JP2008000372A - Control method and controller of gas supply mechanism - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a control method of the gas supply mechanism of an artificial respirator capable of preventing the practical delay of an inspiration supporting operation and supporting a highly similar respiratory operation to the respiratory efforts of a patient. <P>SOLUTION: The point T<SB>onset</SB>of time of starting inspiration is detected and a compensation value H to be lowered with the lapse of time is generated. Then, on the basis of a compensation flow rate F<SB>H</SB>for which the compensation value H is added to a detected flow rate F, the target pressure Pin of a support gas is computed. Thus, the inspiration support operation by a gas supply servo mechanism is practically started from the point T<SB>onset</SB>of time of starting the inspiration and the practical delay of the inspiration support operation is prevented. Also, since the compensation value H is lowered with the lapse of time, after an initial period W<SB>H</SB>elapses in an inspiration period, the influence of the compensation value H included in a compensation flow rate value F<SB>H</SB>is reduced. By reducing the influence of the compensation value H in such a manner, the inspiration support operation is prevented from becoming excessive. <P>COPYRIGHT: (C)2008,JPO&INPIT

Description

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

There are ventilators that support the patient's inspiration when the patient breathes spontaneously. The ventilator has a gas supply servo mechanism (hereinafter simply referred to as a gas supply mechanism) that supplies support gas to the patient's airway during the inspiration period. Proportional assisted ventilation method (proportional assisted ventilation method, Proportional Assist method)
Ventilation, abbreviated PAV method).

A ventilator that realizes the PAV method includes a gas supply mechanism and a control device that controls the gas supply mechanism. The gas supply mechanism supplies support gas to the patient's airway via the inspiratory line. The control device determines a time point when the flow rate of the assist gas supplied to the patient's airway exceeds a predetermined set value of 0 or more as an inspiration trigger time point T _onrg and performs an intake air support operation from the inspiration trigger time point T _onrg. . Further, the control device determines a target pressure value Pin of the support gas discharged from the gas supply mechanism toward the airway based on the flow rate F of the support gas flowing through the airway. For example, in Patent Document 1, the target pressure value Pin is determined by the following equation (a).
Pin = A · ^ Pmus = A (^ R · F + ^ E · ∫F) (a)

Here, Pin indicates the target pressure value of the support gas discharged from the gas supply mechanism toward the airway, and A indicates the amplification gain. ^ Pmus indicates the estimated spontaneous breathing pressure. ^ R indicates the estimated patient airway resistance, and ^ E indicates the estimated patient's lung elastance. F indicates the flow rate of the support gas flowing through the airway in the direction toward the lungs. The ∫F is an integrated value of the flow rate F of the supporting gas flowing through the airway in the direction towards the lungs from the suction start time T _onset, showing the volume of the supporting gas fed into the lungs from the suction start time T _onset .

  As shown in the equation (a), the target pressure value Pin is a value (A · ^ Pmus) obtained by multiplying the estimated spontaneous breathing pressure ^ Pmus by the amplification gain A. By appropriately setting the amplification gain A, the inspiration support operation can be performed with the spontaneous breathing pressure Pmus, that is, with a force proportional to the patient's breathing effort.

Japanese Patent Publication No. 2714288

FIG. 18 is a diagram illustrating a simulation result when the respiratory rate per unit time is small. FIG. 18 (1) shows the time change of the spontaneous breathing pressure Pmus simulating the spontaneous breathing of the patient. FIG. 18 (2) shows the time change of the estimated spontaneous breathing pressure ^ Pmus calculated based on the flow rate F of the support gas. FIG. 18 (3) shows the change over time of the flow rate F of the assisting gas flowing through the patient's airway when the estimated spontaneous breathing pressure ^ Pmus is amplified at a predetermined amplification factor A. The simulation results in FIG. 18 are obtained using the parameters shown in Table 1. FIG. 18 is a simulation of the case where the intake support operation is performed from the intake start time T _onset .

In the simulation result shown in FIG. 18, the respiration rate per minute of the patient is set to 12 times. In the initial state where the control is started, since the flow rate F of the support gas is small, the intake support operation based on the flow rate F is also small, and the substantial intake support operation cannot be performed. For example, a time point when the detected flow rate F _{oneff of the} support gas reaches 1 liter / minute is defined as a support start time point of a substantial intake support operation, and this time point is defined as a real support start time point T _onff .

As shown in FIG. 18, in the simulation considering delay elements such as airway resistance, the substantial support start time T _oneff is delayed with respect to the rising start time T _onset of the spontaneous breathing pressure Pmus. The time delay Td between the rising start time T _onset of the spontaneous breathing pressure Pmus and the substantial support start time T _oneff is about 100 msec in FIG. As described above, in the simulation result shown in FIG. 18, even if the support for the inhalation operation is started from the start time of the inspiration, the substantial inhalation support operation by the ventilator is delayed by about 100 msec from the start time of the inspiration.

FIG. 19 is a diagram illustrating a simulation result simulating a state in which the lung is in dynamic hyperinflation. FIG. 19 (1) shows the time change of the spontaneous breathing pressure Pmus simulating the spontaneous breathing of the patient. FIG. 19 (2) shows the time change of the estimated spontaneous breathing pressure ^ Pmus calculated based on the flow rate F of the support gas. FIG. 19 (3) shows the time change of the flow rate F of the assisting gas flowing through the patient's airway when the estimated spontaneous breathing pressure ^ Pmus is amplified at a predetermined amplification factor A. The simulation results in FIG. 19 are obtained using the parameters shown in Table 1 described above. FIG. 19 is a simulation of the case where the intake support operation is performed from the intake start time T _onset .

The simulation results shown in FIG. 19 indicate that the lung condition is dynamic overexpansion (Dynamic
Hyperinflation) is simulated. Dynamic _{overexpansion} indicates a state in which the pressure in the lung is higher than the airway pressure at the inspiration start time T _onset , that is, a case where residual pressure exists in the lung. In this case, at the start time T _{onset of} inspiration, the support gas flows through the airway in the direction of escaping from the lungs. When the patient's respiration rate per unit time is large and the compliance C, which is the reciprocal of the airway pressure R and elastance, is large, dynamic hyperinflation tends to occur.

As shown in FIG. 19, in a state where the patient's lungs are in dynamic hyperinflation, the substantial support start time T _oneff is greatly delayed from the rise start time T _onset of the spontaneous breathing pressure Pmus. Further, the difference between the spontaneous breathing pressure Pmus and the estimated spontaneous breathing pressure ^ Pmus is large. In the simulation result shown in FIG. 19, the time delay Td between the rise start time T _onset of the spontaneous breathing pressure Pmus and the substantial support start time T _oneff is about 800 msec. As described above, in the simulation result shown in FIG. 19, when dynamic overexpansion occurs, even if the support for the inspiratory operation is started from the inspiration start time, the substantial inspiratory support operation by the ventilator is performed from the inspiration start time. It will be delayed by about 800 msec.

  Further, the pressure difference ΔPmus between the maximum value of the spontaneous breathing pressure Pmus and the maximum value of the estimated spontaneous breathing pressure ^ Pmus increases. The lower the similarity between the estimated spontaneous breathing pressure ^ Pmus and the actual spontaneous breathing pressure Pmus, the target pressure Pin (= A · ^ Pmus) is a value obtained by proportionally amplifying the spontaneous breathing pressure Pmus (A · Pmus). ).

FIG. 20 is a graph schematically showing a time change between the spontaneous breathing pressure Pmus and the flow rate F of the assisting gas flowing through the airway. FIG. 20 (1) shows the time change of the spontaneous breathing pressure Pmus, and FIG. 20 (2) shows the time change of the flow rate F of the support gas. FIG. 20 is a simulation of the case where the intake support operation is performed from the intake start time T _onset .

  In FIG. 20 (1), the spontaneous breathing pressure Pmus is indicated by a solid line. The estimated spontaneous breathing pressure ^ Pmus (= ^ R · F + ^ E · ∫F) is indicated by a broken line. Further, the flow-induced pressure value ^ R · F and the volume-derived pressure value ^ E · ∫F constituting the estimated spontaneous breathing pressure ^ Pmus are indicated by a one-dot chain line and a two-dot chain line, respectively. The spontaneous breathing pressure Pmus is proportional to the magnitude of the patient's breathing effort during inspiration. Accordingly, the patient's respiratory effort increases and the spontaneous breathing pressure Pmus increases.

As shown in FIG. 20, in the state where the lung is in dynamic hyperinflation, the flow rate of the assisting gas flowing through the airway to the lung becomes 0 or less at the inspiration start time T _onset when the patient started the breathing effort. In other words, the patient may begin a breathing effort with the support gas flowing in the direction that is expelled from the lungs. In this case, in the initial period from the inhalation start time T _onset , the flow rate of the support gas becomes zero or less, and the spontaneous breathing pressure Pmus cannot be estimated from the flow rate F of the support gas. Only after the initial period has elapsed can the substantial intake assist operation be performed. The substantial support start time T _oneff has a time delay Td with respect to the intake start time T _onset . Further, a pressure difference ΔPmus is generated at the maximum value between the spontaneous breathing pressure Pmus and the estimated spontaneous breathing pressure ^ Pmus.

As described above, when the delay Td occurs at the real inspiration support time point T _oneff , the similarity between the time change waveform of the spontaneous breathing pressure Pmus and the time change waveform of the target pressure value Pin becomes low. Therefore, even if the estimated spontaneous breathing pressure ^ Pmus is multiplied by the amplification gain A to obtain the target pressure Pin, it is a support for a breathing operation having a low correlation with the patient's breathing effort. Specifically, the breathing motion cannot be supported for the period from the inspiration start time T _onset to the inspiration trigger time T _oneff . In addition, the synchronization between the patient's inhalation operation and the substantial inhalation support operation of the ventilator is reduced. Further, the estimated spontaneous respiration pressure ^ Pmus changes later than the spontaneous respiration pressure Pmus, and the maximum value of the estimated spontaneous respiration pressure ^ Pmus is lower than the spontaneous respiration pressure Pmus, which increases the burden on the patient. End up.

FIG. 21 is a graph showing temporal changes in the spontaneous breathing pressure Pmus and the flow rate F of the assisting gas flowing through the airway. FIG. 21 shows a case where the spontaneous respiration pressure ^ Pmus is determined by the following equation (b), and the inspiratory support operation by the ventilator is performed from the inspiration start time T _onset . FIG. 21 is a simulation of the case where the intake support operation is performed from the intake start time T _onset .
^ Pmus = ^ R · (F + | F _onset |)
+ ^ E · ∫ (F− | F _onset |) (b)

Here, symbols common to the above-described equation (a) have the same meaning as in equation (a). Further, | F _onset | indicates the flow rate F of the support gas flowing through the airway at the inspiration start time T _onset . The flow rate F is a value representing the flow rate F of the support gas at the inspiration start time T _onset as an absolute value in order to make the flow rate value in the direction of flowing toward the lung a positive value. In FIG. 21 (2), the detected flow rate F is indicated by a solid line. Further, a value (F + | F _onset |) obtained by adding the flow rate F _onset at the start of inspiration, which is a constant value, to the detected flow rate F is indicated by a broken line.

In FIG. 21 (1), the spontaneous breathing pressure Pmus is indicated by a solid line. The estimated spontaneous breathing pressure ^ Pmus is indicated by a broken line. Further, the flow rate-derived pressure value ^ R · (F + | F _onset |) and the volume-derived pressure value ^ E · ∫ (F + | F _onset |) constituting the estimated spontaneous breathing pressure ^ Pmus are indicated by a one-dot chain line and a two-dot chain line, respectively. . In this case, the actual support start time T _oneff is prevented from being delayed with respect to the inhalation start time T _onset at which the patient started the inhalation operation, and the inspiratory support operation by the ventilator is started from the time near the inspiration start time T _onset. Can be substantially started.

In the case shown in FIG. 21, the estimated spontaneous breathing pressure ^ Pmus is determined based on a value (F + | F _onset |) obtained by adding the flow rate F _onset at the start of inspiration, which is a constant value, to the detected flow rate F. The estimated spontaneous breathing pressure ^ Pmus has an effect of adding the flow rate F _onset at the start of inspiration to the detected flow rate F over the entire inhalation period. Due to this influence, the estimated spontaneous breathing pressure ^ Pmus becomes extremely larger than the spontaneous breathing pressure Pmus throughout the inspiration period. In this case, the similarity between the time change waveform of the spontaneous breathing pressure Pmus and the time change waveform of the target pressure value Pin becomes low. Therefore, there is a possibility that the respiration action may be different from a value proportional to the spontaneous breathing pressure Pmus.

As described above, even if the control start time for performing the inspiratory support operation is set to the inhalation start time T _onset at which the patient starts the inspiratory operation based on the conventional technique, the substantial support start time T _{onset with} respect to the inhalation start time T _onset . _Due to the delay in _oneff , as shown in FIG. 20, there is a possibility that the respiratory support by the ventilator may support the respiratory action having a low correlation with the spontaneous respiratory pressure Pmus. There's a problem. In addition, when equation (b) is used to accelerate the real support start time T _oneff , there is a problem that the spontaneous breathing support becomes excessive. In addition to the PAV method, the same problem occurs when the ventilator is controlled based on the flow rate F of the support gas.

Therefore, an object of the present invention is to prevent a substantial delay in the start of support with respect to the inspiration start time point T _{onset and} to prevent the intake support operation from becoming excessive. It is an object of the present invention to provide a control method and control device for a gas supply mechanism of a ventilator capable of providing support.

