanism during the inspiratory period when the patient breathes spontaneously
Assist Ventilation (abbreviated as PAV method).

  FIG. 22 is a block diagram showing the entire system 5 including a prior art ventilator 1 and a patient 2. A ventilator 1 that realizes the PAV method includes a gas supply mechanism 3 and a control device 4 that controls the gas supply mechanism 3. The control device 4 detects the flow rate of the support gas and determines the discharge pressure of the gas supply mechanism 3 based on the detected flow rate.

The control device 4 multiplies the first calculated value (K _fa · F) obtained by multiplying the flow rate gain K _fa by the flow rate F of the support gas and the volume V of the support gas supplied into the patient's lungs by the volume gain K _va. The calculated second calculated value (K _va · F / s) is obtained, and the first calculated value (K _fa · F) and the second calculated value (K _va · F / s) are added to calculate the target pressure Pin. . The flow rate gain K _fa is a value obtained by multiplying the estimated airway resistance R by the assist rate α, and the volume gain K _vg is a value obtained by multiplying the estimated lung elastance ^ E by the assist rate α. Here, let A be the assist rate α when ^ R = R and ^ E = E.

The control device 4 gives a signal representing the calculated target pressure Pin to the gas supply mechanism 3. The gas supply mechanism 3 to which the signal representing the target pressure Pin is given discharges the support gas at the support pressure Pvent based on the target pressure Pin. By appropriately setting the flow rate gain K _fa and the volume gain K _va , the gas supply mechanism 3 gives an assist pressure Pvent obtained by proportionally amplifying the patient's spontaneous breathing pressure Pmus (see, for example, Patent Document 1).

  FIG. 23 is a graph showing the relationship between the ventilation volume Vmus during spontaneous breathing and the ideal ventilation volume Vast during assist breathing in the entire system 5 of the prior art. The support pressure Pvent is amplified with an amplification factor of 1 / (1-A) times the spontaneous breathing pressure Pmus according to the time change of the spontaneous breathing pressure Pmus. As a result, the ventilation amount Vast at the time of assist breathing by the ventilator is amplified to 1 / (1-A) times the ventilation amount Vmus only by spontaneous breathing.

In the above-described prior art, it is necessary to determine the flow rate gain K _fa and the volume gain K _fg after accurately determining the airway resistance R of the patient and the elastance L of the lung. Therefore, another conventional technique is disclosed that determines the airway resistance R of the patient and the elastance E of the lung (for example, Patent Document 2).

Japanese Patent Publication No. 2714288 Japanese National Patent Publication No. 11-502755

  The entire system 5 of the prior art supplies the support gas to the patient at a pressure (Pvent + Pmus) obtained by adding the support pressure Pvent and the spontaneous breathing pressure Pmus in a positive feedback configuration. In the positive feedback configuration, the entire system may become unstable, and a phenomenon in which the assist pressure Pvent is amplified without converging, that is, an over assist called a so-called runaway may occur.

For example, in the runaway, when the flow gain _Kfa and the volume gain _Kva are not appropriate, the patient's condition fluctuates when there is an excessive delay from the time when the target pressure Pin is input until the gas supply mechanism 3 operates. This is likely to occur when other disturbances are given. When runaway occurs, the pressure of the support gas becomes excessive, which may damage the patient's lungs and airways, and the conventional technique has to forcibly stop the assist.

The reason why runaway is likely to occur is that the stability margin of the entire system 5 of the prior art is small and the transient response tends to become unstable. If the stability margin of the entire system 5 is small, even if the entire system is slightly changed, the stability limit may be exceeded and a runaway may occur. Further, it is necessary to set the flow rate gain K _fa and the volume gain K _va so that the entire system 5 does not become unstable, and the degree of freedom in gain selection is low, and it is difficult to adjust an appropriate gain. There's a problem.

Therefore, an object of the present invention is to provide a control device for a gas supply mechanism of a ventilator that can increase the stability margin in the entire system including the ventilator and a patient, and a control method for the gas supply mechanism using the control device. Is to provide.

The present invention 請 Motomeko 1 wherein, in response to a signal representative of the target pressure Pin, in support pressure Pvent support gas containing oxygen, a gas supply ventilator delivered to the patient's airway via the intake channel In the control device that controls the mechanism,
A flow rate detecting means for detecting a flow rate F of the support gas flowing through the intake pipe;
A delay compensation means for calculating the output according to a predetermined adjustment transfer function as a delay compensation pressure ^ Pm with the target pressure Pin as an input;
In response to the delay compensation pressure ^ Pm calculated by the delay compensation means, using the respiratory organ model that models the patient's respiratory organ, the assistance gas with the delay compensation pressure ^ Pm will flow through the inspiratory line A flow rate estimating means for calculating a gas flow rate as an estimated flow rate ^ F;
Deviation calculating means for calculating a flow deviation ΔF between the flow rate F detected by the flow rate detecting means and the estimated flow rate ^ F calculated by the flow rate estimating means;
A control amount calculating means for calculating a target pressure Pin in response to the flow rate deviation ΔF calculated by the deviation calculating means and providing a signal representing the target pressure Pin to the gas supply mechanism and the delay compensating means. ,
The adjustment transfer function includes a first-order lag element and includes a time delay element, and a time delay Lm in the time delay element is set as zero.
The constant Tm when the first-order lag element of the adjusting the transfer function is greater than the time constant Tc approximating the primary delay element included in the transfer function of the gas supply mechanism is employed in the gas supply mechanism, characterized in Rukoto It is a control device.

According to the first aspect of the present invention, the flow rate detection means detects the support gas flow rate F, and the flow rate estimation means calculates the support gas estimated flow rate F. The detected flow rate F of the assisting gas varies depending on the patient's spontaneous breathing pressure Pmus, but the estimated assisting gas flow rate F is not affected by the patient's spontaneous breathing pressure Pmus. Therefore, by obtaining a flow rate deviation ΔF between the detected assist gas flow rate F and the estimated assist gas flow rate F, information on the patient's spontaneous breathing pressure Pmus can be acquired. Here, the spontaneous breathing pressure Pmus is a value obtained by converting a respiratory effort, which is a force that a patient spends on breathing, into a pressure, and an accurate value cannot be measured unless a special method is used. In the present invention, the spontaneous breathing pressure Pmus that may be generated by the breathing effort is virtually determined in response to the flow deviation ΔF without using a special method.

  The control amount calculation means calculates a target pressure Pin for controlling the gas supply mechanism in response to the flow rate deviation ΔF. Therefore, the target pressure Pin is also a pressure corresponding to the patient's spontaneous breathing pressure Pmus. By providing the gas supply mechanism with a signal representing the target pressure Pin calculated in this way, the support gas can be supplied to the patient's airway at the support pressure Pvent corresponding to the spontaneous breathing pressure Pmus that changes sequentially.

  Further, by calculating the target pressure Pin based on the flow rate deviation ΔF, the entire system including the patient and the ventilator is compared with the prior art that calculates the target pressure Pin based only on the detected flow rate F of the assisting gas. This system can be made less likely to have a positive feedback configuration, and the margin for the stability limit of the entire system can be increased. If this causes disturbances, there is an excessive time delay in the gas supply mechanism, the patient's respiratory model cannot be set correctly, the patient's lung and airway conditions change, and the patient's respiratory condition changes Even in such a case, divergence of the support pressure Pvent, that is, so-called runaway can be made difficult to occur. Thus, by providing the support pressure Pvent proportional to the patient's spontaneous breathing pressure Pmus, pressure support according to the patient's breathing timing can be performed, and the burden on the patient can be reduced.

In the present invention, the delay compensation means calculates the delay compensation pressure ^ Pm according to the adjustment transfer function based on the target pressure Pin. The flow rate estimating means calculates the estimated flow rate ^ F based on the delay compensation pressure ^ Pm calculated by the delay compensating means. In this case, the dead time Lm of the dead time element constituting the adjustment transfer function is set to zero. In this way, by setting the time constant Tm of the first-order lag element appropriately with the dead time Lm being zero, a stable region can be obtained when the time constant of the first-order lag element of the adjustment transfer function is changed. Can be increased. In other words, it can maintain a stable state and increase the variable time constant. As a result, the time constant can be increased as much as possible within a stable range where the response of the support pressure Pvent does not vibrate or diverge, and the speed response of the support pressure Pvent can be further improved. In addition, when the respiratory organ model is not accurate with respect to the actual respiratory organ, even when the amplification factor is large, the time constant can be increased as much as possible within the stable range, and the responsiveness of the support pressure Pvent can be increased. This can be further improved. Furthermore, by setting the dead time Lm of the dead time element to zero, the parameters required for the adjustment transfer function can be reduced, and the adjustment transmission is compared with the case where both the dead time Lm and the time constant Tm are adjusted. An appropriate time constant Tm in the function can be easily obtained.
According to a second aspect of the present invention, the time constant Tm of the first-order lag element of the adjustment transfer function is configured to be adjustable.
The time constant Tm of the first-order lag element of the adjustment transfer function is larger than the time constant Tc approximating the first-order lag element included in the transfer function of the gas supply mechanism. As a result, when the transfer function of the entire system including the ventilator and the patient is expressed in the frequency domain, the differential gain in the high frequency response portion can be expressed in a form approximately proportional to Tc. Therefore, an increase in the time constant Tm of the adjustment transfer function corresponds to an increase in the differential gain in the high frequency response portion. When the differential gain in the high-frequency response portion increases, the rapid response of the support pressure Pvent during the patient's inhalation start period can be improved. Further, the time constant Tm of the adjustment transfer function is set to a value larger than the time constant Tc approximating the first-order lag element included in the transfer function of the gas supply mechanism, so that the dead time Lm of the adjustment transfer function is zero. Even if it makes it, it can improve a quick response.
By improving the speed response in this manner, the support pressure Pvent can be accurately amplified in proportion to the spontaneous breathing pressure Pmus that changes with time, and the asynchronous state between the ventilator and the patient can be increased. Can be suppressed. Here, it is preferable that the time constant Tm of the adjustment transfer function is set to be lower than a value at which the transient response of the support pressure Pvent becomes oscillating. This can further reduce the burden imposed on the patient by the ventilator.
According to the second aspect of the present invention, the time constant Tm of the first-order lag element of the adjustment transfer function is a time constant Tc approximating the first-order lag element included in the transfer function of the gas supply mechanism, and the transmission of the gas supply mechanism. It is characterized by being set smaller than a value (Tc + Lc) obtained by adding a dead time Lc approximating a dead time element included in the function.
According to the second aspect of the present invention, the time constant Tm can be increased as much as possible within a stable range in which the response of the support pressure Pvent is not oscillating or diverging, and the speed response of the support pressure Pvent is improved. Can do.

The present invention according to claim 3, constant Tm when the first-order lag element of the adjustment transfer function, depending on the amplification factor proportional amplifying spontaneous respiration pressure Pmus, characterized in that the upper limit is set .

According to the third aspect of the present invention, when the time constant Tm of the first-order lag element of the adjustment transfer function becomes excessively large and exceeds a predetermined upper limit value, the response of the support pressure Pvent becomes oscillating. As the amplification factor for proportionally amplifying the spontaneous breathing pressure Pmus increases, the upper limit value decreases. Accordingly, the upper limit value of the time constant Tm is determined according to the amplification factor, and the time constant Tm is set below the upper limit value, so that the response of the support pressure Pvent becomes oscillatory regardless of the amplification factor. Can be prevented.