The present invention is a control method for controlling a gas supply servomechanism for supplying a support gas containing oxygen to a patient's airway via an inspiratory line for a patient performing a spontaneous breathing operation,
An inspiration start time detection step for detecting an inspiration start time T _{onset at} which the patient's respiratory muscles start inspiration;
A compensation value calculating step of calculating a compensation value H that decreases with the passage of time from the intake start time T _onset ;
A compensation flow value generation step of detecting the flow rate F of the support gas flowing through the patient's airway from the inhalation start time T _onset and generating the compensation flow rate value F _H obtained by adding the compensation value H to the detected flow rate F of the support gas. When,
A control amount calculation step of calculating a target pressure value Pin of the support gas to be discharged from the gas supply servomechanism to the intake pipe from the intake start time T _onset based on the compensation flow rate value F _H ;
And a control command step for giving a command to the gas supply servo mechanism so that the support gas is discharged toward the intake pipe at the target pressure value Pin calculated by the control amount calculation step. This is a control method of the servo mechanism.

According to the present invention, the compensation flow rate value F _H that decreases with the lapse of time after the detection of the intake start time T _onset is generated. Further, the target pressure value Pin of the support gas is calculated based on the compensation flow rate value F _H obtained by adding the compensation value H to the detected flow rate F. Accordingly, even when the detected flow rate F is equal to or less than zero, such as when the assist gas is exhausted from the lung at the inspiration start time T _onset , the gas supply servo mechanism substantially does not respond in response to the generation of the compensation value H. Control can be started. It is possible to prevent a delay in the intake assist operation.

Further, since the compensation value H decreases with the passage of time, after the compensation value has decreased, the compensation flow rate value F _H is less affected by the compensation value H, and the influence of the detected flow rate F of the detected assist gas is greater. Become. By reducing the influence of the compensation value H in this way, it is possible to prevent the support operation from becoming excessive when the control is performed based on the compensation flow rate F _H.

The present invention, compensation value H for calculating by the compensation value calculation step, the maximum value of predetermined intake start time T _onset, it is set to a value that decreases smoothly with time from the intake start time T _onset It is characterized by.

According to the present invention, the compensation value H _reaches the maximum value at the intake start time T _onset and decreases smoothly with the passage of time from the intake start time T _onset . On the other hand, the flow rate of the support gas increases with the passage of time from the inspiration start time T _{onset due} to the increase in the spontaneous breathing pressure Pmus. Therefore, the compensation flow rate value F _H has the greatest influence of the compensation value H at the intake start time T _onset , and the influence of the compensation value H gradually decreases with time. As a result, the compensation flow rate value F _H can be smoothly changed over time.

Further, the present invention is characterized in that the compensation value H calculated in the compensation value calculation step is set to a transient response value obtained by incomplete differentiation of the maximum value from the intake start time T _onset .

According to the present invention, the intake start time T _onset, by incomplete differentiation of the set value predetermined, the compensation value H having a transient response that decreases smoothly with the maximum value and becomes time the intake beginning T _onset Can be easily generated according to an arithmetic expression. Further, by performing incomplete differentiation, the attenuation amount of the compensation value H can be increased with time.

  The present invention is characterized in that the transient change of the compensation value H calculated in the compensation value calculation step is set based on the transient change of the patient's spontaneous breathing pressure Pmus.

According to the present invention, the flow compensation value F _H is set for each transient change in the spontaneous breathing pressure Pmus of the patient, so that inhalation assistance according to the patient's condition can be performed, and the burden on the patient is reduced. be able to.

For example, when the rise time of the patient's spontaneous breathing pressure Pmus is short, the amount of attenuation of the compensation value H over time is set large. When the rise time of the spontaneous breathing pressure Pmus of the patient is long, the attenuation amount of the compensation value H with the passage of time is set small. As a result, the inspiratory support operation can be suitably performed during the period from when the patient starts the inspiratory effort until the support gas flows toward the lung, and the compensated flow rate value after the support gas flows toward the lung The influence of the compensation value H on F _H can be reduced.

Further, the present invention is characterized in that the maximum value of the compensation value H calculated in the compensation value calculation step is set to a value corresponding to the magnitude of the flow rate F _onset of the assist gas flowing through the airway at the intake start time T _onset . And

According to the present invention, the maximum value of the compensation value H is set to be larger as the magnitude of the flow rate F _onset of the assist gas flowing in the direction of being discharged from the patient's lung at the inspiration start time point T _onset . For example, the maximum value of the compensation value H is set to a value proportional to the magnitude (absolute value) of the assist gas flow rate F _onset at the intake start time T _onset . In the present invention, when the detected flow rate is negative, that is, the assist gas flows in the direction of being discharged from the lungs of the patient at the inspiration start time point T _onset , the compensation value H can be determined according to the flow rate. Thus, the compensation flow rate value can be raised from the inspiration start time T _onset , and the compensation value H can be changed according to the patient's condition.

  Further, according to the present invention, the maximum value of the compensation value H calculated in the compensation value calculation step is a flow rate change delay from when the patient's respiratory muscles start the inspiration operation until a change in the flow rate of the support gas according to the inspiration operation occurs. It is characterized by being set to a prescribed value H0 for compensating for.

  Due to the patient's airway resistance or the like, there may be a delay in the flow rate change after the patient's respiratory muscles start the inspiration operation until the change in the assist gas flow rate according to the inspiration operation is detected. According to the present invention, by setting the specified value H0 for compensating for the flow rate change delay as the maximum value of the compensation value, the spontaneous breathing pressure Pmus and the estimated spontaneous breathing pressure ^ Pmus caused by the flow rate change delay are set. Misalignment can be prevented.

Further, in the present invention, the compensation value H calculated in the compensation value calculation step is a value obtained by combining the first compensation value H1 and the second compensation value H2,
The first compensation value H1 is set to a transient response value obtained by incomplete differentiation of a value corresponding to the magnitude of the flow rate F _onset of the support gas flowing through the airway at the inspiration start time point T _onset .
The second compensation value H2 is set to a transient response value obtained by incompletely differentiating the specified value H0 for compensating for the flow rate change delay from the intake start time T _onset until the change in the flow rate of the support gas according to the intake operation occurs. It is characterized by that.

According to the present invention, since the compensation value H is determined based on the magnitude of the flow rate F _onset of the support gas at the inhalation start time T _onset , it is possible to perform inspiration support according to the patient's condition. Furthermore, since the compensation value H is determined based on the specified value H0 for compensating for the flow rate change delay, it is possible to prevent the difference between the spontaneous breathing pressure Pmus and the estimated spontaneous breathing pressure ^ Pmus caused by the flow rate change delay. Can do.

In the control amount calculation process, the present invention
A flow rate setting value (K _FG · F _H ) obtained by multiplying the compensation flow rate value F _H by a predetermined flow rate gain K _FG , and a volume obtained by multiplying the integral value V _H of the compensation flow rate value F _{H by} a predetermined volume gain K _VG. A total value ((K _FG · F _H ) + (K _VG · V _H )) obtained by adding the set value (K _VG · V _H ) is calculated as the target pressure Pin.

According to the present invention, by appropriately setting the flow gain K _FG and the volume gain K _VG , it is possible to supply the patient with a support gas having a pressure obtained by proportionally amplifying the spontaneous breathing pressure Pmus, and according to the so-called PAV method. The patient's breathing motion can be supported. Further, by using the compensation flow rate value F _H , it is possible to calculate the target pressure Pin that compensates for the delay in the inspiratory assist operation, and the detected flow rate F of the assist gas flowing toward the patient's lung at the inspiration start time T _onset is Even if it is less than or equal to zero, the control of the gas supply servo mechanism can be started.

In the control amount calculation process, the present invention
Using the flow rate estimation means that models the respiratory organs of the patient, the target pressure Pin is input, the output according to a predetermined adjustment transfer function is calculated as the delay compensation pressure Pm, and the support gas of the delay compensation pressure ^ Pm is calculated. A flow rate estimating step for estimating the flow rate of support gas to be supplied to the patient's airway when supplied, F;
A deviation calculating step of calculating a flow rate deviation ΔF _H between the compensation flow rate F _H and the estimated flow rate ^ F;
A flow rate setting value (K _FG · ΔF _H ) obtained by multiplying the flow rate deviation ΔF _H by a predetermined flow rate gain K _FG and a volume setting value obtained by multiplying an integral value ΔV _H of the compensation flow rate ΔF _{H by} a predetermined volume gain K _VG. characterized in that it comprises a target pressure calculating step of the _{(K VG} · ΔV _H) sum obtained by adding the _{((K FG · ΔF H)} + (K VG · ΔV H)) the target pressure Pin.

According to the present invention, by appropriately setting the flow gain K _FG and the volume gain K _VG , it is possible to supply the patient with a support gas having a pressure obtained by proportionally amplifying the spontaneous breathing pressure Pmus, and according to the so-called PAV method. The patient's breathing motion can be supported. Further, by obtaining the flow rate deviation ΔF _H from the compensated flow rate F _H and the estimated flow rate ^ F, it is possible to obtain the flow rate deviation ΔF _H that compensates for the delay in the intake assist operation. As a result, the control of the gas supply servomechanism can be started even if the flow rate F of the assisting gas flowing toward the patient's lung at the inspiration start time point T _onset is equal to or less than zero.

The compensation flow rate F _H varies depending on the patient's spontaneous breathing pressure Pmus, but the estimated assist gas flow rate F is not affected by the patient's spontaneous breathing pressure Pmus. Therefore, by obtaining the flow rate deviation ΔF _H between the compensated flow rate F _H and the estimated assist gas flow rate F, it is possible to accurately acquire information on the patient's spontaneous breathing pressure Pmus.

The control amount calculation step, based on the flow rate difference [Delta] F _H, calculates the target pressure Pin for controlling the gas supply servomechanism. By providing the target pressure Pin calculated in this way to the gas supply servomechanism, the support gas can be supplied to the patient's respiratory tract at a pressure corresponding to the spontaneous breathing pressure Pmus that changes sequentially.

Further, by calculating the target pressure Pin based on the flow rate deviation ΔF _H , the entire system including the patient and the ventilator can be compared with the case where the target pressure Pin is calculated based only on the compensation flow rate F _H. The positive feedback configuration can be made difficult, and a margin for the stability limit of the entire system can be increased. If this causes disturbance, if there is a time delay in the gas supply servomechanism, if the patient's respiratory organ model cannot be set accurately, or if the patient's condition changes, etc., runaway will be less likely to occur. Can do. As a result, the burden on the patient can be reduced.

The present invention, compensation value H for calculating by the compensation value calculation step, decreases with time from the intake start time T _onset, comprises a different flow compensation value H _F and volume compensation value H _V each other,
In the compensation flow value generation process,
The compensated flow value F _HF used to determine the flow rate set value in the control amount calculation step to generate a compensated flow value F _HF using the flow compensation value H _F,
The compensated flow value F _HV used to determine the volume setting value in the control amount calculation step, and generating a compensated flow value F _HV by using the volume compensation value H _V.

According to the present invention, the compensation flow rate value F _HF , which is different between the flow rate setting value (K _FG · F _H or K _FG · ΔF _H ) and the volume setting value (K _VG · V _H or K _VG · ΔV _H ), Use _FHV . These compensated flow value _{F HF, F HV} is obtained by using different compensation values _H F, the _{H V} each other. By using different compensation values H _F and H _V , it is possible to increase the parameters relating to adjustment of the breathing motion during the period in which the compensation works effectively. By appropriately setting the parameters, the estimated spontaneous breathing pressure ^ Pmus Can be made closer to the spontaneous breathing pressure Pmus.

Further, the present invention is characterized in that the intake start time detection step detects a time when a differential value obtained by differentiating a change in the flow rate of the support gas with respect to a predetermined threshold value increases as a predetermined threshold value as an intake start time T _onset. .

When the patient starts the inhalation operation, the time change of the support gas flow rate increases beyond a predetermined threshold. This is a phenomenon that occurs both when the support gas is flowing toward the lung and when the support gas is flowing in the direction of escape from the lung. According to the present invention, by detecting the inhalation start time point T _onset based on the time change of the flow rate differential value of the support gas, the support gas flows in the direction in which the support gas is discharged from the lungs at the inhalation start time point T _onset . Even in this case, the intake start time T _onset can be determined.

The present invention further includes an intake trigger detection step of detecting an intake trigger time point T _{ontrg in} which the flow rate F of the support gas flowing through the airway exceeds a predetermined intake start threshold value,
If the intake trigger time T _ontrg is detected before detecting the intake start time T _onset , the control gas calculation step detects the flow rate F of the support gas flowing through the airway, and based on the detected flow rate F of the support gas, The target pressure Pin of the assist gas discharged from the gas supply servo mechanism is calculated.

According to the present invention, since the support for the intake operation can be started from one of the intake start time T _onset and the intake trigger time T _ontrg detected at an earlier timing, the intake start time T _onset is temporarily set. Even when it cannot be detected, the start point of the intake assist operation can be determined.

The present invention also includes an inspiration end time detection step of detecting an inspiration end time T _{offset at} which the patient's respiratory muscles terminate inspiration.
An expiratory trigger detection step of detecting an expiratory trigger time point T _{offtrg at} which the flow rate F of the assisting gas flowing through the airway becomes less than a predetermined inhalation end threshold value;
The control command process is characterized in that the control of the gas supply servo mechanism related to the patient's inspiration support is terminated when it is determined that either the inspiration end time T _offset or the expiration trigger time T _offtrg has been reached.

According to the present invention, one of the intake end _{T offset} and expiration trigger time _{T Offtrg,} from the time it is detected at an early timing, it is possible to end the help of inspiration, if the intake end time _{T offset} Even when it cannot be detected, the end point of the intake assist operation can be determined.

  Further, the present invention is a program for causing a computer to execute the control method of the gas supply servo mechanism.