According to a fourth aspect of the present invention, there is provided a gas supply mechanism for a ventilator for supplying a support gas containing oxygen to an inspiratory line at a support pressure Pvent in response to a signal representing a target pressure Pin. And a control method for controlling using a control device including an interface,
A delay compensation pressure calculation step of calculating, as a delay compensation pressure ^ Pm, an output in accordance with a predetermined adjustment transfer function for compensating for the delay , with the target pressure Pin as an input by an arithmetic processing circuit ;
In response to the delay compensation pressure ^ Pm calculated by the delay compensation pressure calculation step by the arithmetic processing circuit, the support gas of the delay compensation pressure ^ Pm is supplied to the intake pipe using a respiratory organ model that models the respiratory organ of the patient. A flow rate estimating step of calculating the flow rate of the support gas that will flow through the road as an estimated flow rate ^ F;
A deviation calculating step of calculating a flow rate deviation ΔF between the flow rate F of the assisting gas flowing through the intake pipe detected by the flow rate detecting means and the estimated flow rate of assisting gas ^ F calculated by the flow rate estimating step. When,
By the processing circuit, in response to the flow rate difference ΔF computed by the deviation calculating step, thereby calculating the target pressure Pin, through the interface a signal representing the target pressure Pin, the control amount to be supplied to the gas supply mechanism and the calculation step seen including,
The adjustment transfer function used in the delay compensation pressure calculation step includes a first-order lag element and also includes a dead time element, and the time constant Tm in the first-order lag element is determined by the transmission of the gas supply mechanism. Control of the gas supply mechanism using the control device, wherein a value larger than the time constant Tc approximating the first-order lag element included in the function is adopted, and the dead time Lm in the dead time element is set as zero Is the method.

According to the fourth aspect of the present invention, the flow rate F of the support gas is detected in the flow rate detection step, and the estimated flow rate F of the support gas is calculated in the flow rate estimation step. The detected flow rate F of the assisting gas varies depending on the patient's spontaneous breathing pressure Pmus, but the estimated assisting gas flow rate F is not affected by the patient's spontaneous breathing pressure Pmus. Therefore, by obtaining a flow rate deviation ΔF between the detected assist gas flow rate F and the estimated assist gas flow rate F, information on the patient's spontaneous breathing pressure Pmus can be acquired.

  In the control amount calculation step, a target pressure Pin for controlling the gas supply mechanism is calculated based on the flow rate deviation ΔF. Therefore, the target pressure Pin is also a pressure corresponding to the patient's spontaneous breathing pressure Pmus. By providing the gas supply mechanism with a signal representing the target pressure Pin calculated in this way, the support gas can be supplied to the patient's airway at the support pressure Pvent corresponding to the spontaneous breathing pressure Pmus that changes sequentially.

  Further, by calculating the target pressure Pin in response to the flow rate deviation ΔF, the patient and the ventilator are included as compared with the conventional technique that calculates the target pressure Pin based only on the detected flow rate F of the support gas. It is possible to make the whole system less likely to have a positive feedback configuration, and it is possible to increase a margin for the stability limit of the whole system. If this causes disturbances, there is an excessive time delay in the gas supply mechanism, the patient's respiratory model cannot be set correctly, the patient's lung and airway conditions change, and the patient's respiratory condition changes Even in such a case, divergence of the support pressure Pvent, that is, so-called runaway can be made difficult to occur. Thus, by providing the support pressure Pvent proportional to the patient's spontaneous breathing pressure Pmus, pressure support according to the patient's breathing timing can be performed, and the burden on the patient can be reduced.

In the present invention, in the delay compensation pressure calculation step, the delay compensation pressure ^ Pm is calculated according to the adjustment transfer function based on the target pressure Pin. In the flow rate estimation step, the estimated flow rate ^ F is calculated based on the delay compensation pressure ^ Pm calculated in the delay compensation pressure calculation step. In this case, the dead time Lm of the dead time element constituting the adjustment transfer function is set to zero. Thus the dead time Lm of dead time element while zero, by appropriately setting the constant Tm when lag elements, stable in the case of changing the time constant of the lag element of the adjusting the transfer function The possible area can be increased. In other words, it can maintain a stable state and increase the variable time constant. As a result, the time constant can be increased as much as possible within a stable range where the response of the support pressure Pvent does not vibrate or diverge, and the speed response of the support pressure Pvent can be further improved. In addition, when the respiratory organ model is not accurate with respect to the actual respiratory organ, even when the amplification factor is large, the time constant can be increased as much as possible within the stable range, and the responsiveness of the support pressure Pvent can be increased. This can be further improved. Furthermore, by setting the dead time Lm of the dead time element to zero, the parameters required for the adjustment transfer function can be reduced, and the adjustment transmission is compared with the case where both the dead time Lm and the time constant Tm are adjusted. An appropriate time constant Tm in the function can be easily obtained.
The time constant Tm of the first-order lag element of the adjustment transfer function is larger than the time constant Tc approximating the first-order lag element included in the transfer function of the gas supply mechanism. As a result, when the transfer function of the entire system including the ventilator and the patient is expressed in the frequency domain, the differential gain in the high frequency response portion can be expressed in a form approximately proportional to Tc. Therefore, an increase in the time constant Tm of the adjustment transfer function corresponds to an increase in the differential gain in the high frequency response portion. When the differential gain in the high-frequency response portion increases, the rapid response of the support pressure Pvent during the patient's inhalation start period can be improved. Further, by adopting a value larger than the time constant Tm by setting the dead time Lm to zero, the quick response can be improved.
By improving the speed response in this manner, the support pressure Pvent can be accurately amplified in proportion to the spontaneous breathing pressure Pmus that changes with time, and the asynchronous state between the ventilator and the patient can be increased. Can be suppressed. Here, it is preferable that the time constant Tm of the adjustment transfer function is set to be lower than a value at which the transient response of the support pressure Pvent becomes oscillating. This can further reduce the burden imposed on the patient by the ventilator.

According to the first aspect of the present invention, the rapid response of the support pressure Pvent can be improved, the support pressure Pvent can be proportionally amplified with high accuracy, and the asynchronous state can be suppressed. As a result, a support pressure close to ideal can be applied to the patient. That is, the assist pressure can be reduced when the patient's spontaneous breathing pressure is small, and the assist pressure can be increased when the spontaneous breathing pressure is large, thereby reducing the burden on the patient. The time constant Tm of the adjustment transfer function is preferably set lower than a value at which the transient response of the support pressure Pvent becomes oscillating. This can further reduce the burden imposed on the patient by the ventilator.

In addition, by setting the dead time Lm included in the adjustment transfer function to zero, if the respiratory organ model is not accurate with respect to the actual respiratory organ, the time constant can be set within a stable range even when the amplification factor is large. It can be increased as much as possible, and the responsiveness of the support pressure Pvent can be further improved. This can further reduce the load that the ventilator places on the patient. Furthermore, by setting the dead time Lm of the dead time element to zero, it is possible to reduce the parameters required for the transfer function for adjustment, and at an appropriate time compared with the case where both the dead time Lm and the time constant Tm are adjusted. The constant Tm can be easily obtained.
Further, by adopting a value larger than the time constant Tm by setting the dead time Lm to zero, the quick response can be improved.
By improving the speed response in this manner, the support pressure Pvent can be accurately amplified in proportion to the spontaneous breathing pressure Pmus that changes with time, and the asynchronous state between the ventilator and the patient can be increased. Can be suppressed. Here, it is preferable that the time constant Tm of the adjustment transfer function is set to be lower than a value at which the transient response of the support pressure Pvent becomes oscillating. This can further reduce the burden imposed on the patient by the ventilator.
According to the second aspect of the present invention, the time constant Tm can be increased as much as possible within a stable range in which the response of the support pressure Pvent is oscillating and does not diverge, and the speed response of the support pressure Pvent is improved. Can do.

According to the third aspect of the present invention, even when the amplification factor is increased, the time constant is increased as much as possible to improve the responsiveness of the support pressure Pvent, and the response becomes oscillatory. Can be prevented. This can further reduce the load that the ventilator places on the patient.

According to the present invention described in claim 4 , by adjusting the parameters of the adjustment transfer function, the entire system including the ventilator and the patient can be adjusted so as to increase the area where the negative feedback configuration is provided. . By adopting a negative feedback configuration for the entire system, it is possible to increase the stability margin as compared to the case where the entire system has a positive feedback configuration, and to prevent the support pressure Pvent from diverging. For example, by appropriately selecting the parameters, it is possible to improve the quick response of the transient response of the support pressure, and it is possible to prevent the support pressure from becoming oscillating. This can reduce the load that the ventilator places on the patient.

In addition, by setting the dead time Lm included in the adjustment transfer function to zero, if the respiratory organ model is not accurate with respect to the actual respiratory organ, the time constant can be set within a stable range even when the amplification factor is large. It can be increased as much as possible, and the responsiveness of the support pressure Pvent can be further improved. This can further reduce the load that the ventilator places on the patient. Furthermore, by setting the dead time Lm of the dead time element to zero, it is possible to reduce the parameters required for the transfer function for adjustment, and at an appropriate time compared with the case where both the dead time Lm and the time constant Tm are adjusted. The constant Tm can be easily obtained. As the time constant Tm of the adjustment transfer function, a value larger than the time constant Tc approximating the first-order lag element included in the transfer function of the gas supply mechanism is adopted, so that the dead time Lm of the adjustment transfer function is zero. Even if it does, quick response can be improved.

  FIG. 1 is a block diagram showing a ventilator 17 and a patient 18. The ventilator 17 includes a ventilator gas supply mechanism 20 and a control device 21 that controls the gas supply mechanism 20. The gas supply mechanism 20 supplies a support gas 16 containing oxygen to the patient's airway 15. The support gas 16 is, for example, pressurized air in the atmosphere. The gas supply mechanism 20 is gas supply means such as a pump, for example, and can control the pressure of the assist gas to be discharged.

When the patient 18 performs spontaneous breathing, as a control method of the gas supply mechanism 20 during the inspiration period, a proportional assist ventilation method (proportional assist ventilation, Proportional
Assist Ventilation (abbreviated as PAV method). The control device 21 of the present embodiment controls the gas supply mechanism 20 in accordance with the original purpose of the PAV method. The gas supply mechanism 20 supplies the support gas 16 to the patient's airway 15 at a support pressure Pvent proportional to the spontaneous breathing pressure Pmus. The spontaneous breathing pressure Pmus is a value obtained by converting a respiratory effort, which is a force acting from the outside of the lungs caused by the movement of respiratory muscles such as the diaphragm, into pressure. In the present embodiment, the support pressure Pvent is treated as being approximately equal to the discharge pressure of the gas supply mechanism 20.

  The gas supply mechanism 20 controlled according to the PAV method supplies the support gas 16 to the patient at a higher pressure as the patient 18 sucks the support gas 16 more strongly. Further, as the suction force of the patient 18 becomes weaker, the pressure of the support gas 16 to be supplied is lowered, and the supply of the support gas 16 is stopped while the patient finishes sucking the support gas. By controlling the gas supply mechanism 20 in this manner, the support gas 16 can be supplied at a pressure corresponding to the respiratory effort of the patient 18, and the burden on the patient 18 in the breathing operation can be reduced.