  According to the present invention, a program for realizing the above-described method is executed by a computer and the gas supply servomechanism is controlled by the computer, thereby preventing a delay in the intake assistance and taking in the intake air based on the detected flow rate F. It is possible to prevent the intake support operation from becoming excessive as compared with the case where support is provided.

Further, the present invention is a control device for controlling a gas supply servo mechanism for supplying a support gas containing oxygen to a patient's airway via an inspiratory line for a patient performing a spontaneous breathing operation,
An inspiration start time detection means for detecting an inspiration start time T _{onset at} which the patient's respiratory muscles start inspiration;
Compensation signal generating means for generating a compensation signal H _sig indicating a compensation value H that decreases with the passage of time from the intake start time T _onset ;
The flow rate F of the support gas flowing through the airway is detected from the intake start time T _onset , the compensation signal H _sig is superimposed on the flow rate signal F _sig indicating the detected flow rate F of the support gas, and the compensation is added to the detected flow rate F. A compensation flow rate signal generating means for generating a compensation flow rate signal F _Hsig indicating a compensation flow rate F _H to which the value H is added;
A control amount calculation means for calculating a target pressure value Pin of the support gas discharged from the gas supply servomechanism based on the compensation flow rate F _H indicated by the compensation flow rate signal F _Hsig from the intake start time _Tonset ;
Control command means for giving a command to the gas supply servo mechanism so as to discharge the support gas at the target pressure Pin toward the patient's airway at the target pressure value Pin calculated by the control amount calculation means, This is a control device for the gas supply servo mechanism.

According to the present invention, the compensation signal H _sig indicating the compensation value H that decreases with time is generated by the compensation signal generating means from the intake start time T _onset detected by the intake start time detecting means. Next, a compensation flow rate signal F _Hsig indicating a compensation flow rate value F _H obtained by adding the compensation value H and the detected flow rate F is generated by the compensation flow rate signal generating means. Based on the compensation flow rate value F _H indicated by the compensation flow rate signal F _Hsig , the target pressure value Pin of the support gas is calculated.

Further, according to the present invention, in the configuration described above, a signal indicating the inhalation start time T _onset is input from a detection device that detects the inhalation start time T _onset at which the patient's respiratory muscle starts inhalation operation instead of the inhalation start time detection means. And a control device for the gas supply servo mechanism.

According to the present invention, a signal indicating the intake start time T _onset is input from the detection device that detects the intake start time T _onset by the intake start time input means. After this signal is input and the intake start time T _onset is determined, a compensation flow rate signal F _Hsig indicating a compensation flow rate value F _H is generated by the compensation flow rate signal generating means. Based on the compensation flow rate value F _H indicated by the compensation flow rate signal F _Hsig , the target pressure value Pin of the support gas is calculated.

The present invention also provides a gas supply servomechanism for supplying a support gas containing oxygen to a patient's airway via an inspiratory line for a patient performing a spontaneous breathing operation,
And a control device for the gas supply servomechanism.

  According to the present invention, by including the control device described above, the ventilator can prevent a delay in inspiratory support and prevent the inspiratory support operation from becoming excessive.

According to the present invention described in claim 1, since the compensation value H is generated from the intake start time T _onset, it is possible to start the suction assistance operation by the gas supply servomechanism from the intake start time T _onset, the intake assistance operation A substantial delay can be prevented. As a result, the burden on the patient supported by inhalation by the ventilator can be reduced. Further, since the compensation value H decreases with the passage of time, the influence of the compensation value H can be reduced after the decrease compared to before the decrease, and the intake assist operation when the intake assist is performed based on the detected flow rate F. It is possible to prevent the intake support operation from becoming excessive. This can reduce the burden on the patient who is inhaled by the ventilator.

According to the second aspect of the present invention, the compensation flow rate value F _H can be changed smoothly with the passage of time by reducing the compensation value H smoothly with the passage of time from the intake start time T _onset. Can do. As a result, it is possible to prevent the pressure value of the assist gas supplied toward the patient's lung from changing rapidly in a short time, and to reduce the burden on the patient who is inhaled.

  According to the third aspect of the present invention, the compensation value H can be smoothly reduced over time by incomplete differentiation of the predetermined set value. As a result, it is possible to prevent the pressure value of the support gas from changing rapidly in a short time, and to reduce the burden on the patient who is inhaled.

  According to the present invention of claim 4, by setting the transient change of the compensation value H based on the transient change of the patient's spontaneous breathing pressure Pmus, it is possible to perform the inspiration support according to the patient's condition, The burden on the patient can be reduced.

According to the fifth aspect of the present invention, the compensation value H is set based on the flow rate F _{onset of the} support gas flowing through the airway at the intake start time T _onset . As a result, the compensation value H can be changed according to the patient's condition, and the compensation flow value can be prevented from becoming excessive or too small. Inhalation support can be suitably performed until it flows toward the front. For example, even if the patient's breathing motion is in a state of dynamic hyperinflation, inhalation assistance can be performed while preventing a substantial delay in inspiratory assistance of the ventilator.

  According to the sixth aspect of the present invention, the compensation value H is set based on the flow rate change delay from when the respiratory muscle of the patient starts the inspiration operation until the change in the flow rate of the support gas according to the inspiration operation occurs. Is done. As a result, it is possible to prevent an error between the spontaneous breathing pressure Pmus and the estimated spontaneous breathing pressure ^ Pmus due to the flow rate change delay that is substantially constant regardless of the patient's condition.

According to the present invention according to claim 7, and the flow rate F _onset of supporting gas flowing through the airway in the intake beginning T _onset, based on the specified value H0 be predetermined, the compensation value H is set. As a result, inhalation support according to the patient's condition can be performed, and a delay in changing the flow rate can be prevented. As a result, a substantial delay in inspiratory support of the ventilator can be prevented more reliably.

According to the eighth aspect of the present invention, by calculating the target pressure value Pin using the compensation flow rate value F _H , the target pressure Pin compensated for a substantial delay in the intake assist operation can be calculated. Even if the flow rate F _{onset of the} assist gas flowing toward the patient's lung at the start time point T _onset is equal to or less than zero, the assist gas having a pressure proportionally amplified to the spontaneous breathing pressure Pmus corresponding to the patient's breathing effort is supplied to the patient. Control of the gas supply servomechanism can begin to supply. In addition, the patient's inspiratory period and the substantial inspiratory support period by the ventilator can be synchronized.

According to the ninth aspect of the present invention, by calculating the target pressure value Pin using the flow rate deviation ΔF _H , it is possible to calculate the target pressure Pin that compensates for a substantial delay in the intake assist operation, and to start intake Even if the flow rate F of the assist gas flowing toward the patient's lung at the time point T _onset is less than or equal to zero, the gas supply servomechanism is configured to supply the patient with the assist gas having a pressure proportionally amplified to the spontaneous breathing pressure Pmus. Control can begin. In addition, the patient's inspiratory period and the substantial inspiratory support period by the ventilator can be synchronized.

Further, by using the flow rate deviation ΔF _H , the spontaneous breathing pressure Pmus can be accurately estimated. Furthermore, the robust stability of the control system including the patient and the ventilator can be improved, and even if the flow gain K _FG and the volume gain K _VG deviate from the values indicating the patient's condition, the control system is unstable. Can be prevented.

According to the tenth aspect of the present invention, compensation is different between the flow rate setting value (K _FG · F _H or K _FG · ΔF _H ) and the volume setting value (K _VG · V _H or K _VG · ΔV _H ). Flow rate values F _HF and F _HV are used. As a result, the parameters relating to the adjustment of the breathing motion can be increased, and the breathing motion can be flexibly adjusted so that the estimated spontaneous breathing pressure ^ Pmus approaches the spontaneous breathing pressure Pmus. For example, by appropriately setting the two compensation values H _F and H _V , it is possible to prevent the intake operation support during the compensation period from becoming excessive or too small, and to facilitate the desired intake operation support. Can do.

According to the eleventh aspect of the present invention, the time point at which the differential value obtained by differentiating the time change of the flow rate of the support gas increases from a predetermined threshold value is detected as the intake start time point T _onset . In this way the intake start time T _onset, even if flowing in a direction in which the supporting gas is discharged from the lungs, it is possible to start the suction assistance operation by the gas supply servomechanism from the intake start time T _onset, intake Delays in support operations can be prevented.

According to the twelfth aspect of the present invention, the support for the intake operation can be started from the time point detected at the earlier timing of either the intake start time point T _onset or the intake trigger time point T _ontrg . Therefore, even if the inhalation start time T _onset cannot be detected, the start time of the inspiratory support operation can be determined reliably, and the safety of the ventilator can be improved.

According to the thirteenth aspect of the present invention, the support for the inhalation operation can be terminated from the time point at which the inspiration end point T _offset or the expiration trigger point T _offtrg is detected at an earlier timing. Therefore, even if the inhalation end time T _offset cannot be detected, it is possible to reliably determine the end time of the inspiratory support operation, and to improve the safety of the ventilator.

  According to the present invention as set forth in claim 14, by causing a computer to execute a program for realizing the above-described method and controlling the gas supply servomechanism by the computer, it is possible to prevent a delay in intake assistance and to detect a detected flow rate. Compared with the case where the intake support operation is performed based on F, it is possible to prevent the intake support operation from becoming excessive.

According to the fifteenth aspect of the present invention, when the intake start time T _onset is detected by the detection means, the compensation flow rate signal F _Hsig indicating the compensation flow rate value F _H is generated by the compensation flow rate signal generation means. Based on the compensation flow rate value F _H indicated by the compensation flow rate signal F _Hsig , the target pressure value Pin of the support gas is calculated. By giving the calculation result to the gas supply servomechanism, it is possible to prevent a delay in the intake support and to prevent the intake support operation from becoming excessive as compared with the case where the intake support is performed based on the detected flow rate F.

According to the sixteenth aspect of the present invention, when the intake means T _onset is determined by the input means, the compensated flow signal F _Hsig indicating the compensated flow value F _H is generated by the compensated flow signal generating means. Based on the compensation flow rate value F _H indicated by the compensation flow rate signal F _Hsig , the target pressure value Pin of the support gas is calculated. By giving the calculation result to the gas supply servomechanism, it is possible to prevent a delay in the intake support and to prevent the intake support operation from becoming excessive as compared with the case where the intake support is performed based on the detected flow rate F.

  According to the present invention of claim 17, by including the above-described control device, the ventilator prevents a delay in inspiratory support, and inspiratory support operation as compared with the case where the inspiratory support is performed based on the detected flow rate F. Can be prevented from becoming excessive.

  FIG. 1 is a block diagram showing a ventilator 17 and a patient 18. A ventilator 17 according to an embodiment of the present invention controls a gas supply servo mechanism 20 (hereinafter simply referred to as a gas supply mechanism 20) that supplies the assist gas 16 to the airway of the patient 18 and the gas supply mechanism 20. And a patient respiratory state monitoring means 200. The gas supply mechanism 20 supplies the support gas 16 containing oxygen to the patient's airway 15 via the inspiratory line 25. 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.

The patient breathing state monitoring means 200 detects an inhalation start time T _onset when the patient 18 starts an inhalation operation and an inhalation end time T _offset at which the patient 18 finishes the inhalation operation in the respiratory state of the patient 18. The patient breathing state monitoring means 200 detects the flow rate and pressure of the support gas flowing through the inspiratory conduit 25 and the expiratory conduit 26, respectively, and based on the detection result, the inspiratory start time T _onset and the inspiratory end time T _offset To detect. For example, the patient respiratory condition monitoring means 200 can be realized by using the technique disclosed in the international patent application WO 2004/002561 A2. The patient breathing state monitoring means 200 gives signals indicating the detected inhalation start time T _onset and inhalation end time T _offset to the calculation unit of the control device 21.

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, and corresponds to the respiratory effort of the patient 18. Accordingly, when the respiratory effort, which is the work amount for the patient's breathing, increases, the spontaneous breathing pressure Pmus increases. 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 unit 50, an estimation unit 51, a compensation flow rate generation unit 201, a deviation calculation unit 52, and a control amount calculation unit 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 may be referred to as a detected flow rate F. Further, among the detected flow rates F, the flow rate of the support gas flowing in the airway in the direction toward the lung is set to a positive value. Further, among the detected flow rate F, the flow rate of the support gas flowing through the airway in the direction of being discharged from the lung is set to a negative value.

  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. More precisely, the detected flow rate F is obtained by subtracting the flow rate Qe of the gas flowing in the direction of exhausting the exhalation conduit 26 from the patient's lungs from the flow rate Qi of the gas flowing in the direction of the inspiratory conduit 25 toward the patient's airway. The obtained value (Qi−Qe) is obtained. 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 detecting means 50 measures the flow rate of the support gas 16 flowing through the inspiratory line 25 and the expiratory line 26. 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. The exhalation duct 26 is a duct that guides support gas discharged from the patient's airway to the outside of the ventilator. 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 a signal indicating the detected flow rate F to the compensation flow rate generation unit 201 and the patient breathing state monitoring unit 200.

The compensation flow rate generation unit 201 receives a signal indicating the inhalation start time T _onset from the patient breathing state monitoring unit 200 _and a signal indicating the detected flow rate F from the flow rate detection unit 200. Based on these signals, the compensation flow rate generation unit 201 generates a signal indicating a compensation flow rate F _H that compensates for a flow rate change delay with respect to the generation of the spontaneous breathing pressure Pmus. Specifically, the compensation flow rate generation unit 201 generates a predetermined compensation value H, and adds the compensation value H to the detected support gas flow rate F to generate a compensation flow rate value F _H (= F + H). To do.

The compensation flow rate generation unit 201 generates a compensation value H that decreases with the passage of time from the intake start time T _onset . Compensation value H becomes the maximum value at the intake beginning _{T onset,} is set to a value that decreases smoothly with time from the intake start time _{T onset.} The compensation flow rate generation unit 201 gives a signal indicating the generated compensation flow rate F _H to the deviation calculation unit 52. In this embodiment, the magnitude of the compensation value H is the initial period W _H is set to be shorter than the intake period W _in the patient from the suction start time T _onset, the threshold value or values determined in advance. Further, after the initial period _WH has elapsed, the value becomes less than the threshold value. Period compensation value H is generated may be the same as the intake period W _in, it may be a shorter period than the intake period.

  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 calculating means 52 is provided with a signal indicating the compensated flow rate F _H from the compensated flow rate generating means 201 and a signal indicating the estimated flow rate ^ F from the estimating means 51. Based on these signals, the deviation calculating means 52 calculates a flow rate deviation ΔF _H that is a value obtained by subtracting the estimated flow rate ^ F from the compensation flow rate F _H, and gives a signal indicating the calculation result to the control amount calculating means 53. The control amount calculation unit 53 is given a signal indicating the flow rate deviation ΔF _H from the deviation calculation unit 52. The control amount calculation means 53 gives a preset gain (K _FG + K _VG / s) to the flow rate deviation ΔF _H 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 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} .

As an example of the first aspect of the present invention, for example, the time constant Tm of the first-order lag element of the adjustment transfer function and the dead time element e ^{−Lm · s} are measured using the target pressure Pin as an input and the support pressure Pvent as an output. A parameter value simulating the transfer function of the gas supply mechanism 20 is used.

  Further, a parameter value different from the transfer function of the gas supply mechanism 20 may be used for the adjustment transfer function. As an illustration of the second aspect of the present invention, for example, 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.

  As an example of the third aspect of the present invention, the dead time Lm in the dead time element of the adjustment transfer function may be set as zero. The time constant Tm of the temporary delay element of the adjustment transfer function is set to an upper limit value according to the amplification factor that proportionally amplifies the spontaneous breathing pressure Pmus. Further, the adjustment transfer function set in the delay compensation unit 55 may be set in addition to the control element and the parameter of the control element described above.

  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 integrates the estimated flow rate {circumflex over (F)} calculated by the estimated flow rate calculator 57 from the start of supply of the support gas, and calculates the integrated value as the support gas volume {circumflex over (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 this 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 O) / (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.
^ Paw− ^ Palv = ^ R · ^ F (1)
∫ ^ F = ^ V (2)
^ Palv = ^ E · ^ V (3)

  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 {circumflex over (F)} from the start of the supply of the support gas is equal to the calculation volume {circumflex over (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 by the following equation (4).
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 compensation flow rate generation unit 201 generates a signal indicating a compensation flow rate F _H that compensates for a change in flow rate with respect to the intake start time T _onset . The compensation flow rate generation unit 201 includes a signal generator 210, a first waveform shaper 211, a second waveform shaper 212, a flow rate regulator 213, a volume regulator 214, a first adder 215, and a second An adder 216 and a third adder 217 are included. FIG. 3 is a timing chart for explaining signals generated by the compensation flow rate generation means 201.

FIG. 3 (1) shows the time change of the spontaneous breathing pressure Pmus of the patient. FIG. 3 (2) shows a time change of the value of the output signal PVIsig of the patient respiratory condition monitoring means 200. FIGS. 3 (3) and 3 (6) show values of the first signal Sig1 and the second signal Sig2 output from the signal generator 210, respectively. 3 (4) shows a flow compensation value H _F represented by the signal output from the flow controller 213, FIG. 3 (5), the volume compensation value H represented by the signal output from the volume adjuster 214 _V is shown. 3 (7) shows a base compensation value H _d represented by the signal output from the second wave shaper 212.

The signal generator 210 is given a signal indicating the inspiration start time T _onset from the patient respiratory condition monitoring means 200. As shown in FIG. 3B, this signal has a rectangular waveform that rises at the intake start time T _onset and falls at the intake end time T _offset . When determining the intake start time T _onset , the signal generator 210 generates two signals Sig1 and Sig2 from the intake start time T _onset to the intake end time T _offset .

The signal generator 210 acquires the flow rate F _{onset of the} support gas flowing through the airway from the flow rate detection means 50 at the inspiration start time point T _{onset and} outputs a signal indicating a value corresponding to the acquired flow rate F _onset as the first signal Sig1. To do. Specifically, the absolute value of the acquired flow rate F _onset is used. The value indicated by the first signal Sig1 is a rectangular waveform that rises at the intake start time T _onset and falls at the intake end time T _offset as shown in FIG. 3 (3).

Signal generator 210, the larger the absolute value of the detected flow rate _{F onset} in the intake beginning _{T onset,} increasing the value of the first signal Sig1 is shown. In the present embodiment, among the values indicated by the first signal Sig1, the magnitude from the intake start time _Tonset to the intake end time T _offset matches the magnitude of the detected flow rate F _onset at the intake start time T _onset. Is set to the value to be

Further, when the signal generator 210 determines the intake start time T _onset , the signal generator 210 outputs a signal indicating a predetermined specified value H0 as the second signal Sig2. As shown in FIG. 3 (6), the value indicated by the second signal Sig2 is a rectangular waveform that rises at the intake start time T _onset and falls at the intake end time T _offset .

  The signal generator 210 sets the magnitude of the value indicated by the second signal Sig2 according to the specified value H0. The specified value H0 is set to a value for compensating for a flow rate change delay from when the patient's respiratory muscles start the inspiration operation to when the support gas flow rate change according to the inspiration operation occurs. In the present embodiment, the prescribed value H0 is determined to be a value proportional to the time spent until the estimated airway pressure ^ Pmus is calculated after the patient's respiratory muscles start inspiration. The prescribed value H0 is determined based on the patient's airway resistance and ventilator duct resistance, the detection delay of the airway flow rate, and the calculation delay of the estimated airway pressure Pmus. The specified value H0 hardly fluctuates greatly depending on the patient's condition. In the present embodiment, the prescribed value H0 is set to 100 milliliters / second.

The signal generator 210 provides the first signal Sig 1 to the first waveform shaper 211. The first waveform shaper 211 outputs a signal that decreases with the passage of time from the intake start time T _onset based on the value indicated by the first signal Sig1 provided from the signal generator 210. In the present embodiment, the first waveform shaper 211 outputs, as an output signal, a signal indicating a value obtained by incomplete differentiation of the value indicated by the first signal Sig1 as an input value.

  Specifically, in the first waveform generator 211, a transfer function represented by (a1 · s) / (a2 · s + 1) is set. A signal indicating an output value when the value indicated by the first signal Sig1 is input to this transfer function is output from the first waveform shaper 211 as an output signal. Here, a1 and a2 indicate preset coefficients, and in the present embodiment, a1 and a2 are set to 0.5, respectively.

The first waveform shaper 211 gives a shaping signal to the flow rate regulator 213 and the volume regulator 214. Flow regulator 213, a value indicated by the signal provided from the first waveform shaper 211, by multiplying the flow rate adjustment gain K _FH be predetermined, as shown in FIG. 3 (4) shows a flow compensation value H _F Generate a signal. Flow regulator 213, a signal indicative of the generated flow compensation value _{H F,} applied to the first adder 215.

The volume adjuster 214, the value indicating the signal supplied from the first waveform shaper 211, by multiplying the volume adjustment gain K _VH the predetermined, as shown in FIG. 3 (5), a volume compensation value H _V The signal shown is output. The volume adjuster 214, a signal indicating the generated volume compensation value _{H V,} applied to the second adder 216.

In the present embodiment, the flow rate adjustment gain K _FH is set smaller than the volume adjustment gain K _VH (K _FH <K _VH ). Specifically, the flow rate adjustment gain K _FH is set to 0.5, and the volume adjustment gain K _VH is set to 1. This 3 (4), as shown in FIG. 3 (5), a flow compensation value H _F and a volume compensation value H _V shows the different transient response change was. The flow compensation value H _F and volume compensation value H _V is a first compensation value H1 calculated based on the value indicated by the first signal Sig1 signal generator 210.

Further, the signal generator 210 gives the second signal Sig 2 to the second waveform shaper 212. Based on the value indicated by the second signal Sig2 provided from the signal generator 210, the second waveform shaper 212 forms a signal that decreases with the lapse of time from the intake start time T _onset . In the present embodiment, the second waveform shaper 212 outputs a signal indicating a value obtained by incomplete differentiation of the value indicated by the second signal Sig2 as a shaping signal.

Specifically, in the second waveform generator 212, a transfer function represented by (b1 · s) / (b2 · s + 1) is set. A signal indicating an output value when the value indicated by the second signal Sig2 is input to this transfer function is output from the second waveform shaper 212 as an output signal. Here, b1 and b2 indicate preset coefficients, and in the present embodiment, b1 and b2 are respectively set to 0.2. Signal output from the second waveform shaper 212, the base compensation value H _d as the second compensation value is calculated based on the value of the second signal Sig2 signal generator 210. The second waveform generator 212 provides a signal indicating the generated base compensation value H _d to the third adder 217.

The first adder 215 is given a signal indicating the detected flow rate F from the flow rate detection means 50. First adder 215, and the detected flow F indicated by the signal supplied from the flow rate detecting unit 50, by adding the flow rate correction value H _F indicated by the signal supplied from the flow regulator 213, added value (F + H _F) Is supplied to the third adder 217. As described above, the detected flow rate F has a positive value when it flows through the airway and flows out to the lungs. A negative value is defined when flowing from the lungs to the external space through the airway.

The third adder 217 receives the signal indicating the calculation result from the first adder 215 and the signal generated by the second waveform shaper 212. The third adder 217 adds the value (F + H _F ) indicated by the signal provided from the first adder 215 and the base compensation value H _d indicated by the signal provided from the second waveform shaper 212, and calculates the result ( F + H _F + H _d ) is given to the deviation calculating means 52.

The second adder 216 receives a signal indicating the detected flow rate from the flow rate detection means 50. The second adder 216, and the detected flow F indicated by the signal supplied from the flow rate detecting unit 50, by adding the volume correction value H _V indicated by the signal supplied from the volume adjuster 214, added value (F + H _V) Is given to the deviation calculating means 52. As described above, the detected flow rate F has a positive value when it flows through the airway and flows out to the lungs. A negative value is defined when flowing from the lungs to the external space through the airway.

Thus compensated flow generating means 201, a signal indicative of the detected flow F to flow compensation value _{H F} and base compensation value _{H d} and the flow rate term compensation rate obtained by adding the _{F HF (= F + H F} + H d), the detected flow F generating a signal indicating a volume compensation value _H volumetric term compensation flow by adding the _{V F HV (= F + H} V) to. The flow rate compensation flow rate F _HF and the volume term compensation flow rate F _HV are expressed by the following equation (5).
F _HF = F + | F _onset | · K _FH · G ₂₁₁ + H0 · G ₂₁₂
F _HV = F + | F _onset | · K _VH · G ₂₁₁
G ₂₁₁ = (a1 · S) / (a2 · s + 1)
G ₂₁₂ = (b1 · S) / (b2 · s + 1) (5)

Here, F _HF indicates a flow rate compensation flow rate, and F _VF indicates a volume term compensation flow rate. F indicates the flow rate detected by the flow rate detection means 50. Further, | F _onset | represents the absolute value of the detected flow rate at the intake start time T _onset . H0 indicates a predetermined specified value H0. K _FH represents a flow rate adjustment gain, and K _VH represents a volume adjustment gain. G ₂₁₁ and G ₂₁₂ are transfer functions for realizing a transient response change for decreasing the compensation values H _V , H _F , and H _d with the lapse of time. In this embodiment, incomplete differentiation is used. It is done. a1, a2, b1, and b2 are coefficients for performing incomplete differentiation, and s indicates a Laplace operator.

The deviation calculating means 52 includes a first subtracter 218 and a second subtractor 119. The first subtracter 218 outputs a signal indicating a flow rate compensation flow deviation ΔF _HF obtained by subtracting the estimated flow rate ^ F given from the estimation means 51 from the flow rate compensation flow rate F _HF . The second subtractor 119 obtains a volume term compensation flow deviation ΔF _HV by subtracting the estimated flow rate ^ F given from the estimation means 51 from the volume term compensation flow rate F _HV , and a signal indicating the volume term compensation flow rate deviation ΔF _HV is obtained. Output. The deviation calculation means 52 gives a signal indicating the flow rate compensation flow deviation ΔF _HF and a signal indicating the volume term compensation flow deviation ΔF _HV to the control amount calculation means 53.

The control amount calculation means 53 is a first calculation value (K _FG · ΔF _HF ) obtained by multiplying the flow rate compensation flow deviation ΔF _HF calculated by the deviation calculation means 52 by a flow rate gain K _FG that is a coefficient set in advance. A second calculation value (K _VG · ΔF _HV / s) obtained by multiplying a value ΔF _HV / s obtained by sequentially integrating the volume term compensation flow rate deviation ΔF _HV from the start of the supply of the support gas and a volume gain K _VG that is a coefficient set in advance. And the target pressure Pin related to the support pressure Pvent is calculated by adding the first calculation value (K _FG · ΔF _HF ) and the second calculation value (K _VG · ΔF _HV / s). Therefore, the equation indicating the target pressure Pin is expressed by the following equation (6).
Pin = K _FG · (F _HF − ^ F) + K _VG · (F _HV − ^ F) / s (6)

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 compensation flow rate generation unit 201, the deviation calculation unit 52, and the control amount calculation unit 53 described above have been individually described for easy understanding. However, the transfer function may be equivalently converted and arranged. Further, the estimation unit 51, the compensation flow rate generation unit 201, the deviation calculation unit 52, and the control amount calculation unit 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 in the following equation (7).
G (s) ₂₀ = Gc (s) · e ^{−Lc · s} (7)

Here, Gc (s) ₂₀ indicates 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.