  The control device 21 calculates a target pressure Pin corresponding to the patient's spontaneous breathing pressure Pmus and gives the target pressure Pin to the gas supply mechanism 20. The gas supply mechanism 20 provided with the target pressure Pin supplies the support gas 16 to the patient's airway 15 at the support pressure Pvent corresponding to the patient's spontaneous breathing pressure Pmus. In the embodiment of the present invention, the transfer function to which “(s)” is attached indicates that it is a transfer function in the Laplace region, and the value to which “^” is attached is not an actual value but an estimated value or “S” indicates a Laplace operator.

  The control device 21 includes a flow rate detection means 50, an estimation means 51, a deviation calculation means 52, and a control amount calculation means 53. The flow rate detection means 50 detects the flow rate F of the support gas 16 actually supplied to the patient's airway 15. Hereinafter, the flow rate detected by the flow rate detection means 50 is referred to as a detected flow rate F. The detected flow rate F is the flow rate of the gas flowing from the gas supply mechanism 20 through the inspiratory conduit 25, and is approximated to be equal to the flow rate of the gas flowing through the patient's airway. Since this detected flow rate F varies due to the influence of the spontaneous breathing pressure Pmus, it becomes the flow rate of the respiratory system in a state where the spontaneous breathing pressure Pmus is applied.

  The flow rate detection means 50 measures the flow rate of the support gas 16 flowing through the intake pipe line 25. The intake pipe 25 is a pipe that guides the support gas 16 from the pressure source of the gas supply mechanism 20 to the patient's airway. For example, the flow rate detection means 50 is realized by a differential pressure type flow meter. When the flow rate detection unit 50 detects the flow rate F of the support gas, the flow rate detection unit 50 gives the detected flow rate F to the deviation calculation unit 52.

  The estimation means 51 has an observer 54 indicating a respiratory organ model that is modeled by simulating a patient's respiratory tract. The observer 54 serves as a flow rate estimating means for calculating the flow rate {circumflex over (F)} of assisting gas that will be supplied to the patient. Specifically, the observer 54 determines the amount of the support gas flowing through the intake pipe 25 when the support gas having the predetermined delay compensation pressure ^ Pm is supplied to the intake pipe 25 in the absence of the spontaneous breathing pressure Pmus. Estimate the flow rate.

  Hereinafter, the flow rate estimated by the estimation means 51 is referred to as an estimated flow rate ^ F. The estimated flow rate {circumflex over (F)} is a delay compensation pressure {circumflex over (P)}, which will be described later, and is the flow rate of the respiratory system when support gas is given to the respiratory system. When the estimation unit 51 estimates the flow rate {circumflex over (F)} of the support gas, the estimation unit 51 gives a signal representing the estimated flow rate {circumflex over (F)} to the deviation calculation unit 52.

The deviation calculation means 52 calculates a flow rate deviation ΔF that is a value obtained by subtracting the estimated flow rate ^ F from the detected flow rate F, and gives the calculation result to the control amount calculation means 53. The control amount calculator 53 gives a preset gain (K _FG + K _VG / s) to the flow rate deviation ΔF, and calculates a target pressure Pin for generating the support pressure Pvent.

  The control amount calculation means 53 gives a signal representing the calculated target pressure Pin to the estimation means 51 and the gas supply mechanism 20, respectively. The gas supply mechanism 20 supplies the support gas 16 to the patient's airway 15 at the discharge pressure based on the signal representing the target pressure Pin given from the control amount calculation means 53, that is, the support pressure Pvent. Further, the estimating means 51 sequentially calculates the estimated flow rate ^ F based on the signal representing the target pressure Pin given from the control amount calculating means 53.

  FIG. 2 is a block diagram specifically showing the entire system 14 according to the embodiment of the present invention. The estimation unit 51 further includes a delay compensation unit 55 in addition to the observer 54. The delay compensation unit 55 is provided to compensate for delay elements such as a primary delay element and a dead time element of each component constituting the entire system 14 such as a delay element of the gas supply mechanism 20 and a delay element of the air circuit. It is done. The delay compensation unit 55 is a delay compensation unit that calculates an output in accordance with an adjustment transfer function that can be parameter-adjusted with the target pressure Pin as an input, and the control characteristics of the entire system 14 including the ventilator 17 and the patient 18. Provided to improve.

  The delay compensation unit 55 receives the target pressure Pin as an input, and calculates an output according to an adjustment transfer function capable of parameter adjustment as a delay compensation pressure ^ Pm. The delay compensation unit 55 gives a signal representing the calculated delay compensation pressure ^ Pm to the observer 54. In the present embodiment, the adjustment transfer function includes a control element that approximates the transfer function of the gas supply mechanism 20 that is measured with the target pressure Pin as an input and the support pressure Pvent as an output. Have different parameters for determining the control element.

In the present embodiment, the transfer function of the gas supply mechanism 20 is represented by the product of the first-order lag element Gc (s) and the dead time element e ^{−Lc · s} . The adjustment transfer function Gm (s) · e ^{−Lm · s} is also expressed by the product of the first-order lag element Gm (s) and the dead time element e ^{−Lm · s} . The time constant Tm of the first-order lag element of the adjustment transfer function is set larger than the time constant Tc of the first-order lag element of the gas supply mechanism 20 (Tm> Tc). The dead time Lm of the transfer function for adjustment is set to substantially the same value as the dead time Lc of the gas supply mechanism 20 (Lm≈Lc or Lm = Lc). In the present invention, “substantially the same” includes the same case.

  The observer 54 estimates the estimated flow rate ＦF of the support gas when the support gas is supplied to the airway 15 of the patient 18 with the delay compensation pressure ＰPm based on the respiratory organ model of the patient. The observer 54 includes a subtractor 56, an estimated flow rate calculator 57, a support gas volume calculator 58, and an alveolar pressure calculator 59.

  The subtractor 56 receives a signal representing the delay compensation pressure ^ Pm from the delay compensation unit 55 and a signal representing the calculated alveolar pressure ^ Palv from the alveolar pressure calculation unit 59. The subtractor 56 subtracts the calculated alveolar pressure ^ Palv from the delay compensation pressure ^ Pm, and gives a signal representing the value to the estimated flow rate calculator 57. The calculated alveolar pressure ^ Palv will be described later.

  The estimated flow rate calculator 57 divides the subtracted value subtracted by the subtractor 56 by a preset estimated airway resistance ^ R, and calculates the divided value as an estimated flow rate ^ F. The estimated flow rate calculator 57 gives a signal representing the calculation result to a deviation calculation means 52 and a support gas volume calculator 58 described later.

  The estimated airway resistance ^ R is a value obtained by estimating the airway resistance R of the patient, and is set in advance by, for example, a medical staff. In addition, the estimated airway resistance R may be set in advance by a detection value detected by the measuring device. Further, in the entire system 14 of the present embodiment, the estimated airway resistance ^ R does not need to be exactly matched to the actual patient's airway resistance R.

  The support gas volume calculator 58 sequentially accumulates the estimated flow rate ^ F calculated by the estimated flow rate calculator 57 from the support gas supply start time, and calculates the integrated value as the support gas volume ^ V. The assist gas volume calculator 58 is a so-called integrator. Hereinafter, the volume of the support gas calculated by the support gas volume calculator 58 is referred to as a calculation volume ^ V, and is distinguished from the actual volume V of the support gas.

  The alveolar pressure calculator 59 multiplies the calculation volume {circumflex over (V)} by a preset lung estimated elastance {circumflex over (E)} E, and calculates the multiplication value as the calculated alveolar pressure ^ Palv. The alveolar pressure calculator 59 gives the calculated calculated alveolar pressure ^ Palv to the subtractor 56. The calculated alveolar pressure ^ Palv is a value obtained by estimating the pressure in the alveoli, and is distinguished from the actual alveolar pressure Palv.

  The estimated elastance ^ E of the lung is a value obtained by estimating the elastance E representing the elastic force of the patient's lung, and is preset by, for example, a medical staff. Further, the estimated lung elastance ^ E may be set in advance by a detection value detected by a measuring instrument such as a ventilation mechanics inspection apparatus. Also, in the overall system 14 of the present embodiment, the lung elastance ^ E does not have to be exactly matched to the actual patient's lung elastance E.

When the support gas 16 flows through the airway 15, a pressure loss approximately proportional to the flow rate F of the support gas 16 occurs, and the pressure in the lung becomes lower than the airway pressure Paw. The airway resistance R represents the relationship between the flow rate F of the support gas 16 and the pressure loss. A value (F · R) obtained by multiplying the flow rate F of the support gas 16 by the airway resistance R is a loss pressure caused by the duct resistance of the airway 15. For example, the general airway resistance R is 5 to 30 (cmH ₂ 0) / (liter / second). However, the airway resistance R varies greatly depending on the patient's condition.

When the support gas 16 is supplied into the lung, the alveolar pressure Palv increases almost in proportion to the increase in the volume V of the support gas 16 supplied into the lung. The lung elastance E represents the relationship between the volume V of the support gas 16 and the alveolar pressure Palv. A value (V · E) obtained by multiplying the volume V of the support gas 16 by the lung elastance E is the alveolar pressure Palv. The alveolar pressure Palv is a pressure against the inflow of the support gas 16. For example, typical lung elastance E is 1/20 to 1/50 (milliliter) / (cmH ₂ O). However, lung elastance E varies greatly depending on the patient's condition.

  Based on such characteristics of the respiratory tract, a respiratory organ model possessed by the observer 54 is set. The respiratory organ model possessed by the observer 54 is set to the following relationship.

  That is, the respiratory organ model of the observer 54 is a model of the respiratory organ of the patient when the spontaneous respiratory pressure Pmus is zero. In this model, the value obtained by subtracting the calculated alveolar pressure ^ Palv from the delay compensation pressure ^ Pm is equal to the value obtained by multiplying the estimated flow rate ^ F and the estimated airway resistance ^ R. Further, the value obtained by integrating the estimated flow rate ^ F from the support gas supply start time is equal to the calculation volume ^ V. The calculated alveolar pressure ^ Palv is equal to a value obtained by multiplying the estimated lung elastance ^ E and the calculated volume ^ V.

Therefore, when the calculated airway pressure ^ Paw is an input value and the estimated flow rate ^ F is an output value, the transfer function G (s) ₅₄ of the observer 54 is expressed as follows.
G (s) ₅₄ = s / (^ R · s + ^ E) (4)

  Here, ^ R represents the estimated airway resistance, and ^ E represents the estimated elastance of the lung. The other expressions have the same meaning with respect to the symbols shown in the above expressions. The respiratory organ model possessed by such an observer 54 is an example of implementation, and may be another model obtained by modeling the respiratory organ of a patient.