Further, the following equation (8) shows an 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.
G (s) ₅₅ = Gm (s) · e ^{−Lm · s} (8)

Here, Gc (s) ₅₅ indicates a transfer function excluding the dead time element, and means a first-order lag element in the present embodiment. 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.

  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 about (1 + B) times the spontaneous breathing pressure Pmus according to the time change of the spontaneous breathing pressure Pmus. It is amplified by. 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 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.

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 _H. The compensation flow rate F _H 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. Therefore, the flow deviation ΔF _H 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 deviation ΔF _H 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. it can.

The entire system 14 including the ventilator 17 and the patient of the present invention 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, its transfer function G ( s) ₁₄ is represented by a predetermined arithmetic expression. In the ventilator 17, when the detected flow rate F after the inspiration start time T _onset is an input value and the pressure target value Pin is an output value, the transfer function G (s) ₁₇ of the ventilator ₁₇ is determined in advance. It can be obtained by an arithmetic expression.

FIG. 4 is a graph showing temporal changes in the spontaneous breathing pressure Pmus and the flow rate F of the assisting gas flowing through the airway. FIG. 4 (1) shows the time change of the spontaneous breathing pressure Pmus, and FIG. 4 (2) shows the time change of the flow rate F of the support gas. In FIG. 4 (1), the spontaneous breathing pressure Pmus is indicated by a solid line. In addition, the estimated spontaneous breathing pressure ^ Pmus (= R · F _HF + E · ∫F _HV ) estimated by the calculation is indicated by a broken line. Further, the flow-induced pressure value R · F _HF and the volume-induced pressure value E · ∫F _HV constituting the estimated spontaneous breathing pressure ^ Pmus are shown by a one-dot chain line and a two-dot chain line, respectively. In FIG. 4B, the detected flow rate F is indicated by a solid line. Further, the flow rate compensation flow rate F _HF is indicated by a one-dot chain line, and the volume term compensation flow rate F _HV is indicated by a two-dot chain line.

As shown in FIG. 4, the flow rate compensation flow rate F _HF and the volume term compensation flow rate F _HV are values obtained by adding compensation values H _F , H _V , and H _d to the detected flow rate F. Therefore, the compensation flow rates F _HF and F _HV are zero or more and can be values that rise with the passage of time from the intake start time T _onset . Even if there is residual pressure in the lung at the inspiration start time T _onset and the support gas is discharged from the lung, as shown in FIG. 4 (2), the compensation flow rate from the inspiration start time T _onset F _H can be generated. Thus, the estimated spontaneous breathing pressure ^ Pmus that increases from the inspiration start time T _onset can be calculated.

In the present embodiment, the compensation values H _F , H _V , and H _d are set to values that smoothly decrease with time. Accordingly, the compensation flow rate values F _HF and F _HV gradually decrease with time from the intake start time T _onset and approach the detected flow rate F. Thus the after the initial period W _H of the intake period of time, it is possible to suppress the influence of the compensation value H. Further, as shown in FIG. 4A, the similarity of the transient waveform between the spontaneous breathing pressure Pmus and the estimated spontaneous breathing pressure ^ Pmus can be increased. Thus, by bringing the estimated spontaneous breathing pressure ^ Pmus close to the actual spontaneous breathing pressure Pmus, the gas supply servomechanism can be controlled according to the estimated spontaneous breathing pressure Pmus close to the spontaneous breathing pressure Pmus, The burden can be reduced.

Further, since each compensation value H _F , H _V , H _d is set to an output value that has been subjected to incomplete differentiation, each compensation value H _F , H _V , H _d can be prevented from changing rapidly. . The compensation flow values F _HF and F _HV obtained from the compensation values H _F , H _V and H _d are also prevented from changing rapidly in a short time. As a result, the estimated spontaneous breathing pressure ^ Pmus can be increased smoothly, and the burden on the patient received during inhalation support can be reduced. Moreover, the settling time can be changed according to the detected flow rate F _onset at the intake start time T _onset by setting the compensation value H to a value to be subjected to the incomplete differentiation calculation process. This makes it possible to compensate more according to the patient's condition.

Further, in this embodiment, based on the flow rate _{F onset} of supporting gas flowing through the airway in the intake beginning _{T onset,} flow compensation value _{H F} and volume compensation value _{H V} is set. As a result, the compensation value H can be changed according to the patient's condition, and the compensation flow rate value F _H can be prevented from becoming excessive or too small. Inhalation support can be suitably performed until it flows toward the lungs. For example, even when the patient's breathing motion is in a state of dynamic hyperinflation, inhalation assistance can be performed while preventing the inhalation start delay of the ventilator.

In the present embodiment, the base compensation value _Hd is set based on the flow rate change delay from when the respiratory muscle of the patient starts the inspiration operation until the change in the flow rate of the support gas according to the inspiration operation occurs. . As a result, it is possible to prevent an error between the spontaneous breathing pressure Pmus and the estimated spontaneous breathing pressure ^ Pmus due to the flow rate change delay that is substantially constant regardless of the patient's condition.

For example, the flow rate adjustment gain K _FH and the volume adjustment gain K _VH are set so that the flow rate compensation value F _H becomes zero at the intake start time T _onset . In this embodiment, the volume term compensation flow rate F _HV are determined by the volume compensation value H _V, the volume compensation value H _V is set to 1. In contrast, the flow rate term compensation flow value F _HF are determined by the flow compensation value H _F together with the base compensation value H _d, in a state in which the flow compensation value H _F and base compensation value H _d is combined intake The flow rate compensation value F _H is set to be zero at the start time point T _onset . Thus flow compensation value H _F is preferably set to a value smaller than the volume compensation value H _V.

The transfer function G ₂₁₁ of the first waveform shaper 211 to determine the attenuation of the flow compensation value H _F and volume compensation value H _V is preferably determined according to the constant Tp when the patient's spontaneous breathing pressure Pmus . For example, when the time constant Tp is small, that is, when the rise time is short, the parameter of the transfer function _G211 is determined so as to attenuate quickly. When the time constant Tp is large, that is, when the rise time is long, the parameters of the transfer function _G211 are determined so as to attenuate slowly. As a result, the compensation flow rate value F _H can be increased smoothly with time in the intake period, and the influence of the compensation value H can be reduced. As an example, the parameters a1, a2 is set to the transfer function _{G 211} is set to a value proportional constant Tp time of the patient's spontaneous breathing pressure Pmus. In addition, it is preferable that the parameters a1 and a2 are obtained corresponding to the time constant Tp. As described above, the compensation value H is preferably set to a maximum value, an attenuation time, an attenuation amount per time, and the like based on a transient change in the spontaneous breathing pressure Pmus of the patient.

In the present embodiment, there is a base compensation value Hd that occurs regardless of the flow rate _Fonset of the assist gas flowing through the airway at the intake start time T _onset . This makes it possible to eliminate also the intake support delay flow F _onset of the supporting gas in the intake beginning T _onset occurs even if zero. In this case, the transfer function _{G 212} of the second waveform shaper 212 to determine the attenuation amount of the base compensating value _{H d,} when the flow rate _{F onset} of the supporting gas in the intake beginning _{T onset} is zero, the compensation flow rate value The value is set to a value that can prevent the compensation value from becoming zero and can reduce the influence of the compensation value. As a result, even if the lung state is not dynamic overexpansion, a delay in the inspiratory support operation can be suppressed.

In the present embodiment, the target pressure Pin compensated for the delay in the intake assist operation can be calculated by obtaining the target pressure value Pin using the flow rate deviation ΔF _H , and the patient's patient at the intake start time T _onset can be calculated. Even when the flow rate F of the assist gas flowing toward the lung is less than or equal to zero, the control of the gas supply servo mechanism can be started so as to supply the assist gas with a pressure obtained by proportionally amplifying the spontaneous breathing pressure Pmus to the patient. . In addition, the patient's inspiratory period and the inspiratory support period by the ventilator can be synchronized.

Further, by using the flow rate deviation ΔF _H , the spontaneous breathing pressure Pmus can be accurately estimated. Furthermore, the robust stability of the control system including the patient and the ventilator can be improved, and even if the flow gain K _FG and the volume gain K _VG deviate from the values indicating the patient's condition, the control system is unstable. Can be prevented.

Further, according to the present embodiment, different compensation flow rate values F _HF and F _HV are used for the flow rate setting value (K _FG · ΔF _H ) and the volume setting value (K _VG · ΔV _H ). As a result, it is possible to increase the parameters relating to the adjustment of the inspiration support, and to flexibly adjust the respiration operation so that the estimated spontaneous respiration pressure ^ Pmus approaches the spontaneous respiration pressure Pmus. For example, by appropriately setting the two compensation values F _HF and F _HV , it is possible to prevent the support for the intake operation from becoming excessive or too small, and to facilitate the support for the desired intake operation. .

FIG. 5 is a diagram illustrating a simulation result when the respiration rate per unit time is small. FIG. 5 (1) shows the time change of the spontaneous breathing pressure Pmus simulating the spontaneous breathing of the patient. FIG. 5 (2) shows the time change of the estimated spontaneous breathing pressure ^ Pmus calculated based on the flow rate F of the support gas. FIG. 5 (3) shows the change over time in the flow rate F of the assisting gas flowing through the patient's airway when the estimated spontaneous breathing pressure ^ Pmus is amplified at a predetermined amplification factor. In the simulation result shown in FIG. 5, the respiration rate per minute of the patient is set to 12 times, and the maximum value of the spontaneous breathing pressure Pmus is set to 3 cmH ₂ O. Moreover, simulation was performed using the parameter values shown in Table 2, Table 3, and Table 4 below, and the estimated spontaneous breathing pressure Pmus was obtained based on the compensation flow value F _H. The simulation result of FIG. 5 corresponds to the simulation result of FIG. Figure 18 is that the seeking estimated spontaneous respiration pressure ^ Pmus using a detection flow F, Figure 5 is that obtaining the estimated spontaneous respiration pressure ^ Pmus using a compensated flow F _H deviation [Delta] F _H determined from Is different.

As shown in FIG. 5, the detected flow _{F onset} in the intake beginning _{T onset} is the case was almost zero, the first compensation value H1 generated based on the detected flow rate _{F onset} in the intake beginning _{T onset,} i.e. flow compensation value H _F and volume compensation value H _V becomes almost zero. On the other hand, the base compensation value H _d determined based on the specified value H0 is determined regardless of the detected flow rate F _onset at the intake start time T _onset . Therefore even if it was detected flow rate F _onset substantially zero in the intake beginning T _onset as shown in FIG. 5, the respiratory muscles of the patient starts to inspiration, supporting gas in accordance with the intake operation The flow rate change delay until the flow rate change occurs can be prevented. In the present embodiment, as shown in FIG. 5, the estimated spontaneous breathing pressure ^ Pmus can be generated almost simultaneously with the start of the spontaneous breathing pressure Pmus.

As a result, in the present embodiment, the rising delay of the estimated spontaneous pressure ^ Pmus with respect to the intake start time T _onset can be made substantially zero. Accordingly, as shown in FIG. 18, when the estimated spontaneous pressure ^ Pmus is estimated based on the detected flow rate, a delay Td of 100 msec is generated, whereas in the present embodiment, a substantial difference with respect to the intake start time T _{onset is} generated. The delay can be prevented, the inspiration support operation can be started, and the burden on the patient can be reduced.

The base compensation value H _d Since the decay with time, after a lapse of the initial period W _H is compensated flow F _H, the influence of the base compensation value H _d becomes smaller. Therefore, as shown in FIGS. 5 (1) and 5 (2), the waveforms of the spontaneous respiratory pressure Pmus and the estimated spontaneous respiratory pressure ^ Pmus can be made similar. For example, the maximum value of the spontaneous breathing pressure Pmus and the maximum value of the estimated spontaneous breathing pressure ^ Pmus can both be 2.8 cmH ₂ O. As a result, the inspiration support operation can be performed with an amplification factor having a high correlation with the spontaneous breathing pressure Pmus.

FIG. 6 is a diagram showing a simulation result simulating a state in which the lung is in dynamic overexpansion. FIG. 6 (1) shows the time change of the spontaneous breathing pressure Pmus simulating the spontaneous breathing of the patient. FIG. 6 (2) shows the time change of the estimated spontaneous breathing pressure ^ Pmus calculated based on the flow rate F of the support gas. FIG. 6 (3) shows the time change of the flow rate F of the assisting gas flowing through the patient's airway when the estimated spontaneous breathing pressure ^ Pmus is amplified at a predetermined amplification factor. In the simulation result shown in FIG. 6, the respiration rate per minute of the patient is set to 30 times, and the maximum value of the spontaneous breathing pressure Pmus is set to 10 cmH ₂ O. Moreover, simulation was performed using the parameter values shown in Tables 2 to 4, and the estimated spontaneous respiratory pressure Pmus was obtained based on the compensation flow rate F _H. The simulation result of FIG. 6 corresponds to the simulation result of FIG. Figure 19 is that the seeking estimated spontaneous respiration pressure ^ Pmus using a detection flow F, 6 that obtains the estimated spontaneous respiration pressure ^ Pmus using a compensated flow F _H deviation [Delta] F _H determined from Is different.