The control amount calculation means 53 has a first calculation value (K _FG · ΔF) obtained by multiplying the flow rate deviation ΔF calculated by the deviation calculation means 52 by a flow rate gain K _FG that is a coefficient set in advance, and the support gas supply start time. To obtain a second calculated value (K _VG · ΔF / s) obtained by multiplying a value obtained by sequentially integrating the flow rate deviation ΔF by a volume gain K _VG that is a coefficient set in advance, and obtain a first calculated value (K _FG · ΔF) And the target pressure Pin related to the support pressure Pvent is calculated by adding the second calculation value (K _VG · ΔF / s). When the flow rate deviation ΔF is an input value and the target pressure Pin is an output value, the transfer function G (s) ₅₃ of the control amount calculation means 53 is shown below.
G (s) ₅₃ = K _FG + K _VG / s (5)

Here, K _FG represents a flow rate gain, and K _VG represents a volume gain. The other expressions have the same meaning with respect to the symbols shown in the above expressions. For example, the flow gain K _FG is set to a value (^ R · β _FG ) obtained by multiplying the estimated airway resistance ^ R by a predetermined flow amplification gain β _FG , and the volume gain K _VG is the estimated lung elastance ^ E. _Is multiplied by a predetermined volume amplification gain β _VG (^ E · β _VG ). When the flow rate amplification gain β _FG and the volume amplification gain β _VG are set to the same value, they are simply referred to as amplification gain β. Further, B represents the amplification gain β when ^ R = R and ^ E = E. By adjusting the flow rate gain K _FG in this way, the quick response to the spontaneous breathing pressure Pmus can be further improved, and by adjusting the volume gain K _VG , the steady gain of the target pressure Pin can be adjusted. . By making it possible to individually adjust the flow rate gain _KFG and the volume gain _KVG , it is possible to set the target pressure Pin by improving the control characteristics such as quick response and attenuation in combination with the steady gain.

  The estimation unit 51, the deviation calculation unit 52, and the control amount calculation unit 53 described above have been individually described for easy understanding, but the transfer functions may be equivalently converted and arranged. Further, the estimating means 51, the deviation calculating means 52, and the control amount calculating means 53 may be realized by a computer capable of numerical calculation executing a predetermined operation program.

In one embodiment of the present invention, the transfer function of the gas supply mechanism 20 includes a dead time element. In FIG. 2, the transfer function Gc (s) excluding the time ^delay element and the transfer function e ^{−Lc · s of the} time ^delay element in the transfer function of the gas supply mechanism 20 are individually illustrated. A transfer function G (s) ₂₀ that approximates the characteristics of the gas supply mechanism 20 when the target pressure Pin is an input value and the support pressure Pvent is an output value is shown below.
G (s) ₂₀ = Gc (s) · e ^{−Lc · s} (6)

  Here, Gc (s) represents a transfer function of the gas supply mechanism 20 excluding the dead time element, and means a first-order lag element in the present embodiment. Therefore, Gc (s) indicates 1 / (Tc · s + 1), and Tc is a time constant of the first-order lag element. E indicates the base of the natural logarithm, and Lc indicates a dead time required for the gas supply mechanism 20 to start adjusting the support pressure Pvent after the target pressure Pin is given. The other expressions have the same meaning with respect to the symbols shown in the above expressions.

The adjustment transfer function G (s) ₅₅ of the delay compensation unit 55 when the target pressure Pin is an input value and the delay compensation pressure ^ Pm is an output value is shown below.
G (s) ₅₅ = Gm (s) · e ^{−Lm · s} (7)

  Here, Gc (s) represents a transfer function excluding the time delay element, and in the present embodiment, means a first-order lag element. Therefore, Gm (s) indicates 1 / (Tm · s + 1), and Tm is a time constant of the first-order lag element. Lm represents a dead time in the adjustment transfer function.

  Further, in the actual patient's respiratory tract, the spontaneous breathing pressure Pmus is given in addition to the support pressure Pvent, which is different from the respiratory organ model of the observer 54. In the embodiment of the present invention, since the pressure loss in the intake pipe is small, the support pressure Pvent that is the discharge pressure of the gas supply mechanism 20 and the actual airway pressure Paw of the patient are approximated and handled.

  FIG. 3 is a graph showing the relationship between the ventilation volume Vmus during spontaneous breathing and the ideal ventilation volume Vast during assist breathing in the entire system 14 of the present invention. When it is considered that the transfer functions of the gas supply mechanism 20 and the delay compensation unit 55 are equal to each other, the support pressure Pvent is an amplification factor of (1 + B) times the spontaneous breathing pressure Pmus according to the time change of the spontaneous breathing pressure Pmus. Amplified. Ventilation is equal to the volume of support gas flowing into the lungs. Here, B indicates the amplification gain β when ^ R = R and ^ E = E as described above. The ventilation volume Vast during assist breathing by the ventilator 17 of the present embodiment is amplified to (1 + B) times the ventilation volume Vmus during spontaneous breathing.

  When the patient's condition is switched from the expiration period to the inspiration period, the patient operates respiratory muscles such as the diaphragm. As a result, the ventilation volume Vmus and the spontaneous breathing pressure Pmus during spontaneous breathing gradually increase with time, and gradually decrease when reaching a certain peak value P1. Then, the patient's condition is switched from the inspiration period to the expiration period.

  Normally, the patient's ventilation volume Vmus and spontaneous breathing pressure Pmus during spontaneous breathing during the inspiratory period first draws a gently increasing curve as a waveform with respect to time, and then draws a rapid decreasing curve when the peak value is reached from the local maximum. . However, depending on the patient's condition, the patient's ventilation volume Vmus and spontaneous breathing pressure Pmus vary greatly, and the peak value P1 and the inspiration period W1 vary.

  The gas supply mechanism 20 controlled by the control device 21 discharges the support gas at the support pressure Pvent so that the airway pressure Paw is proportionally amplified based on the amplification gain predetermined for the spontaneous breathing pressure Pmus of the patient. For example, when the peak value P1 of the spontaneous breathing pressure Pmus is small and the inhalation period W1 is short, the support pressure Pvent is set so that the peak value P2 of the airway pressure Paw is small and the period W2 in which the support gas is supplied becomes short. Be controlled. Similarly, when the peak value P1 of the spontaneous breathing pressure Pmus is large and the inhalation period W1 is long, the support pressure Pvent is set such that the peak value P2 of the airway pressure Paw is large and the period W2 during which the support gas is supplied becomes long. Is controlled.

  As described above, according to the control device 21 of the present embodiment, the support pressure Pvent is determined based on the flow rate deviation ΔF. The detected flow rate F varies depending on the patient's spontaneous breathing pressure Pmus, but the estimated flow rate ^ F is not affected by the patient's spontaneous breathing pressure Pmus. Accordingly, the flow rate deviation ΔF is a value obtained by extracting a change in the spontaneous breathing pressure Pmus. As a result, the spontaneous breathing pressure Pmus, which is normally difficult to detect, can be estimated, and can be configured as a disturbance observer when the spontaneous breathing pressure Pmus is regarded as a disturbance. Thus, by calculating the target pressure Pin according to the flow rate deviation ΔF related to the spontaneous breathing pressure Pmus, the assisting gas can be supplied to the patient at the assisting pressure Pvent that follows the spontaneous breathing pressure Pmus almost in real time. .

  Hereinafter, a method for controlling the gas supply mechanism 20 according to the present embodiment, in which a signal indicating the delay compensation pressure ^ Pm calculated using the delay compensation unit 55 is input to the observer 54 is referred to as an improved estimation type PAV method. . On the other hand, the control method of the gas supply mechanism 20 of the comparative example in which the support pressure Pvent is detected by the pressure detection means without using the delay compensation unit 55 and the detection result is input to the observer 54 is an estimated PAV method. Called. A control method of the gas supply mechanism 20 using the transfer function shown in FIG. 20 is referred to as a conventional PAV method.

FIG. 4 is a block diagram showing equivalent conversion of the entire system 14 when the improved estimation type PAV method is used. FIG. 4 shows a transfer function having the spontaneous breathing pressure Pmus as an input and the support pressure Pvent as an output. As shown in FIG. 4, in the improved estimation type PAV method, [{Gm (s) · e ^{−Lm · s} } / {Gc (s) · e ^{−Lc · s} } × {R · s + E} / {^ R When the feedback gain of s + ^ E} -1] is positive, a negative feedback configuration is used, and when the feedback gain is negative, a positive feedback configuration is used.

Therefore, even when {R · s + E} / {^ R · s + ^ E} is less than 1, {Gm (s) · e ^{−Lm · s} } / {Gc (s) · e ^{−Lc · s} } × { The negative feedback configuration can be maintained by adjusting the adjustment transfer function Gm (s) · e ^{−Lm · s} so that R · s + E} / {＾ R · s + ^{帰 還} E} exceeds 1. For example, in this embodiment, even if airway resistance R and lung elastance E change significantly due to a change in the patient's condition, and {R · s + E} / {^ R · s + ^ E} is less than 1, positive feedback By preventing the configuration, the area of the negative feedback configuration can be expanded.

Further, the entire system 14 using the improved estimation type PAV method takes the spontaneous breathing pressure Pmus as an input value, and an addition value of the spontaneous breathing pressure Pmus and the support pressure Pvent as an output value, the transfer function G (s) ₁₄ Is represented by the following equation.

ここで、各記号については、上述する記号にそれぞれ対応する。

Here, each symbol corresponds to the symbol described above.

If Gc (s) = 1, ^ R = R, ^ E = E, _KFG = ^ R · B, and _KVG = ^ E · B, the entire system 14 using the improved estimation type PAV method is spontaneous. The pressure (Pmus + Pvent) obtained by adding the respiration pressure Pmus and the support pressure Pvent is amplified to (1 + C) times the spontaneous respiration pressure Pmus. Here, C is represented by the following equation.

If the estimated airway resistance ^ R and the estimated elastance ^ E can be accurately estimated, that is, if (R · s + E) / (^ R · s + E) is 1, {Gm (s) · e ^{−Lm · As long as s} } / {Gc (s) · e− ^{Lc · s} } is greater than 1, the assist pressure Pvent can always be negatively fed back with respect to the spontaneous breathing pressure Pmus. Also, {Gm (s) · e− ^{Lm · s} } / {Gc (s) · e− ^{Lc · s} } is determined from the amplification gain B inputted in advance by being as close to 1 as possible. An amplification factor substantially equal to the amplification factor (1 + B) can be obtained.

Further, even when the estimated airway resistance ^ R and the estimated elastance ^ E cannot be accurately estimated and (R · s + E) / (^ R · s + E) is less than 1, as described above, Gm (s) · e ^{−Lm · s} } / {Gc (s) · e ^{−Lc · s} } × {R · s + E} / {^ R · s + ^ E} is greater than 1. By adjusting the time constant Tm and the dead time Lm, the support pressure Pvent can be stably controlled while maintaining the negative feedback configuration.

FIG. 5 is a block diagram showing equivalent conversion of the entire system 14 when the estimation type PAV method is used. In the estimated PAV method, a signal indicating the assist pressure Pvent detected by the pressure detection means is input to the observer 54. In this case, in FIG. 2, it is equivalent to the case where the transfer function of the gas supply mechanism 20 is equal to the adjustment transfer function (Gm (s) · e ^{−Lm · s} = Gc (s) · e ^{−Lc · s} ). It can be approximated as being.

  FIG. 5 shows a transfer function with the spontaneous breathing pressure Pmus as an input and the support pressure Pvent as an output. As shown in FIG. 5, in the estimation type PAV method, when the feedback gain of [{R · s + E} / {^ R · s + ^ E} −1] is positive, a negative feedback configuration is established, and the feedback gain is If it is negative, a positive feedback configuration is used. For example, if the airway resistance R and lung elastance E change significantly due to changes in the patient's condition, and {R · s + E} / {^ R · s + ^ E} is less than 1, a positive feedback configuration results. The system 14 cannot be configured as a negative feedback configuration.

The entire system 14 using the estimation type PAV method takes the spontaneous breathing pressure Pmus as an input value, and if the sum of the spontaneous breathing pressure Pmus and the support pressure Pvent is an output value, the transfer function G (s) ₁₄ is expressed as follows: Shown in

As described above, when the estimated PAV method is used and {circumflex over (R)} R and {circumflex over (E)} E> {{R · s + E} / {^ R · s + ^ E} is less than 1, positive It becomes a feedback configuration and stability is lowered. On the other hand, when the improved estimation type PAV method according to this embodiment is used, even if {R · s + E} / {^ R · s + ^ E} is less than 1, adjustment is performed. {Gm (s) · e− ^{Lm · s} } / {Gc (s) · e− ^{Lc · s} } × {R · s + E} / {^ R · s + ^ E} By making the value exceed 1, the negative feedback configuration can be maintained and the stability can be improved.

  It is impossible for the estimated airway resistance ^ R and the estimated elastance ^ E to be exactly the same as the actual airway resistance R and the actual elastance E. When there is a deviation (^ R ≠ R, ^ E ≠ E), it is normal.

  In the entire system 14 using the improved estimation type PAV method, the time constant Tm and the dead time Lm, which are parameters of the adjustment transfer function, are appropriately selected, and compared with the overall system 5 using the estimation type PAV method. Thus, the area that becomes the negative feedback configuration can be further increased and the positive feedback configuration can be made difficult. Therefore, even if the estimated values ^ R and ^ E are not accurate, the margin of stability as the entire system 14 Is expensive.

  As described above, the improved estimation type PAV method can increase the margin of stability compared to the estimation type PAV method, so that even if the patient's state changes, setting errors of the estimated values ^ R and ^ E. Even if there is an error, even if the gain is set large or a disturbance or the like acts, the runaway can be made more difficult to occur.

  In this embodiment, the adjustment transfer function has a first-order lag element and a dead time element, which are control elements approximating the transfer function of the gas supply mechanism 20. Therefore, the delay compensation pressure ^ Pm can be set according to the time change of the support pressure Pvent. As a result, the value obtained by dividing the adjustment transfer function by the transfer function of the gas supply mechanism can be made close to 1, and the amplification factor of the support pressure Pvent with respect to the support pressure Pmus is set as an amplification gain by the control amount calculation means. And the difference between the actual amplification factor and the actual amplification factor can be reduced.

FIG. 6 is a block diagram showing equivalent conversion of the entire system 5 when the conventional PAV method is used. FIG. 6 shows a transfer function with the spontaneous breathing pressure Pmus as an input and the support pressure Pvent as an output. As shown in FIG. 6, in the conventional type PAV method, always a positive feedback configuration, the flow gain K _fa and volume gain K _Va is not appropriate, there is a possibility that the support pressure Pvent diverges.

When the entire system 5 using the conventional PAV method has the spontaneous breathing pressure Pmus as an input value and the sum of the spontaneous breathing pressure Pmus and the support pressure Pvent as an output value, the transfer function G (s) ₅ is given by It is expressed by the following formula.

  In the entire system 5 using the conventional PAV method, the pressure (Pmus + Pvent) obtained by adding the spontaneous breathing pressure Pmus and the support pressure Pvent is amplified to 1 / (1-A) times the spontaneous breathing pressure Pmus. In this case, if A <0 or A> 1, the spontaneous breathing pressure Pmus cannot be amplified.

  On the other hand, in the improved estimation type PAV method, the parameter of the adjustment transfer function is appropriately set even if the estimation of the respiratory organ model is not a little accurate, and the adjustment transfer is made to the transfer function of the gas supply mechanism 20. By changing the function differently, the amplification can be performed as long as the amplification gain B is larger than zero. Therefore, the entire system 14 using the improved estimation type PAV method can increase the degree of freedom in gain selection.

  FIG. 7 is a graph showing experimental results using the improved estimation type PAV method according to the present embodiment. FIG. 8 is a graph showing experimental results using the estimated PAV method as a comparative example. FIG. 9 is a graph showing experimental results using the conventional PAV method. In the experiments shown in FIG. 7 to FIG. 9, spontaneous measurement was performed when a respirator was connected to a simulation apparatus that simulated a patient's respiratory organ, and the simulation apparatus performed a respiratory simulation operation that simulated the patient's respiratory action. The measurement result of respiratory pressure Pmus and support pressure Pvent is shown. In FIG. 7 to FIG. 9, the spontaneous breathing pressure Pmus is indicated by a solid line, the support pressure Pvent is indicated by a broken line, and the simulator is operated under the same conditions as the patient's inspiratory period and the temporal change of the spontaneous breathing pressure.

  Table 1 shows various parameters used in the experiment. As shown in Table 1, the estimated airway resistance ^ R and the estimated elastance ^ E are set to be the same as the actual airway resistance R and the actual elastance E, and the time constant Tm of the adjustment transfer function is The experiment is performed with the time constant Lc of the transfer function of the gas supply mechanism 20 being larger than the time constant Tc of the transfer function (Tm> Tc) and the dead time Lm of the transfer function of the gas supply mechanism 20 being the same as the dead time Lc of the transfer function. Went.

  As shown in FIGS. 7 to 9, when the improved estimation type PAV method is used, the quick response at the intake start timing can be improved as compared with the case where the conventional type PAV method and the estimation type PAV method are used. The support pressure Pvent can be made to suitably follow the change in the spontaneous breathing pressure Pmus. Specifically, when an arbitrary time W2 at the time of rising has elapsed since the start of the increase in the spontaneous breathing pressure ^ Pm, the increase amount X11 of the support pressure Pvent of the improved estimated PAV method is expressed as the estimated PAV method and The increase amount X12, 13 of the support pressure Pvent of the conventional PAV method can be made larger.

When R = ^ R, E = ^ E, and the high frequency response portion in the frequency domain, that is, s is very large, when the equation (8) is modified, the transfer function G (s) ₁₄ of the entire system is expressed as Tm It can be approximated with s.

  In this case, the time constant Tm acts as a differential operation of a proportional integral differential (PID) feedback operation, and the differential gain is proportional to Tm / Tc. Therefore, by increasing the time constant Tm and satisfying Tm> Tc, the differential gain can be increased, and as a result, the rapid response at the time of rising of the spontaneous breathing pressure Pmus can be improved. In the low frequency response portion, the differential gain is reduced by the influence of parameters other than the time constant Tm. As a result, the differential gain at the end of the intake period can be reduced, and overshoot can be suppressed.

  In addition, due to the improvement in speed response, the asynchronous periods W31 and W32 when the estimated PAV method and the improved estimated PAV method are used are made smaller than the asynchronous period W33 when the conventional PAV method is used. be able to. Here, the asynchronous period is the difference between the time when the spontaneous breathing pressure Pmus begins to decrease and the time when the support pressure Pvent begins to decrease. Further, in the conventional PAV method and the estimated PAV method, the support pressure Pvent suddenly increases from the time when the spontaneous breathing pressure Pmus starts to decrease (X22, X23). On the other hand, in the improved estimation type PAV method, the increase in the support pressure Pvent is small from the time when the spontaneous breathing pressure Pmus starts to decrease (X21).

  Thus, in the improved estimation type PAV method, the rapid response at the time of rising of the spontaneous breathing pressure Pmus can be improved, the asynchronous period W3 can be shortened, and the increase in the support pressure Pvent after inspiration is reduced, The burden that the ventilator gives to the patient can be reduced.

  Further, by setting the dead time Lm included in the transfer function for adjustment and the dead time Lc included in the transfer function of the gas supply mechanism to substantially the same value (Lm≈Lc or Lm = Lc), the first order of the transfer function for adjustment When the time constant of the delay element is changed, the stable region can be increased. In other words, the range of time constants that can be changed can be expanded while maintaining a stable state. As a result, the time constant can be increased as much as possible within a stable range where the response of the support pressure Pvent does not vibrate or diverge, and the speed response of the support pressure Pvent can be further improved.

  Furthermore, in the estimated PAV method, the support pressure Pvent increases while vibrating, whereas in the improved estimated PAV method, the vibration of the support pressure Pvent can be suppressed. This further reduces the load that the ventilator places on the patient. In the improved estimated PAV method, the pressure detecting means can be made unnecessary, the manufacturing cost can be reduced, the failure due to the failure of the pressure detecting means can be eliminated, and the reliability of the ventilator Can be improved.

  FIG. 10 shows experimental results comparing the support pressure Pvent response waveforms of the improved estimated PAV method and the conventional PAV method. FIG. 11 shows experimental results comparing the assist gas flow rate F response waveforms of the improved estimated PAV method and the conventional PAV method. 10 and 11 show experimental results in which the dead time Lm of the adjustment transfer function used in the improved estimation type PAV method is 10 msec and the time constant Tm is 5, 10, and 20. As the time constant Tc and the dead time Lc of the gas supply mechanism 20, the values set in Table 1 are used.

  As is apparent from FIG. 10, the improved estimation type PAV method can improve the speed response as compared with the conventional PAV method. Further, in the improved estimation type PAV method, the quick response is further improved in proportion to the increase of the time constant Tm. When the time constant Tm is excessive, the support pressure Pvent becomes oscillating.

As is apparent from FIG. 11, in the improved estimation type PAV method, the time constant Tm is increased and the flow rate F of the support gas becomes oscillatory. When the flow rate F of the support gas becomes oscillating, it means that the load applied to the patient increases, which is not preferable. When the response of the support gas flow rate F is oscillating, it is possible to prevent the support gas flow rate F from oscillating by decreasing the flow rate gain _KFG . As a result, the amplification factor can be increased without causing the flow rate of the support gas to vibrate.

Further, the entire system 14 according to the present embodiment can be made less likely to be a positive feedback configuration and has improved stability, so that the degree of freedom in parameter selection is great, the amplification gain β is amplified, and the flow rate gain K _{Even if the FG} and the volume gain _KVG are changed, the runaway is unlikely to occur and can be suitably adjusted.

  FIG. 12 is a simulation result showing the response of the support pressure Pvent when the amplification factor for the spontaneous breathing pressure Pmus and the time constant Tm of the adjustment transfer function are changed in the improved estimation type PAV method. A respiratory model and a model simulating the respiratory organ model are configured by a program, and the gain and the time constant Tm of the adjustment transfer function are changed.

  The simulation result shown in FIG. 12 is the same as the experimental result shown in FIG. Specifically, in the improved estimation type PAV method, the time constant Tm can be increased and the quick response can be improved. When the time constant Tm of the adjustment transfer function is excessive, the support pressure Pvent becomes oscillatory. .

  In addition, when the amplification factor is large, the time constant Tm of the adjustment transfer function with which the support pressure Pvent is oscillating is smaller than when the amplification factor is small. For example, in the simulation result of this embodiment, when the amplification factor is four times, the response of the support pressure Pvent becomes oscillating when the time constant Tm is 33 msec, and when the amplification factor is six, the adjustment transmission is performed. When the time constant Tm of the function is 25 msec, the response of the support pressure Pvent becomes oscillating.

When the time constant Tm of the adjustment transfer function in which the support pressure Pvent is oscillating is defined as the oscillating time constant Tm1, the oscillating time constant Tm1 generally has the following relationship.
Amplification gain β × (Tm1 / Tc)> D (12)
Here, D is a predetermined constant, and is 6 in the present embodiment. When the time constant Tm of the adjustment transfer function that does not make the support pressure Pvent vibrate is the non-vibration time constant Tm2, the non-vibration time constant Tm2 generally has the following relationship.
Tm2 <D × Tc / amplification gain β (13)

  In the present embodiment, the time constant Tm of the adjustment transfer function is determined so as to satisfy the expression (13). In other words, the upper limit of the time constant Tm of the first-order lag element of the adjustment transfer function is set according to the amplification factor for proportionally amplifying the spontaneous breathing pressure Pmus. By setting the time constant Tm of the upper limit value that satisfies the expression (13), the response of the support pressure Pvent does not vibrate, and the quick response can be improved as much as possible.