As shown in FIG. 6, in a state where the patient's respiration rate per minute is large and the compliance C, which is the reciprocal of the airway pressure R and elastance, is large, the patient's lung condition tends to be dynamic hyperinflation. If dynamic hyperinflation occurs, there is residual pressure in the lungs at the intake beginning _{T onset,} the detected flow _{F onset} in the intake beginning _{T onset,} but negative, that is, the direction to assist gas discharged from the lungs Flows. In this case, the absolute value of the detected flow rate _{F onset} in the intake beginning _{T onset | F onset |} depending on the flow rate compensation value _{H F} and volume compensation value _{H V} is generated. Flow compensation value H _F and volume compensation value H _V is the intake beginning T _onset, the greater the flow rate of the assist gas flowing in the discharge direction to be discharged from the lungs is set to a large value. On the other hand, the base compensation value H _d is determined regardless of the detected flow rate F _onset at the intake start time T _onset .

In the present embodiment, even when the flow aid gas in the discharge direction at the intake beginning T _onset, by flow compensation value H _F and volume compensation value H _V is set, with respect to the intake start time T _onset Thus, the rising delay of the estimated spontaneous pressure ^ Pmus can be made almost zero. Accordingly, as shown in FIG. 19, when the estimated spontaneous pressure ^ Pmus is estimated based on the detected flow rate F, the substantial intake assist time t _oneff has a delay Td of about 800 msec with respect to the intake start time T _onset . On the other hand, in the present embodiment, the substantial inspiration assist time t _oneff has a very small delay with respect to the inhalation start time T _onset , so that the inspiration support operation can be started and the burden on the patient can be reduced. .

The flow compensation value _{H F,} volume compensation value _{H V} and base compensation value _{H d} Since the decay with time, with time, compensated flow _{F H,} the influence of the compensation value _{H F, H V, H d} Less. Therefore, as shown in FIGS. 6 (1) and 6 (2), the waveforms of the spontaneous breathing pressure Pmus and the estimated spontaneous breathing pressure ^ Pmus can be made similar. For example, the maximum value of the spontaneous breathing pressure Pmus is 9.5 cmH ₂ O, and the maximum value of the estimated spontaneous breathing pressure ^ Pmus can be about 10 cmH ₂ O, and the difference is small. As a result, the inspiration support operation can be performed at a high amplification factor similar to the spontaneous breathing pressure Pmus.

  FIG. 7 is a diagram illustrating a simulation result simulating a state in which the lung is in dynamic overexpansion. FIG. 7 (1) shows the time change of the spontaneous breathing pressure Pmus simulating the spontaneous breathing of the patient. FIG. 7 (2) shows the time change of the estimated spontaneous breathing pressure ^ Pmus calculated based on the flow rate F of the support gas. FIG. 7 (3) shows the change over time in the flow rate F of the assisting gas flowing through the patient's airway when the estimated spontaneous breathing pressure ^ Pmus is amplified at a predetermined amplification factor.

In the simulation result shown in FIG. 7, the respiration rate per minute of the patient is set to 30 times, and the maximum value of the spontaneous breathing pressure Pmus is set to 5 cmH ₂ O. Moreover, simulation was performed using the parameter values shown in Tables 2 to 4, and the estimated spontaneous breathing pressure Pmus was obtained based on the compensation flow rate F _H. The simulation result of FIG. 7 corresponds to the simulation result of FIG. Figure 7 is that the maximum value of the spontaneous respiration pressure Pmus is set to 5 cmH ₂ O, 6 maximum value of the spontaneous respiration pressure Pmus is set to 10 cm H ₂ O. As shown in FIG. 7, even if the maximum value of the spontaneous breathing pressure Pmus is set to 5 cmH ₂ O, the estimated spontaneous breathing pressure similar to the spontaneous breathing pressure Pmus can be calculated without changing the parameter of the control element. .

  As shown in FIGS. 6 and 7, even if the patient's condition changes and the maximum value of the spontaneous breathing pressure Pmus changes, in this embodiment, the estimated spontaneous breathing pressure similar to the patient's spontaneous breathing pressure Pmus is used. ^ Pmus can be calculated respectively.

Thus, in the present embodiment, the gas supply servo mechanism is controlled using the compensation flow rate F _H slightly changed with respect to the detected flow rate F. The difference between the compensated flow rate value F _H and the detected flow rate F decreases with time, and the compensated flow rate value F _H approaches the detected flow rate F with time. By using the compensated flow rate value F _H instead of the detected flow rate F given as an input value, it is not necessary to change the feedback gain and control elements of the entire system including the patient and the ventilator, and inspiratory support delay is prevented. be able to. Therefore, it is not necessary to change the feedback gain and control elements of the entire system, the stability of the entire system can be maintained, and even if the compensation flow rate value F _H is used, the entire system becomes unstable. Can be prevented. That is, robust stability and quick response can be maintained.

  FIG. 8 is a diagram illustrating a simulation result simulating a state in which the lung is in dynamic hyperinflation. FIG. 8 (1) shows the time change of the spontaneous breathing pressure Pmus simulating the spontaneous breathing of the patient. FIG. 8 (2) shows the time change of the estimated spontaneous breathing pressure ^ Pmus calculated based on the flow rate F of the support gas. FIG. 8 (3) shows the change over time in the flow rate F of the assisting gas flowing through the patient's airway when the estimated spontaneous breathing pressure ^ Pmus is amplified at a predetermined amplification factor.

Simulation results shown in FIG. 8, the number of breaths per minute of the patient is set to 30 times, simulation is performed using the parameter values shown in Table 1 to obtain the estimated spontaneous breathing pressure Pmus based on compensated flow F _H It was. The simulation result shown in FIG. 8 shows that the flow rate adjustment gain K _FH is set to a value equal to the volume adjustment gain K _VH , that is, K _FH = K _VH = 1. A simulation was performed, and an estimated spontaneous breathing pressure Pmus was obtained based on the compensation flow rate value F _H.

As shown in FIG. 8, when the parameters of the present embodiment are used, when the flow rate adjustment gain K _FH is made the same as the volume adjustment gain K _VH , the rise of the estimated spontaneous respiration pressure ^ Pmus becomes the spontaneous respiration pressure Pmus. Compared with the simulation result shown in FIG. 6, the similarity between the spontaneous breathing pressure Pmus and the estimated spontaneous breathing pressure ^ Pmus is faster. In the present embodiment, as shown in FIG. 2, the flow rate adjustment gain K _FH and the volume adjustment gain K _VH can be individually adjusted, so that the flow rate adjustment gain K _FH and the volume adjustment gain K _VH are optimal. By setting to a small value, the similarity between the spontaneous respiratory pressure Pmus and the estimated spontaneous respiratory pressure ^ Pmus can be increased as shown in FIG.

  FIG. 9 is a diagram illustrating a simulation result simulating a state in which the lung is in dynamic hyperinflation. FIG. 9 (1) shows the time change of the spontaneous breathing pressure Pmus simulating the spontaneous breathing of the patient. FIG. 9 (2) shows the time change of the estimated spontaneous breathing pressure ^ Pmus calculated based on the flow rate F of the support gas. FIG. 9 (3) shows the change over time of the flow rate F of the assisting gas flowing through the patient's airway when the estimated spontaneous breathing pressure ^ Pmus is amplified at a predetermined amplification factor.

In the simulation results shown in FIG. 9, the respiratory rate per minute of the patient was set to 30 times, and the estimated spontaneous respiratory pressure Pmus was obtained based on the compensation flow rate F _H. The simulation results shown in FIG. 9 show that the time constant Tp of the spontaneous breathing pressure Pmus is 0.1, and the other parameters are simulated using the parameter values shown in Tables 2 to 4 to obtain the compensation flow rate value F _H. Based on this, the estimated spontaneous breathing pressure Pmus was obtained. That is, the case where the spontaneous breathing pressure Pmus rises in a shorter time than the simulation result shown in FIG.

In this case, with the adjustment gain (K _FH = 0.5, K _VH = 1) for each compensation value shown in Table 4, the estimated spontaneous breathing pressure ^ Pms is slightly larger than the spontaneous breathing pressure Pmus. At the same time, the rising is delayed with respect to the spontaneous breathing pressure Pmus. If the maximum value of the estimated spontaneous breathing pressure ^ Pmus is increased, by reducing the volume adjustment gain K _VH, the maximum value of the estimated spontaneous breathing pressure ^ Pms, it can be brought close to the maximum value of the spontaneous respiration pressure Pmus . In the case where the rising of the estimated spontaneous breathing pressure ^ Pmus is delayed, by increasing the flow rate adjusting gain K _FH, it is possible to prevent delay in the rise of the estimated spontaneous breathing pressure ^ Pms. In the present embodiment, the flow rate adjustment gain K _FH and the volume adjustment gain K _VH can be set individually, so that a compensation value corresponding to the patient's breathing state can be generated flexibly.

  FIG. 10 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 airway pressure detection means 61, an input means 39, a display means 40, and a servo amplifier 48 which is an amplifier circuit.

  The flow rate detecting means 50 converts the flow rate of the gas flowing through the inspiratory line 25 and the expiratory line 26 of the gas supply mechanism 20 into an electric signal, and gives the electric signal to the control device main body 33. The airway pressure detection means 61 converts the pressure of the gas flowing through the inspiratory line 25 and the expiratory line 26 of the gas supply mechanism 20 into an electric signal, and gives the electric signal to the control device main body 33.

  The ventilator main body 17 a includes a mounting member 24, a pump 22, a pump actuator 31, an exhalation valve 27, an inspiratory conduit 25, an expiratory conduit 26, and an expiratory valve actuator 32. In the present embodiment, the gas supply mechanism 20 that includes the pump 22 and the pump actuator 31 and can adjust the pressure of the discharged assist gas is configured.

  The attachment member 24 is attached to the patient to supply gas to the patient's airway, and an attachment hole communicating with the patient's airway is formed. The inspiratory duct 25 is a duct connecting the pump 22 and the mounting member 24, and guides the gas supplied from the pump 22 to the patient's airway. The exhalation duct 26 is a duct that connects the mounting member 24 and the discharge place 23 outside the apparatus, and guides the gas flowing out from the patient's airway to the discharge place 23 outside the apparatus.

  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, a nurse, or an administrator 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 and the estimated spontaneous breathing pressure ^ Pmus. 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 such as the estimated spontaneous breathing pressure ^ Pmus on the display screen. The servo amplifier 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 storage unit 103, and a storage unit 104. The interface 101 receives signals from the flow rate detection means 50 and the airway pressure detection means 61 to be connected, and gives the signals 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 compensation flow rate generation unit 201, the deviation The calculating means 52, the control amount calculating means 53, and the patient respiratory state monitoring means 200 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. 11 is a flowchart showing the operation of the calculation unit 102 of the control device main body 33. First, in step s0, the calculation unit 102 applies the estimated airway resistance ^ R, the estimated elastance ^ E, the transfer function of the gas supply mechanism 20, the transfer function of the delay compensation unit 55, the flow rate gain K _FG , and the volume gain to the control device body 33. Parameters such as K _VG , second and second waveform shapers 211, flow rate adjustment gain K _FH , volume adjustment gain K _{VH, and the} like are input, and target pressure Pin, estimated flow rate ^ F, and compensation value H can be calculated. When calculation preparation is possible, the process proceeds to step s1.

In step s1, the calculation unit 102 determines whether or not it is the inhalation start time T _onset by executing a state monitoring program that realizes the patient state monitoring unit 200. Specifically, the calculation unit 102 determines whether or not the intake start time T _onset has been reached according to a predetermined calculation formula based on signals given from the flow rate detection means 50 and the airway pressure detection means 61. When it is determined that the patient's breathing state has reached the inspiration start time point T _onset by the inhalation start time point detection step, the process proceeds to step s2.

In step s <b> 2, the calculation unit 102 starts generating the compensation flow rate H by executing a compensation flow rate generation program that realizes the compensation flow rate generation unit 201. Specifically, as described with respect to FIG. 2, the flow rate compensation value H _F, starts generating the volume compensation values H _V and base compensation value H _d. When the compensation value calculation process for sequentially calculating the compensation values H _F , H _V , and H _d is started in this way, the process proceeds to step s3.

In step s3, the calculation unit 102 executes the compensation flow rate generation program for realizing the compensation flow rate generation unit 201, thereby sequentially executing the compensation values H _F , H _V , and H _d that are sequentially calculated. _{F 1} , H _V , H _d and the first detected flow rate F among the detected flow rates F that are sequentially detected are added to generate a compensation flow rate value. In this embodiment as described above, to produce a flow compensation value H _F, and the base compensation value H _d, the flow rate term compensation flow value F _HF which is the sum of the detected flow F. The generating and the volume compensation value H _V, the detected flow F and the volume term compensation flow rate F _HV which is the sum of. When the compensation flow value generation process for generating these compensation flow values F _HF and F _HV is completed, the process proceeds to step s4.

In step s4, the calculation unit 102 executes the deviation calculation program that realizes the deviation calculation means 52, and thereby the estimated flow rate ^ F obtained from the previously calculated target pressure Pin and the compensation flow rate value generated in step s3. A deviation ΔF _H between F _HF and F _HV is calculated. Specifically, a flow rate compensation deviation ΔF _HF that is a deviation between the flow rate compensation flow value F _HF and the estimated flow rate F is calculated. Further, a volume term compensation deviation ΔF _HV that is a deviation between the volume term compensation flow rate value F _HV and the estimated flow rate ＦF is calculated. When the compensation flow rate deviation is obtained in this way, the process proceeds to step s5.

In step s5, the calculation unit 102 executes a control amount calculation program that realizes the control amount calculation means 53, and thereby, based on the flow rate compensation deviation ΔF _HF and the volume term compensation deviation ΔF _HV , the gas supply servo mechanism. To calculate the target pressure Pin of the support gas to be discharged to the intake pipe. When the target pressure Pin is calculated as the control amount in this way, a signal representing the target pressure Pin is given to the gas supply mechanism 20, and the process proceeds to step s6.

  In step s6, the calculation unit 102 executes an estimation program for realizing the estimation unit 51, thereby performing an 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. The estimated flow rate {circumflex over (F)} that will be supplied to the patient when Pmus does not exist is calculated, and the process proceeds to step s7.

In step s7, the calculation unit 102 determines whether it is the inhalation end time T _offset by executing a state monitoring program for realizing the patient state monitoring unit 200. Specifically, the calculation unit 102 determines whether or not the intake end time T _offset has been reached according to a predetermined calculation formula based on signals given from the flow rate detection means 50 and the airway pressure detection means 61. If it is determined in the inhalation end time detection step that the patient's breathing state has reached the inspiration end time T _offset as described above, the process proceeds to step s8.

  In step s8, the calculation unit 102 performs a control operation of the servo supply mechanism during expiration. For example, a command is given to the servo supply mechanism so that the airway pressure becomes a preset pressure at expiration. For example, in order to prevent collapse of the patient's lungs, the exhalation set pressure is maintained to be higher than the positive end-expiratory pressure (abbreviation: PEEP pressure), which is higher than the atmospheric pressure, in step s1. Return to.

If it is determined in step s1 that the patient's breathing state has not reached the inspiration start time point T _onset , the process proceeds to step s8 to perform an expiration control operation. When the calculation unit 102 determines that a predetermined end condition is satisfied while performing the operations of steps s1 to s7, the calculation unit 102 ends the control operation of the gas supply servo mechanism. For example, when it is determined that a predetermined end condition is satisfied by giving an end command by the input means 39, the control operation is ended.

  As described above, the flow rate estimation unit 51, the compensation flow rate generation unit 201, the deviation calculation unit 52, the control amount calculation unit 53, and the patient respiratory state monitoring unit 200 may be realized by executing software predetermined by the computer. . A program for realizing the above-described method and a storage medium storing the program are also included in the embodiment of the present invention. Further, the above-described control device alone and a ventilator including the control device and the ventilator main body are also included in the present invention.

  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. 11, or may be a ventilator that supplies support gas via a pipe.

Further, in this embodiment, it takes in the detected flow F _onset in the intake start time T _onset every inspiration, generating a compensation value for each inspiration. As a result, a compensation flow rate corresponding to a change in the respiratory state can be generated, and even if the patient's state changes, it is possible to prevent inspiration support delay more accurately.

FIG. 12 is a graph showing the principle for _obtaining the intake start time T _onset . FIG. 12 (1) is a graph showing the time change of the detected flow rate F, and FIG. 12 (2) is a graph showing the time change of the differential value Fdt obtained by differentiating the time change of the detected flow rate F.

When the patient starts the inhalation operation, the differential value Fdt, which is a time change in the flow rate of the support gas, increases beyond the predetermined inhalation start threshold value X1. This is a phenomenon that occurs both when the support gas is flowing in the direction toward the lungs and when the support gas is flowing in the direction of being discharged from the lungs. In the present embodiment, the time point when the differential value Fdt of the support gas flow rate F exceeds the intake start threshold value X1 is detected as the intake start time point T _onset based on the time change of the differential value Fdt of the support gas flow rate F. To do. As a result, the inspiration start time T _onset can be determined even when the support gas flows in the direction in which the assist gas is discharged from the lungs at the inspiration start time T _onset .

In the intake start time detection step of step s1 shown in FIG. 11, a time when the deviation shown below exceeds a predetermined intake start threshold value X1 may be determined as the intake start time T _onset . For example, the deviation is (1) a difference between two old and new detected values of two detected flow rates F detected at any time, and (2) an estimated value obtained by extrapolating a predicted time lapse value of the detected flow rate F and an actual value. The difference between the values and (3) the time change rate of the detected flow rate F may be calculated as needed, and any of the time change rate differences for each calculation may be used.

In addition to the detected flow rate F, the inspiration start time T _onset may be obtained based on a differential value that is a temporal change amount of the spontaneous breath estimated value represented by (＾ R · F + ＾ E · ∫F). Moreover, you may use the technique disclosed by the international patent application (WO 2004/002561 A2) mentioned above.

Similarly, the time point when the change in the flow rate of the support gas or the differential value obtained by differentiating the estimated spontaneous breathing value falls below the predetermined inhalation end threshold value X2 can be detected as the inhalation end point _Toffset .

The detection of the intake start time _{T onset} by performing for periods there will be air start time _{T onset,} it can be prevented detection error of the intake air start time _{T onset.} Further, a mask 300 is set for the patient to exclude a change in the differential value of the detected flow rate F caused by disturbances such as coughing and sneezing from the detection determination of the inhalation start time point T _onset . As a result, the intake start time T _onset can be accurately determined.

Also in the same manner, the detection of the intake end _{T Offset} by performing for periods there will be an intake end _{T offset,} it is possible to prevent the detection error of the intake end _{T offset.} In addition, a mask is set for the patient to exclude a change in the differential value of the detected flow rate F caused by disturbance such as coughing and sneezing from the detection determination of the inhalation end time T _offset . As a result, the intake end time T _offset can be accurately determined.

  In addition to obtaining based on the detected flow rate, the patient respiratory state monitoring unit 200 may be realized by using a detection unit that detects a change state that changes when the patient starts inspiration. For example, the patient's inhalation start time may be obtained by causing a pressure sensor or a position sensor for detecting the movement of the diaphragm to enter the lung using a catheter and directly obtaining the pressure in the lung. In this way, the patient respiratory state monitoring unit 200 may be realized using a detection unit that can enter the lung.

FIG. 13 is a flowchart illustrating the operation of the calculation unit 102 of the control device main body 33 according to the modified example of the first embodiment. As a modification of the first embodiment of the present invention, the intake trigger time T _ontrg is used together with the intake start time T _onset to determine the start time of the intake assist operation. Further, the expiration trigger time T _offtrg is used together with the inspiration start time T _{onset for} the determination of the end time of the inspiration support operation. The operation of the operation unit 102 according to the modified example has steps similar to those in the flowchart of the first embodiment shown in FIG. 11, and description of similar steps is omitted.

First, in step a0, the same operation as in step s0 described above is performed, and the process proceeds to step a1. In step a1, the intake start time T _onset is detected in the same manner as in step s1, and if detected, the process proceeds to step a3. If not detected, the process proceeds to step a2. In step a2, it is determined whether or not the flow rate F of the assisting gas flowing through the airway exceeds a predetermined intake trigger set value F _ontrg that is equal to or greater than 0, and the time when it exceeds the set value F _ontrg is determined as the intake trigger time T _ontrg . . For example, the intake trigger set value is set to 1 liter / minute. If it is determined in step a2 that the intake trigger time T _ontrg has been reached, the process proceeds to step a3.

Steps a3 to a7 perform the same operations as steps s2 to s6 shown in FIG. In step a7, when the calculation unit 102 estimates the flow rate of the support gas, the calculation unit 102 proceeds to step a8. In step a8, the arithmetic unit 102 determines whether or not the intake end time T _offset is reached. If it is not determined that the patient's breathing state has reached the inhalation end time point T _offset , the process proceeds to step a9. In step a9, the calculation unit 102 determines whether or not the flow rate F of the assisting gas flowing through the airway has decreased _{below a} predetermined exhalation trigger set value F _offtrg, and determines when the exhalation trigger set value F _offtrg has decreased. _Judge as time T _offtrg . For example, the expiration trigger set value is set to a value that is 10% of the maximum value of the detected flow rate F. If it is determined in step a9 that the expiration trigger time point _Tofftrg has been reached, the process proceeds to step a10.

In step a10, the same operation as in step s8 is performed. When the operation of step a10 is completed, the process returns to step a1. Further, when the intake start time T _onset is detected in step a1, the process proceeds to step a3. If the intake trigger time T _offset is not detected in step a2, the process proceeds to step a10. In step a8, when the intake end time T _offset is detected, the process proceeds to step a10. If the expiration trigger time point T _offset is not detected in step a9, the process returns to step a4.

As described above, when the intake trigger time T _ontrg is detected before the intake start time T _onset is detected, the flow rate F of the support gas flowing through the airway is detected in the control amount calculation step, and the detected flow rate F of the support gas is detected. Based on the above, the target pressure Pin of the assist gas discharged from the gas supply servo mechanism is calculated. If it is determined that either the end of inspiration T _offset or the expiration trigger time T _offtrg has been reached, the control of the gas supply servo mechanism relating to the patient's inspiration support is ended.

As a result, the support for the intake operation can be started from either the intake start time T _onset or the intake trigger time T _ontrg detected at an earlier timing. Further, the support of the inhalation operation can be ended from the time point detected at an earlier timing, either the inspiration end time point T _offset or the expiration trigger time point T _offtrg . Therefore, even if the inhalation start time T _onset or the inhalation end time T _offset cannot be detected, the start time of the inspiration support operation can be determined reliably, and the safety of the ventilator can be improved.

FIG. 14 is a block diagram specifically showing the entire system 14 of the second embodiment of the present invention. The control device main body in the second embodiment shown in FIG. 14 has a flow rate setting value (K _FG · F _H ) obtained by multiplying a compensation flow rate value F _H by a predetermined flow rate gain K _FG and an integral value of the compensation flow rate value F _H. A total value ((K _FG · F _H ) + (K _VG · V _H )) obtained by adding a volume setting value (K _VG · V _H ) obtained by multiplying V _H by a predetermined volume gain K _VG is the target pressure Pin. Calculate as In this case, the estimation unit 51 and the deviation calculation unit 52 are removed from the ventilator shown in FIG. 1, and the configuration of the control amount calculation unit 53 is different. The control amount calculation means 53 obtains the flow rate compensation flow rate F _HF and the volume term compensation flow rate F _HV as in the first embodiment.

The control amount calculation means 53 is configured to compensate the first term value (K _FG · F _HF ) obtained by multiplying the flow rate term compensation flow rate F _HF by a flow rate gain K _FG that is a coefficient set in advance, and the volume term compensation from the start of the supply of the support gas. A second calculated value (K _VG · F _HV / s) obtained by multiplying a value F _HV / s obtained by sequentially integrating the compensation flow rate F _HV with a volume gain K _VG that is a preset coefficient is obtained, and a first calculated value ( K _FG · ΔF _HF ) and the second calculated value (K _VG · ΔF _HV / s) are added to calculate a target pressure Pin related to the support pressure Pvent. Therefore, the equation indicating the target pressure Pin is expressed by the following equation (9).
Pin = K _FG · F _HF + K _VG · F _HV / s (9)

Here, the symbols shown in the above formula have the same meaning. For example, the flow gain K _FG is set to a value (^ R · α _FG ) obtained by multiplying the estimated airway resistance ^ R by a predetermined flow amplification gain α _FG , and the volume gain K _VG is the estimated lung elastance ^ E. _Is multiplied by a predetermined volume amplification gain α _VG (^ E · α _VG ). When the flow rate amplification gain α _FG and the volume amplification gain α _VG are set to the same value, they are simply referred to as amplification gain α. Further, A represents the amplification gain α when ^ R = R and ^ E = E. In this case, 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 about 1 / (1 of the spontaneous breathing pressure Pmus according to the time change of the spontaneous breathing pressure Pmus. -A) Amplification is performed at a double amplification factor.

In the embodiment shown in FIG. 14, the entire system 14 including the ventilator 17 of the present invention and the patient takes the spontaneous breathing pressure Pmus as an input value, and outputs an added value of the spontaneous breathing pressure Pmus and the support pressure Pvent as an output value. Then, the transfer function G (s) ₁₄ is expressed by a predetermined formula.

In the present embodiment, by calculating the target pressure value Pin using the compensation flow rate value F _H , it is possible to calculate the target pressure Pin that compensates for the delay in the intake assist operation, and the patient's patient at the intake start time T _onset can be calculated. Even if the flow rate F _{onset of the} assist gas flowing toward the lung is less than or equal to zero, the gas supply servo is supplied so that the assist gas having a pressure proportionally amplified to the spontaneous breathing pressure Pmus corresponding to the patient's breathing effort is supplied to the patient. Control of the mechanism can be started. In addition, the patient's inspiratory period and the inspiratory support period by the ventilator can be synchronized.

  FIG. 15 is a flowchart showing the operation of the calculation unit 102 of the control device main body 33 in the second embodiment. The operation of the calculation unit 102 of the second embodiment has steps similar to the flowchart of the first embodiment shown in FIG. 11, and the description of the same steps is omitted.

First, in step b0, the same operation as in step s0 described above is performed, and the process proceeds to step b1. In step b1, the intake start time T _onset is detected in the same manner as in step s1, and if detected, the process proceeds to step b2. Steps b2 and b3 perform the same operations as steps s3 and s4 shown in FIG. When the compensation flow rate is generated in step b3, the process proceeds to step b4.

  In step b4, the target pressure Pin is estimated according to the arithmetic expression described above. When the target pressure Pin is calculated as a control amount as described above, a signal representing the target pressure Pin is given to the gas supply mechanism 20, and the process proceeds to step b5. In steps b5 and b6, the same operations as in steps s7 and s8 are performed. Even in the second embodiment or the like having a control system different from that in the first embodiment, when the control amount is determined based on the detected flow rate F, the same effect as in the first embodiment can be obtained. In this way, the present invention can be used for all control methods for controlling a ventilator based on the detected flow rate, and inspiratory assistance is achieved by controlling the ventilator using the compensation flow rate described above instead of the detected flow rate. It is possible to prevent a delay in intake support by preventing excess.

FIG. 16 is a diagram illustrating another example of the correction value H. In the present embodiment described above, the compensation value H is determined from the result of the incomplete differentiation operation, but is not limited to this. Compensation value H has a large value in the initial period W _H that is set shorter than the inhalation period period W _in the patient, the smaller the value after a lapse of the initial period W _H, its waveform, FIG. 16 (1) As shown in FIG. 16 (4), any waveform may be used.