  For example, the time constant Tm of the adjustment transfer function is set to 1/2 of the vibration time constant Tm1. In this embodiment, the time constant Tm of the adjustment transfer function is set to 1.5 to 2 times the time constant Tc of the gas supply mechanism 20.

  When pressure detecting means for detecting the support pressure Pvent is provided, the estimation means 51 is given a signal indicating the support pressure Pvent from the pressure detection means, and determines that the support pressure Pvent is oscillatory based on the signal. Then, the time constant Tm of the adjustment transfer function may be adjusted to be lowered, and if it is determined that it is not vibrational, the time constant Tm of the adjustment transfer function may be adjusted to be increased. This can more reliably prevent the support pressure Pvent from becoming vibrational.

  The time constant Tm of the adjustment transfer function may be changed at the start of intake and the end of intake. For example, by increasing the time constant Tm at the start of intake and decreasing the time constant Tm at the end of intake, it is possible to improve the quick response at the start of intake and reduce the overshoot amount at the end of intake.

  As described above, when the improved estimation type PAV method according to the embodiment of the present invention is used, in the entire system 14, the estimated airway resistance ^ R and the estimated elastance ^ for determining the state of the patient's respiratory period In addition to E and the amplification factor of the support pressure Pvent, parameters for adjusting the control characteristics of the entire system are set to be adjustable.

As numerical values input to the control device, the estimated airway resistance ^ R and the estimated elastance ^ E are values determined according to the patient's condition, and the flow rate gain K _VG and the volume gain K _VG are This is a constant determined medically according to the pathological condition. In the present embodiment, the control characteristics of the entire system 14 can be improved by further changing the parameters of the adjustment transfer function.

  Specifically, the time constant Tm of the primary transfer element of the adjustment transfer function and the dead time Lm of the dead time element can be set. In this case, by appropriately setting the time constant Tm and the dead time Lm and increasing the area of the negative feedback configuration, the stability margin can be improved and the instability can be made difficult. Therefore, even if the conventional PAV method and the estimated PAV method have a positive feedback configuration, the improved estimated PAV method of the present embodiment can stabilize the control system as a negative feedback configuration. . This can prevent runaway. As a result, a control method for the gas supply mechanism that further reduces the burden on the patient can be realized.

  In addition, by appropriately selecting the parameters of the adjustment transfer function, the responsiveness of the support pressure Pvent can be improved, and the support pressure Pvent can be amplified in proportion to the spontaneous breathing pressure Pmus with high accuracy. Asynchronous state between the ventilator and the patient can be suppressed. In addition, the support pressure Pvent can be prevented from becoming vibrational. Thus, the rapid response and robust stability of the support pressure Pvent can be improved, and the load on the patient can be reduced.

For example, even if the estimated airway resistance ^ R and the estimated elastance ^ E are slightly deviated from the actual airway resistance R and elastance E, it is possible to increase the stability margin of the entire system as described above. Therefore, the entire system 14 is prevented from becoming unstable, and the amplification gains K _FG and K _VG for amplifying the support pressure Pvent can be increased. In particular, as described above, even when {circumflex over (R)} R and {circumflex over (E)} E, it is possible to adopt a negative feedback configuration and set a larger stability margin. Therefore, even if the airway resistance R and the elastance E are not accurately obtained, the possibility of a runaway is reduced, and the gas supply mechanism 20 can be controlled. In the present embodiment, the parameters of the adjustment transfer function are set to be adjustable. As a result, even when the control characteristics vary for each gas supply mechanism 20 or when the control characteristics change due to changes over time, the control characteristics can be stabilized by appropriately adjusting the parameters.

  FIG. 13 is a block diagram illustrating an example of the ventilator 17. The control device 21 includes a control device main body 33 including a computer, a flow rate detection means 50, an input means 39, a display means 40, and servo amplifiers 47 and 48 which are amplifier circuits. The control device 21 may further include airway pressure detection means 61.

  The flow rate detection means 50 converts the flow rate of the gas flowing through the intake pipe 25 of the gas supply mechanism 20 into an electrical signal, and gives the electrical signal to the control device main body 33. The input means 39 includes an estimated airway resistance R ^ and estimated elastance ^ E, an amplification gain β, a time constant Tm of the delay compensation unit 55, and a dead time Lm from a doctor and a nurse or a manager who manages the gas supply mechanism 20. Etc. are entered. The input means 39 gives a signal indicating the input information to the control device body 33.

  The display means 40 is an informing means for informing the patient's airway pressure. Based on the display command signal received from the control device main body 33, the display means 40 displays a waveform indicating a temporal change in the spontaneous breathing pressure Pmus of the patient on the display screen.

  The amplifier circuit 48 gives a signal indicating the target pressure Pin calculated by the control device main body 33 to the pump actuator 31. The pump actuator 31 controls the pump based on a signal indicating the target pressure Pin, and the discharge pressure of the gas supply mechanism 20 is feedback-controlled.

The control device main body 33 includes an interface 101, a calculation unit 102, a primary storage unit 103, and a storage unit 104. The interface 101 receives a signal from the connected flow rate detection means 50 and gives the signal to the calculation unit 102. The storage unit 104 stores a program to be executed by the control device main body 33, and the calculation unit 102 reads and executes the program stored in the storage unit 104, whereby the estimation unit 51, the deviation calculation unit 52, the control amount The calculation means 53 can be realized. As a result, the control device main body 33 can control the gas supply mechanism 20 described above. The storage unit 104 may be a computer-readable recording medium such as a compact disc. The arithmetic unit 102 is realized by an arithmetic processing circuit such as a CPU, and executes an operation according to an operation program stored in the storage unit 104.

FIG. 14 is a flowchart showing the operation of the control device main body 33. First, in step s0, the control device main body 33 has parameters such as the estimated airway resistance ^ R, the estimated elastance ^ E, the transfer function of the gas supply mechanism 20, the flow gain _KFG, and the volume gain _KVG. When the input and the target pressure Pin and the estimated flow rate {circumflex over (F)} are ready for calculation, the process proceeds to step s1.

  In step s1, the control device main body 33 executes the operation of the deviation calculating means 52, and the deviation ΔF between the estimated flow rate ^ F obtained from the previously calculated target pressure Pin and the flow rate F given from the flow rate detecting means 50. Is calculated. When the flow rate deviation ΔF is calculated, the process proceeds to step s2.

  In step s2, the control device main body 33 executes the operation of the control amount calculation means 53, and calculates the target pressure Pin based on the flow rate deviation ΔF. When the target pressure Pin is calculated, a signal representing the target pressure Pin is given to the gas supply mechanism 20, and the process proceeds to step s3.

  In step s3, the control device main body 33 executes the operation of the estimating means 51, performs the operation corresponding to the delay compensation unit 55 and the observer 54 based on the signal representing the target pressure Pin, and the spontaneous breathing pressure Pmus exists. If not, the estimated flow rate {circumflex over (F)} that will be supplied to the patient is calculated, and the process proceeds to step s4.

  In step s4, the control device main body 33 determines whether or not a predetermined end condition is satisfied. For example, when the end command is not given by the input means 39, it is determined to continue the control of the gas supply mechanism 20, and the process returns to step s1. In step s1, the flow rate deviation ΔF is calculated again using the estimated flow rate F calculated in step s3 and the flow rate F given from the flow rate detection means 50. In step s4, when determining that the predetermined end condition is satisfied, the control device main body 33 proceeds to step s5 and ends the control operation.

  As described above, the delay compensation unit, the flow rate estimation unit, the deviation calculation unit, and the control amount calculation unit described above may be realized by executing software predetermined by a computer. Further, the gas supply mechanism 20 is not particularly limited as long as the pressure of the assisting gas to be discharged can be controlled by the control device 21 and an inhalation conduit 25 for guiding the assisting gas to the patient's airway 15 is formed. For example, the gas supply mechanism 20 may be a ventilator having a bellows type pump as shown in FIG. 13, but may be a ventilator that supplies support gas via a pipe.

  The above-described embodiment of the present invention is an example of the present invention, and the configuration can be changed within the scope of the invention. For example, the above-described block diagram is merely an example of the present invention, and may be equivalently converted if the same effect can be obtained. The time constant Tm and the dead time Lm of the adjustment transfer function described above are merely examples, and are not limited to these values. For example, the dead time Lm of the transfer function for adjustment and the dead time Lc of the transfer function approximating the gas supply mechanism 20 may be different values. Further, the time constant Tm of the adjustment transfer function may be smaller than the time constant Tc of the transfer function approximating the gas supply mechanism 20.

  In the present embodiment, the parameters in the adjustment transfer function can be changed, but the case where the parameters of the adjustment transfer function are fixed is also included in the present invention. In this case, the adjustment transfer function is determined so that the entire system has a negative feedback configuration. As a result, the same effects as those of the above-described embodiment can be obtained. In this case, the adjustment transfer function includes a control element approximating the transfer function of the gas supply mechanism, so that the value obtained by dividing the adjustment transfer function by the transfer function of the gas supply mechanism can be close to 1, and the support pressure Regarding the amplification factor of the support pressure Pvent with respect to Pmus, the difference between the amplification factor set by the control amount calculation means and the actual amplification factor can be reduced.

  Further, the time constant Tm of the first-order lag element constituting the adjustment transfer function is set to be larger than the time constant Tc approximating the first-order lag element included in the gas supply mechanism (Tm> Tc). As a result, the quick response can be improved and the asynchronous state can be suppressed. In addition, the support pressure Pvent can be accurately amplified in proportion to the spontaneous breathing pressure Pmus that changes with the passage of time. Furthermore, by setting the dead time Lm included in the adjustment transfer function and the dead time Lc included in the transfer function of the gas supply mechanism to substantially the same value (Lm≈Lc or Lm = Lc), the first order of the adjustment transfer function When the time constant of the delay element is changed, the stable region can be increased. In other words, it can maintain a stable state and increase the variable time constant. As a result, the time constant can be increased as much as possible within a stable range where the response of the support pressure Pvent does not vibrate or diverge, and the speed response of the support pressure Pvent can be further improved.

  Further, the upper limit value of the time constant Tm of the transfer function for adjustment is determined based on the equation (13) so that the assist pressure Pvent does not diverge according to the amplification factor of the assist pressure Pvent determined by the ventilator manager. It is preferable. As a result, even if the amplification factor is changed, the support pressure Pvent can be prevented from becoming oscillating, and the quick response can be improved as much as possible.

In the present embodiment, the adjustment transfer function includes the first-order lag element Gm (s) and the dead time element e ^{−Lm · s} . However, the present invention is not limited to this. For example, the adjustment transfer function may include a multi-order delay element obtained by approximating the transfer function of the gas supply mechanism with a rational function. In this case, the multi-order delay element Gm (s) of the adjustment transfer function is expressed by the following equation.

  The adjustment transfer function may include a proportional element k in the equation (14). In this case, the adjustment transfer function is expressed by the following equation. By setting the proportional element, the amplification factor can also be set by the adjustment transfer function.

  The proportional element k may be 0.8. Further, the adjustment transfer function may include a dead time element in the above-described equation (14) or (15).

  FIG. 15 is a graph showing the concept of changes in the evaluation value F of robust stability and rapid response when the time constant Tm of the first-order lag element and the dead time Lm of the dead time element are changed. As the evaluation value F, for the control of the support pressure Pvent, a value obtained by evaluating the two controllability of the robust stability and the quick response which are in a trade-off relationship is adopted. Here, FIG. 15 is intended to facilitate understanding of changes in the evaluation value F of robust stability and rapid response when the time constant Tm of the first-order lag element and the time delay Lm of the time delay element are changed. It is used for the above, and does not coincide with the actual change state.

  As shown in FIG. 15, when the time constant Tm, dead time Lm, and evaluation value F are respectively set on the three-dimensional coordinate axes, the entire system using the improved PAV method of the present embodiment shown in the equation (2) is used. In this case, it is estimated that there will be two points where the evaluation value F becomes the maximum value together with the experimental result using the simulation device simulating the respiratory organ of the patient, although there is a deviation from the calculation result by the simulation. .

  Therefore, by setting the combination of the time constant Tm constituting the first maximum point M1 or the second maximum point M2 and the dead time Lm as a parameter of the adjustment transfer function, the robust stability and the quick response are balanced, and Each can be relatively good.

  In the present embodiment, in the vicinity of the first maximum point M1, the value (Tm + Lm) obtained by adding the time constant Tm and the dead time Lm in the adjustment transfer function is the first-order lag element included in the transfer function of the gas supply mechanism. It is set smaller than the value (Tc + Lc) obtained by adding the approximate time constant Tc and the dead time Lc approximating the dead time element.

Further, a dead time Lc and a time constant Tc approximating a dead time element included in the transfer function of the gas supply mechanism are set to Lc = 10 msec and Tc = 24 msec, respectively. When the estimated airway resistance R is 20 (cmH ₂ 0) / (liter / second), the estimated elastance is 1/30 (milliliter) / (cmH ₂ O), and the amplification factor is 4 times, the improved PAV method is In the experimental results of the present embodiment used, the dead time Lm of the adjustment transfer function was 4 msec, and the time constant Tm was 10 msec.

Similarly, in the vicinity of the second maximum point M2, the value (Tm + Lm) obtained by adding the time constant Tm and the dead time Lm in the adjustment transfer function approximates the first-order lag element included in the transfer function of the gas supply mechanism. It is set smaller than a value (Tc + Lc) obtained by adding the constant Tc and the dead time Lc approximating the dead time element. In the vicinity of the second maximum point M2, the dead time Lm included in the adjustment transfer function is set to zero as described above. As described above, the dead time Lc and the time constant Tc approximating the dead time element included in the transfer function of the gas supply mechanism are set to Lc = 10 msec and Tc = 24 msec, respectively. When the estimated airway resistance R is 20 (cmH ₂ 0) / (liter / second), the estimated elastance is 1/30 (milliliter) / (cmH ₂ O), and the amplification factor is 4 times, the improved PAV method is In the present embodiment used, the dead time Lm of the adjustment transfer function is 0 msec and the time constant Tm is 10 msec. As the time constant Tm constituting the second maximum point M2, a value larger than the time constant Tm constituting the first maximum point M1 is adopted as much as the dead time Lm is made zero.

  Here, at the first maximum point M1, the combination of the dead time Lm and the time constant Tm varies as the parameters of the entire system including the patient and the ventilator change. On the other hand, even if the parameter shown in Table 1 changes, the value of the time constant Tm varies at the second maximum point M2 while the dead time Lm is kept at zero.

  Thus, the time constant of the first-order lag element of the adjustment transfer function is changed by setting the parameters of the adjustment transfer function to the time constant Tm and the dead time Lm constituting the first maximum point M1 or the second maximum point M2. In this case, the stable area can be increased. In other words, it can maintain a stable state and increase the variable time constant. As a result, the time constant can be increased as much as possible within a stable range where the response of the support pressure Pvent does not vibrate or diverge, and the speed response of the support pressure Pvent can be further improved. Further, if the respiratory organ model is not accurate with respect to the actual respiratory organ, even if the amplification factor is large, the dead time Lm included in the adjustment transfer function and the dead time Lc included in the transfer function of the gas supply mechanism. As compared with a case where the difference is excessively different, the time constant can be increased as much as possible in the stable range, and the quick response of the support pressure Pvent can be further improved.

  Further, with the time constant Tm and the dead time Lm constituting the second maximum point M2, the dead time Lm of the dead time element is set to zero regardless of changes in the parameters of the entire system including the patient and the ventilator. The parameters necessary for the transfer function for adjustment can be reduced, and an appropriate time constant Tm can be easily obtained by manual or computer calculation, as compared with the case where both the dead time Lm and the time constant Tm are adjusted. be able to.

  In this embodiment, the evaluation value F is defined as a linear sum (F = F1 + F2) of the robust stability evaluation value F1 related to robust stability and the rapid response evaluation value F2 related to rapid response. In the present embodiment, the robust stability evaluation value F1 related to the robust stability is expressed by equation (16).

Here, F1 is a robust stability evaluation value. The A ₁ is a robust stability weight constant to determine the robust stability evaluation value F1. Further, another term | S | multiplied by the robust stability A ₁ is a value for evaluating the stability margin, and is set as a modulus margin in the present embodiment. The modulus margin is expressed by the equation (17) when the loop transfer function in the entire system including the ventilator and the patient is expressed by L (jω). That is, in the Nyquist diagram, the vector trajectory of the loop transfer function in the entire system and the stability limit are represented by the closest distance. In the present embodiment, the robust stability evaluation value F1 related to robust stability is determined based on the equations (16) and (17). However, it may be determined based on another evaluation equation.

  In the present embodiment, the quick response evaluation value F2 related to the quick response is expressed by the equation (18).

Here, F2 is a rapid response evaluation value. A ₂ is a rapid response weight constant for determining the rapid response evaluation value F2. Further, another term (1 / Tresp) multiplied by the quick response weight constant A ₂ is a value for evaluating the quick response, and in this embodiment, when the step response is considered, the support pressure The value obtained by dividing Pvent by the spontaneous breathing pressure Pmus (= Pvent / Pmus) is represented by the reciprocal of the time Tresp (1 / Tresp) until reaching 0.632. In the present embodiment, the quick response evaluation value F2 related to the quick response is determined based on such a relationship, but may be determined based on another evaluation formula.

As described above, since robust stability and rapid response are in a trade-off relationship, when robust stability is improved, rapid response is reduced. For example, when it is desired to obtain a comprehensive evaluation value F that weakens the robust stability and improves the quick response, the quick response weight constant A ₂ is made larger than the robust stability weight constant A ₁ described above. It is possible to obtain a well-balanced comprehensive evaluation value with low robust stability and high responsiveness. Thus, by appropriately allocating the balance between the weight constants A ₁ and A ₂ , it is possible to obtain an evaluation value in consideration of the balance between robust stability and rapid response. In the present embodiment, the determination is made based on the above-described equations (16) to (18). However, this determination is an example, and an evaluation value related to robust stability and an evaluation related to rapid response are determined according to other arithmetic expressions. It is good also as a comprehensive evaluation value using the value which united the value.

  When the improved PAV method is used by using the time constant Tm and the dead time Lm constituting the first maximum point M1 and the second maximum point M2 obtained in this way, it is more stable than the simple estimation type PAV. The stability and speed can be improved compared to the improved PAV method using a time constant Tm and time delay Lm selected at random, as well as improving the stability and speed. it can.

Table 2 is a table showing experimental comparison results when using the time constant Tm and the dead time Lm constituting the first maximum point M1 and the second maximum point M2, respectively. In Table 2, a dead time Lc and a time constant Tc approximating a dead time element included in the transfer function of the gas supply mechanism are set to Lc = 10 msec and Tc = 24 msec, respectively. In addition, the estimated airway resistance R is 20 (cmH ₂ 0) / (liter / second), the estimated elastance is 1/30 (milliliter) / (cmH ₂ O), the amplification factor is quadrupled, and the airway resistance is measured by a simulator. The experiment was conducted by changing R and lung elastance E. In this case, the two local maximum points M1 and M2 are values obtained by adding the time constant Tm and the dead time Lm in the adjustment transfer function (Tm + Lm) to approximate the first-order lag element included in the transfer function of the gas supply mechanism. It is set smaller than a value (Tc + Lc) obtained by adding the time constant Tc and the dead time Lm approximating the dead time element.

  As shown in Table 2, in the present embodiment, the first maximum point M1 is configured by using the time constant Tm and the dead time Lm constituting the second maximum point M2 (Lm = 0, Tm = 20). It can be seen that robust stability and quick response are improved as compared with the case where the time constant Tm and the dead time Lm are used (Lm = 4, Tm = 10). That is, in the present embodiment, robust stability and quick response can be further improved when the dead time Lm is zero.

  FIG. 16 is a graph showing the support pressure Paw and the ventilation flow rate Qi using the time constant Tm and the dead time Lm constituting the second maximum point M2. FIG. 16 compares the case where the improved PAV method is used with the case where the conventional PAV method is used.

In FIG. 16, a dead time Lc and a time constant Tc approximating a dead time element included in the transfer function of the gas supply mechanism are set to Lc = 10 msec and Tc = 24 msec, respectively. The 20 _{(cm H} 2 0) and the estimated airway resistance R and airway resistance R in the simulator / (l / sec), and 1/30 (mL) and elastance E and the estimated elastance of the simulator / (cm H ₂ O), the amplification factor was tripled, the time constant Tm of the adjustment transfer function was 20 msec, and the dead time Lm of the adjustment transfer function was 0 msec.

The first set time δT1, which is the time difference from the time T2 when the patient finishes the spontaneous inspiration to the time when the ventilation flow rate Qi shifts to the state before reaching the value Q0 of the time T1 before the start of the spontaneous inspiration, Assuming the degree of asynchrony, the first set time δT _new 1 in the improved estimated PAV method according to the present embodiment can be made smaller than the first set time δT _org 1 in the conventional PAV method. . As a result, the period of asynchronous breath can be shortened and the load on the patient can be reduced.

Further, the airway pressure Paw can be approximated to a waveform shape obtained by amplifying the patient's spontaneous breathing pressure Pmus by a predetermined amplification factor of three. Specifically, in the improved estimation type PAV method, the rapid response of the airway pressure Paw to the patient's spontaneous breathing pressure Pmus can be improved as compared with the conventional PAV method. Further, the second set time δT2, which is the time difference from the time T2 when the patient finishes the spontaneous inspiration to the time when the airway pressure Paw starts to decrease, is set as the degree of exhalation asynchronization, and the improved estimation type according to the present embodiment The second set time δT _new 2 in the PAV method can be made smaller than the second set time δT _org 2 in the conventional PAV method. This also makes it possible to shorten the period of exhaled breath and reduce the load on the patient.

  In the present embodiment, the time constant Tm and the dead time Lm of the adjustment transfer function that constitutes one of the first maximum point M1 and the second maximum point M2 are used in this way. The parameter is not limited to this.

  For example, as described above, the time constant Tm of the adjustment transfer function may be larger (Tm> Tc) than the time constant Tc that approximates the first-order lag element included in the transfer function of the gas supply mechanism. The dead time Lm of the adjustment transfer function may be substantially the same as the dead time Lc approximating the dead time element included in the transfer function of the gas supply mechanism (Lm≈Lc or Lm = Lc). In addition to the determination method described above, the time constant Tm and the dead time Lm of the adjustment transfer function may be determined as appropriate according to changes in the parameters of the entire system including the patient and the ventilator. Also, other forms of the adjustment transfer function can be used.