For example, as shown in FIG. 16A, the compensation value H may be a so-called sawtooth waveform in which a time change attenuates in a linear function with the passage of time from the intake start time T _onset . Further, as shown in FIG. 16 (2), the amount of attenuation attenuated from the intake start time T _onset does not have to be constant, and attenuates over time according to other functions such as quadratic function changes and exponential function changes. May be. The compensation value H, it is preferable that the maximum value in the intake beginning T _onset, as shown in FIG. 16 (3), may be a maximum value when the time from the intake start time T _onset has elapsed . As shown in FIG. 16 (4), the compensation value H may have a rectangular waveform in its time-varying waveform. These compensation values H indicates a large value in an initial period W _H that is set shorter than the inhalation period period W _in the patient, a value lower than past the initial period W _H to the magnitude of the initial period W _H Set to Each compensation value H shown in FIGS. 16 (1) to 16 (4) shows a value of zero when the initial period _WH has elapsed, but is not limited thereto. For example, the compensation value H gradually decreases with time, and may be set to a value equal to or greater than zero even after the initial period _WH has elapsed.

The compensation value H, as shown in FIG. 16 (1) to 16 (3), the maximum value of predetermined intake start time _{T onset,} decreasing values smoothly with the elapsed time from the intake start time _{T onset} It is preferable to set to. As a result, it is possible to prevent the pressure value of the assist gas supplied toward the patient's lung from changing rapidly in a short time, and to reduce the burden on the patient who is inhaled. Moreover, you may use the compensation value H shown other than FIG. 16 (1)-(4).

FIG. 17 is a block diagram showing a ventilator according to another embodiment of the present invention. As shown in FIG. 17, the ventilator intake start through the input means from the patient respiratory status monitoring device 200 for detecting an intake start time _{T onset} and the intake end _{T offset} the patient time _{T onset} and the intake end point T A signal indicating the _offset may be provided. As a result, the same effects as those of the first and second embodiments can be obtained.

The above-described embodiment is merely an example of the invention, and the configuration can be changed within the scope of the invention. For example, the block diagrams shown in FIGS. 2 and 14 are merely examples of the present invention, and the configuration may be changed if the same effect can be obtained. For example, in this embodiment, the base compensating value H _d, the flow compensation value H _F, was used and the volume compensation value H _V, may be used any one. For example the base compensation value H _d alone, or a flow compensation value H _F, may be used only as a volume compensation value H _V as the compensation value. In the present embodiment, the flow rate adjustment gain K _FH and the volume adjustment gain K _VH are used separately, but a common adjustment gain may be used. In the above-described embodiment, the case of using the PAV method has been described. However, the present invention can be applied to any control method other than the PAV method and controlling the gas supply servo mechanism based on the detected flow rate.

  In addition, the configuration other than the compensation flow rate generation can be changed as appropriate. For example, in this embodiment, the transfer function of the delay compensation unit 55 may be a function that simulates the transfer function of the gas supply mechanism, or a transfer function different from the transfer function of the gas supply mechanism may be used.

  Preliminary values may be used for the estimated airway resistance ^ R and the estimated elastance ^ E, but the airway resistance R and the elastance E are measured and determined based on the measured values. Also good. Further, the airway resistance R and the elastance E obtained based on the technique disclosed in JP-T-11-502755 may be input to the control device.

  Each parameter shown in Tables 2 to 4 is preferably configured to be changeable by input means. Further, the current parameter value, estimated spontaneous breathing pressure ^ Pmus, detected flow rate, compensated flow rate, etc. may be displayed on the display means.

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. 4 is a timing chart for explaining signals generated by a compensation flow rate generation unit 201. It is a graph which shows the time change of the spontaneous-respiration pressure Pmus and the flow volume F of the assistance gas which flows through an airway. It is a figure which shows the simulation result in case there are few respiration rates per unit time. It is a figure which shows the simulation result which simulated the state from which a lung becomes dynamic hyperinflation. It is a figure which shows the simulation result which simulated the state from which a lung becomes dynamic hyperinflation. It is a figure which shows the simulation result which simulated the state from which a lung becomes dynamic hyperinflation. It is a figure which shows the simulation result which simulated the state from which a lung becomes dynamic hyperinflation. 2 is a block diagram showing an example of a ventilator 17. FIG. 4 is a flowchart showing the operation of the calculation unit 102 of the control device body 33. It is a graph which shows the principle for calculating _| requiring the inspiration start time _Tonset . It is a flowchart which shows operation | movement of the calculating part 102 of the control apparatus main body 33 of embodiment of the modification of 1st Embodiment. It is a block diagram which shows concretely the whole system 14 of 2nd Embodiment of this invention. It is a flowchart which shows operation | movement of the calculating part 102 of the control apparatus main body 33 in 2nd Embodiment. It is a figure which shows the other example of the correction value H. It is a block diagram which shows the ventilator which is other embodiment of this invention. It is a figure which shows the simulation result in case there are few respiration rates per unit time. It is a figure which shows the simulation result in case there are many respiration rates per unit time. It is a graph which shows typically the time change of spontaneous-respiration pressure Pmus and flow volume F of the support gas which flows through an airway.

It is a graph which shows typically the time change of spontaneous-respiration pressure Pmus and flow volume F of the support gas which flows through an airway.

Explanation of symbols

DESCRIPTION OF SYMBOLS 14 Whole system 17 Ventilator 20 Gas supply mechanism 21 Control apparatus 50 Flow rate detection means 51 Flow rate estimation means 52 Deviation calculation means 53 Control amount calculation means 200 Patient respiratory state monitoring means 201 Compensation flow rate generation means Pmus Spontaneous respiration pressure ^ Pmus estimation Spontaneous breathing pressure

Claims (17)

A control method for controlling a gas supply servo mechanism for supplying a support gas containing oxygen to a patient's airway via an inspiratory line for a patient performing a spontaneous breathing operation,
An inspiration start time detection step for detecting an inspiration start time T _{onset at} which the patient's respiratory muscles start inspiration;
A compensation value calculating step of calculating a compensation value H that decreases with the passage of time from the intake start time T _onset ;
A compensation flow value generation step of detecting the flow rate F of the support gas flowing through the patient's airway from the inhalation start time T _onset and generating the compensation flow rate value F _H obtained by adding the compensation value H to the detected flow rate F of the support gas. When,
A control amount calculation step of calculating a target pressure value Pin of the support gas to be discharged from the gas supply servomechanism to the intake pipe from the intake start time T _onset based on the compensation flow rate value F _H ;
And a control command step for giving a command to the gas supply servo mechanism so that the support gas is discharged toward the intake pipe at the target pressure value Pin calculated by the control amount calculation step. Servo mechanism control method. The compensation value H calculated in the compensation value calculation step is set to a value that is predetermined in advance at the intake start time T _onset , and is set to a value that decreases smoothly with the passage of time from the intake start time T _onset. Item 2. A method for controlling a gas supply servomechanism according to Item 1. 3. A control method for a gas supply servo mechanism according to claim 2, wherein the compensation value H calculated in the compensation value calculation step is set to a transient response value obtained by incomplete differentiation of the maximum value from the intake start time T _onset. .   4. The method for controlling a gas supply servo mechanism according to claim 2, wherein the transient change of the compensation value H calculated in the compensation value calculating step is set based on the transient change of the spontaneous breathing pressure Pmus of the patient. The maximum value of the compensation value H calculated in the compensation value calculation step is set to a value corresponding to the magnitude of the flow rate F _onset of the support gas flowing through the airway at the intake start time T _onset . The control method of the gas supply servo mechanism as described in any one of -4.   The maximum value of the compensation value H calculated in the compensation value calculation step is to compensate for a flow rate change delay from when the respiratory muscle of the patient starts the inspiration operation until a change in the flow rate of the support gas according to the inspiration operation occurs. The control method for the gas supply servomechanism according to any one of claims 2 to 4, wherein the control value is set to a specified value H0. The compensation value H calculated in the compensation value calculation step is a value that is a combination of the first compensation value H1 and the second compensation value H2,
The first compensation value H1 is set to a transient response value obtained by incomplete differentiation of a value corresponding to the magnitude of the flow rate F _onset of the support gas flowing through the airway at the intake start time T _onset ,
The second compensation value H2 is set to a transient response value obtained by incomplete differentiation of the specified value H0 for compensating for the flow rate change delay from the intake start time T _onset until the change in flow rate of the support gas according to the intake operation occurs. The method of controlling a gas supply servomechanism according to claim 1. In the control amount calculation process,
A flow rate setting value (K _FG · F _H ) obtained by multiplying the compensation flow rate value F _H by a predetermined flow rate gain K _FG , and a volume obtained by multiplying the integral value V _H of the compensation flow rate value F _{H by} a predetermined volume gain K _VG. A total value ((K _FG · F _H ) + (K _VG · V _H )) obtained by adding a set value (K _VG · V _H ) is calculated as the target pressure Pin. The control method of the gas supply servo mechanism as described in any one of these. In the control amount calculation process,
Using the flow rate estimation means that models the respiratory organs of the patient, the target pressure Pin is input, the output according to a predetermined adjustment transfer function is calculated as the delay compensation pressure Pm, and the support gas of the delay compensation pressure ^ Pm is calculated. A flow rate estimating step for estimating the flow rate of support gas to be supplied to the patient's airway when supplied, F;
A deviation calculating step of calculating a flow rate deviation ΔF _H between the compensation flow rate F _H and the estimated flow rate ^ F;
A flow rate setting value (K _FG · ΔF _H ) obtained by multiplying the flow rate deviation ΔF _H by a predetermined flow rate gain K _FG and a volume setting value obtained by multiplying an integral value ΔV _H of the compensation flow rate ΔF _{H by} a predetermined volume gain K _{VG. (K VG} · ΔV _H) and the total value obtained by adding the _{((K FG · ΔF H)} + (K VG · ΔV H)) claims, characterized in that it comprises a target pressure calculating step of said target pressure Pin Item 8. A method for controlling a gas supply servo mechanism according to any one of Items 1 to 7. Compensation value H for calculating by the compensation value calculation step, decreases with time from the intake start time T _onset, comprises a different flow compensation value H _F and volume compensation value H _V each other,
In the compensation flow value generation process,
The compensated flow value F _HF used to determine the flow rate set value in the control amount calculation step to generate a compensated flow value F _HF using the flow compensation value H _F,
The compensated flow value F _HV used to determine the volume setting value in the control amount calculation step, claim 8, characterized in that generating the compensated flow value F _HV by using the volume compensation value H _V or 9 The gas supply servomechanism control method described. The inhalation start time detection step detects a time when a differential value obtained by differentiating the time change of the flow rate of the support gas increases from a predetermined threshold value as an inhalation start time T _onset. The control method of the gas supply servo mechanism as described in any one of these. An intake trigger detection step of detecting an intake trigger time point T _{ontrg in} which the flow rate F of the support gas flowing through the airway exceeds a predetermined intake start threshold value;
If the intake trigger time T _ontrg is detected before detecting the intake start time T _onset , the control gas calculation step detects the flow rate F of the support gas flowing through the airway, and based on the detected flow rate F of the support gas, The method for controlling a gas supply servomechanism according to any one of claims 1 to 11, wherein a target pressure Pin of the support gas discharged from the gas supply servomechanism is calculated. An inspiration end time detection step of detecting an inspiration end time T _{offset at} which the patient's respiratory muscle ends the inspiration operation;
An expiratory trigger detection step of detecting an expiratory trigger time point T _{offtrg at} which the flow rate F of the assisting gas flowing through the airway becomes less than a predetermined inhalation end threshold value;
_13. The control command step ends the control of the gas supply servo mechanism related to the patient's inspiration support when it is determined that either the inspiration end time T _offset or the expiration trigger time T _offtrg has been reached. The control method of the gas supply servo mechanism as described in any one of these.   The program for making a computer perform the control method of the gas supply servo mechanism as described in any one of Claims 1-13. A control device for controlling a gas supply servo mechanism for supplying a support gas containing oxygen to a patient's airway via an inspiratory line for a patient performing a spontaneous breathing operation,
An inspiration start time detection means for detecting an inspiration start time T _{onset at} which the patient's respiratory muscles start inspiration;
Compensation signal generating means for generating a compensation signal H _sig indicating a compensation value H that decreases with the passage of time from the intake start time T _onset ;
The flow rate F of the support gas flowing through the airway is detected from the intake start time T _onset , the compensation signal H _sig is superimposed on the flow rate signal F _sig indicating the detected flow rate F of the support gas, and the compensation is added to the detected flow rate F. A compensation flow rate signal generating means for generating a compensation flow rate signal F _Hsig indicating a compensation flow rate F _H to which the value H is added;
A control amount calculation means for calculating a target pressure value Pin of the support gas discharged from the gas supply servomechanism based on the compensation flow rate F _H indicated by the compensation flow rate signal F _Hsig from the intake start time _Tonset ;
Control command means for giving a command to the gas supply servo mechanism so as to discharge the support gas at the target pressure Pin toward the patient's airway at the target pressure value Pin calculated by the control amount calculation means, Control device for the gas supply servo mechanism. In the configuration according to claim 15, instead of the inhalation start time detection means, a signal indicating the inhalation start time T _onset is input from a detection device that detects the inhalation start time T _onset at which the patient's respiratory muscle starts the inhalation operation. And a control device for the gas supply servomechanism. A gas supply servomechanism for supplying a support gas containing oxygen to the patient's airway via an inspiratory line for a patient performing spontaneous breathing operation;
A ventilator comprising a control device for a gas supply servomechanism according to claim 15 or 16.

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