  FIG. 17 is a block diagram showing the entire system 13 according to still another embodiment of the present invention. The entire system 13 illustrated in FIG. 17 has the same configuration as the entire system 14 illustrated in FIG. 2 except that a part of the configuration of the estimation unit 51 is different. Therefore, the description of the same configuration is omitted, and the reference numerals corresponding to the entire system 14 in FIG.

  The estimation means 51 further includes a detection delay calculator 60. The detection delay calculator 60 has a detection means model that is modeled by simulating the flow rate detection means 50. In this case, when the estimated flow rate {circumflex over (F)} is given from the observer 54, the detection delay computing unit 60 computes the estimated flow rate {circumflex over (F)} based on the detection delay of the flow rate detection means 50, and gives this computation result to the deviation computation means 52. . The deviation calculator 52 subtracts the detected flow rate detected by the flow rate detector 50 from the estimated flow rate F calculated by the detection delay calculator 60 to calculate a flow rate deviation ΔF. As a result, the estimation means 51 is supplied to the patient's airway based on the time characteristic from when the flow rate F of the support gas supplied to the patient's airway is detected until the flow rate detection means 50 outputs the detection result. The flow rate {circumflex over (F)} of the assist gas to be calculated can be calculated more accurately. Therefore, the asynchronous state can be more reliably prevented.

  FIG. 18 is a block diagram showing the entire system 12 according to still another embodiment of the present invention. The entire system 12 shown in FIG. 18 has the same configuration as the entire system 14 shown in FIG. 2 except that a part of the configuration of the estimation means 51 is different. Therefore, the description of the same configuration is omitted, and the reference numerals corresponding to the entire system 14 in FIG.

Instead of the delay compensation unit 55, a pressure detection means 61 is provided. The pressure detector 61 detects an airway pressure Paw, which is a pressure in the patient's airway. Then, the pressure detection unit 61 gives the detected airway pressure Paw to the predetermined delay compensation unit 55. The delay compensation unit has an adjustment transfer function for improving the quick response of the support pressure Pvent. When the airway pressure Paw is given as an input, the delay compensation unit calculates the delay compensation pressure ^ Pm and gives the calculated delay compensation pressure ^ Pm to the observer 54. Here, the delay compensation pressure ^ Pm is set to {Gm (s) · e ^−Lms } / {Gc (s) · e ^−Lcs } using the above-described symbols. Even if it does in this way, the effect mentioned above can be achieved.

  FIG. 19 is a block diagram showing the entire system 10 according to still another embodiment of the present invention. The overall system 10 shown in FIG. 19 is the same as the overall system 14 shown in FIG. 2 except that the estimated airway resistance ^ R and estimated elastance ^ E set in the estimating means 51 are different. Have Therefore, the description of the same configuration is omitted, and the reference numerals corresponding to the entire system 14 in FIG.

  FIG. 20 is a graph for explaining the airway resistance R. When the flow of the assisting gas flowing in the airway becomes a laminar flow, the flow velocity changes linearly in proportion to the airway pressure Paw. However, in reality, the airway repeats branching and the thickness is not uniform, so the flow of the support gas becomes turbulent. Therefore, the estimated airway resistance ^ R is set in consideration of the turbulent resistance.

The estimated airway resistance {circumflex over (R)} set in the estimating means 51 includes a first resistance coefficient {circumflex over (R)} _T set constant regardless of the flow rate of the support gas, and a flow rate of the support gas calculated by the estimated flow rate calculator. This is a value obtained by adding the second resistance coefficient ^ K _T depending on F. The first resistance coefficient {circumflex over (R) _} and the second resistance coefficient {circumflex over (K)} _T are set to coefficients corresponding to the airway resistance of the patient. Further, the resistance of the entire respiratory organ including not only the lung but also the thorax may be set as the estimated airway resistance ^ R. The airway resistance R may be approximated by an estimated airway resistance ^ R expressed by another approximate expression.

  FIG. 21 is a graph for explaining lung compliance. The estimated elastance ^ E set in the estimating means 51 is a value based on the assist gas volume ^ V calculated by the assist gas volume calculator and is the reciprocal of the lung compliance C. The compliance C increases nonlinearly with the increase in the volume V of the assist gas during the patient's inspiration period, and has a saturation characteristic and a hysteresis characteristic.

  The alveolar pressure calculator 59 obtains in advance information indicating the relationship between the compliance C and the volume of the support gas, so that the alveolar pressure Palv can be accurately calculated even when the case where the compliance is nonlinear is taken into consideration. It can be calculated.

  Thus, since the observer has a non-linear respiratory tube model, the estimated flow rate {circumflex over (F)} can be estimated more accurately. As a result, the spontaneous breathing pressure Pmus can be accurately estimated, and the support pressure Pvent can be determined according to the estimated spontaneous breathing pressure Pmus.

  The estimated airway resistance R and estimated elastance E may be appropriately set by a doctor, but the airway resistance R and elastance E measured in advance by a measuring instrument may be used. Further, airway resistance ^ R and elastance ^ E obtained by the estimation method disclosed in Patent Document 2 may be used.

2 is a block diagram showing a ventilator 17 and a patient 18. FIG. It is a block diagram which shows concretely the whole system | strain 14 of one Embodiment of this invention. It is a graph which shows the relationship between the ventilation volume Vmus at the time of spontaneous respiration in the whole system 14 of this invention, and the ideal ventilation volume Vast at the time of assist breathing. It is a block diagram which shows the whole system 14 at the time of using the improved estimation type PAV method. It is a block diagram which shows the whole system | strain 14 at the time of using an estimation type PAV method. It is a block diagram which shows the whole system | strain 14 at the time of using the conventional PAV method. It is a graph which shows the experimental result using improved estimation type PAV method. It is a graph which shows the experimental result using a presumed PAV method. It is a graph which shows the experimental result using the conventional PAV method. It is the experimental result which compared the support pressure Pvent response waveform of the improved estimation type PAV method and the conventional type PAV method. It is the experimental result which compared the flow volume F response waveform of the improved estimation type PAV method and the conventional type PAV method. It is a simulation result which shows the response of the support pressure Pvent when the amplification factor with respect to the spontaneous-respiration pressure Pmus and the time constant Tm of the transfer function for adjustment are changed in the improved estimation type PAV method. 2 is a block diagram showing an example of a ventilator 17. FIG. 4 is a flowchart showing the operation of the control device main body 33. It is a graph which shows the concept of the change of the evaluation value F of robust stability and quick response in the case of changing the time constant Tm of the primary delay element and the dead time Lm of the dead time element. It is a graph which shows assist pressure Paw and ventilation flow Qi using time constant Tm and dead time Lm which constitute the 2nd maximum point M2. It is a block diagram which shows the whole system | strain 13 of further another embodiment of this invention. It is a block diagram which shows the whole system | strain 12 of further another embodiment of this invention. It is a block diagram which shows the whole system | strain 10 of further another embodiment of this invention. It is a graph for demonstrating airway resistance R. FIG.

It is a graph for demonstrating the compliance of a lung. 1 is a block diagram showing an entire system 5 including a prior art ventilator 1 and a patient 2. FIG. It is a graph which shows the relationship between the ventilation volume Vmus at the time of spontaneous respiration in the whole system 5 of a prior art, and the ideal ventilation volume Vast at the time of assist breathing.

Explanation of symbols

14 Overall system 17 Ventilator 20 Gas supply mechanism 21 Control device 50 Flow rate detection means 51 Flow rate estimation means 52 Deviation calculation means 53 Control amount calculation means 54 Observer Pmus Spontaneous respiration pressure Pin Target pressure Pvent Support pressure F Actual support gas Flow rate ^ F Estimated support gas flow rate ΔF Flow rate deviation

Claims (4)

In a control device for controlling a gas supply mechanism for a ventilator that supplies a support gas containing oxygen at a support pressure Pvent to an airway of a patient via an inspiratory line in response to a signal representing a target pressure Pin.
A flow rate detecting means for detecting a flow rate F of the support gas flowing through the intake pipe;
A delay compensation means for calculating the output according to a predetermined adjustment transfer function as a delay compensation pressure ^ Pm with the target pressure Pin as an input;
In response to the delay compensation pressure ^ Pm calculated by the delay compensation means, using the respiratory organ model that models the patient's respiratory organ, the assistance gas with the delay compensation pressure ^ Pm will flow through the inspiratory line A flow rate estimating means for calculating a gas flow rate as an estimated flow rate ^ F;
Deviation calculating means for calculating a flow deviation ΔF between the flow rate F detected by the flow rate detecting means and the estimated flow rate ^ F calculated by the flow rate estimating means;
A control amount calculating means for calculating a target pressure Pin in response to the flow rate deviation ΔF calculated by the deviation calculating means and providing a signal representing the target pressure Pin to the gas supply mechanism and the delay compensating means. ,
The adjustment transfer function includes a first-order lag element and includes a time delay element, and a time delay Lm in the time delay element is set as zero.
The time constant Tm of the first-order lag element of the adjustment transfer function is larger than the time constant Tc approximating the first-order lag element included in the transfer function of the gas supply mechanism. Control device. The time constant Tm of the first-order lag element of the adjustment transfer function approximates the time constant Tc that approximates the first-order lag element included in the transfer function of the gas supply mechanism and the time delay element included in the transfer function of the gas supply mechanism. 2. The control device for a gas supply mechanism according to claim 1, wherein the control device is set to be smaller than a value (Tc + Lc) obtained by adding the dead time Lc. 3. The gas supply mechanism according to claim 1, wherein the time constant Tm of the first-order lag element of the adjustment transfer function is set to an upper limit according to an amplification factor for proportionally amplifying the spontaneous breathing pressure Pmus. Control device. In response to a signal representing the target pressure Pin, a gas supply mechanism for a ventilator that supplies a support gas containing oxygen at the support pressure Pvent to the inspiratory line using a control device including an arithmetic processing circuit and an interface A control method for controlling,
A delay compensation pressure calculation step of calculating, as a delay compensation pressure ^ Pm, an output in accordance with a predetermined adjustment transfer function for compensating for the delay, with the target pressure Pin as an input by an arithmetic processing circuit;
In response to the delay compensation pressure ^ Pm calculated by the delay compensation pressure calculation step by the arithmetic processing circuit, the support gas of the delay compensation pressure ^ Pm is supplied to the intake pipe using a respiratory organ model that models the respiratory organ of the patient. A flow rate estimating step for calculating the flow rate of the support gas that will flow through the road as an estimated flow rate ^ F;
A deviation calculating step of calculating a flow rate deviation ΔF between the flow rate F of the assisting gas flowing through the intake pipe detected by the flow rate detecting means and the estimated flow rate of assisting gas ^ F calculated by the flow rate estimating step. When,
The control processing circuit calculates a target pressure Pin in response to the flow rate deviation ΔF calculated by the deviation calculating step, and gives a control amount for giving a signal representing the target pressure Pin to the gas supply mechanism via the interface. Calculation process,
The adjustment transfer function used in the delay compensation pressure calculation step includes a first-order lag element and also includes a dead time element, and the time constant Tm in the first-order lag element is determined by the transmission of the gas supply mechanism. Control of the gas supply mechanism using the control device, wherein a value larger than the time constant Tc approximating the first-order lag element included in the function is adopted, and the dead time Lm in the dead time element is set as zero Method.

Images