Patent Application
20220330837
The present invention pertains to a device for determining an estimate or representation for a respiratory and/or a cardiogenic signal from a signal determined by means of measured values, which determined signal results from a superimposition of the cardiac activity and the breathing and/or ventilation of a patient.
A “signal” shall be defined below as the curve or course in the time range or also in the frequency range of a directly or indirectly measurable indicator, which is variable over time and which is correlated with a physical variable. This physical variable is linked in this case with the cardiac activity and/or with the spontaneous (intrinsic) breathing of a patient and/or with a mechanical ventilation of the patient and is generated by at least one signal source in the body of the patient or by a ventilator. A “respiratory signal” is correlated with the spontaneous breathing and/or with a mechanical ventilation of the patient, and a “cardiogenic signal” is correlated with the cardiac activity of the patient.
The respiratory signal is, in particular, an indicator for the breathing pressure or an indicator for the flow of breathing air relative to the lungs of the patient, wherein this flow of breathing air is effected by the breathing pressure, and the breathing pressure and hence also the breathing flow is caused by the spontaneous breathing of the patient, by mechanical ventilation performed by a ventilator or by a superimposition of the spontaneous breathing and the mechanical ventilation. For example, the pressure in the airway, in the esophagus or in the stomach or an electromyogram may be used as an indicator for the breathing pressure, as a rule, as a pressure difference relative to the ambient pressure. The breathing air flow brings about a change over time in the lung filling level of the patient.
A possible application of the present invention is the control of a mechanical ventilator. This ventilator assists the spontaneous breathing of a patient. The ventilator shall perform ventilation strokes synchronized with the spontaneous breathing of the patient in order for the patient not to breathe against the ventilator. In order to synchronize the ventilator automatically with the spontaneous breathing of the patient, a respiratory signal is needed.
In many situations this respiratory signal cannot be measured in an isolated manner from the cardiogenic signal. Rather, only a sum signal can be obtained, which results from a superimposition of the breathing and/or ventilation and the cardiac activity of the patient. The influence of the cardiac activity on the sum signal shall at least approximately consequently be compensated by calculation in this application.
Conversely, it is frequently desirable to obtain and to use a cardiogenic signal, for example, an improved ECG signal. Only a sum signal, which results from a superimposition of the cardiac activity with the breathing and/or with the mechanical ventilation of the patient, is often available for this task as well. The influence of the breathing and/or ventilation on the sum signal shall be compensated at least approximately in this application. Even if the patient is fully sedated and is ventilated exclusively mechanically, i.e., the patient's intrinsic spontaneous breathing is greatly or even completely suppressed, the ventilation does influence the cariogenic signal.
In the first application, the respiratory signal is the wanted signal and the cardiogenic signal is an unwanted signal to be at least approximately compensated by calculation. In the second application, the cardiogenic signal is the wanted signal and the respiratory signal is the unwanted signal.
An approach to generating a wanted signal from a sum signal of a patient is described in M. Ungureanu and W. M. Wolf: Basic Aspects Concerning the Event-Synchronous Interference Canceller,” IEEE Transactions on Biomedical Engineering, Vol. 53, No. 11 (2006), pp. 2240-2247.
A process and a device for generating two data signals, wherein the first data signal describes an activity of a muscle responsible for inhalation and the second data signal describes an activity of a muscle relevant for exhalation, are described in DE 10 2015 015 296 A1. Two surface myography sensors detect two EMG signals. A cardiac signal component in the EMG signals is suppressed by calculation. In addition, the breathing activity of the patient is determined. A computer detects on the basis of the detected breathing activity when the patient is inhaling and when the patient is exhaling. A first separated signal and a second separated signal are determined on the basis of the two EMG signals.
A process for automatically controlling a ventilation system is described in DE 10 2007 062 214 83. A breathing activity signal uEMG(t) is recorded with electrodes on the surface of the chest in the process known from DE 10 2007 062 214 B3. To make electromyographic signals representing the breathing activity from the recorded electrode signals, the electrode signals must be subjected to a preprocessing; in particular, ECG signal components, which dominate the overall signal in terms of the signal height, must be removed. Filtering as well as an envelope detection may preferably be carried out for this purpose. The enveloping curve detection is preferably carried out by forming the absolute value or by raising to the second power and subsequent low-pass filtering of the electrode signals. Electromyographic signals, which represent the breathing activity and which can be used to control the ventilation drive of the ventilator, are obtained after this preprocessing, as this is described, e.g., in DE 10 2007 062 214 B3.
A medical sensor device 11 is described in DE 10 2009 035 018 A1. Electrodes 12 on the chest of a patient generate electrical signals, from which an electromyogram (sEMG) is generated. An array with an acceleration sensor 6 and with a microphone 7 generates a mechanomyogram (MMG). The measured signals contain an ECG component, which is suppressed by calculation by filtering.
WO 2005/096924 A1 describes a ventilation system (positive pressure ventilation device), which ventilates a patient as a function of EMG signals. Electrodes on the skin of the patient (skin surface electrodes) provide signals, in which the EMG signal being sought is superimposed by an ECG signal. The ECG component is removed from the measured signal by calculation, so that an EMG signal from which unneeded components have been removed (moving average electromyogram signal) is generated. This signal is displayed (displaying).
A process for generating signals of a fetus in the womb, especially the heartbeat activity of the fetus (fetal heart rate), is described in US 2007/0191728 A1. Electrodes 20, 21 and 22 on the abdomen of the pregnant mother measure a superimposition of ECG and EMG signals. The ECG signals are separated from the EMG signals by calculation, and the fetal signals are distinguished from the signals of the pregnant mother by calculation. EP 2371412 A1 shows a device for the mechanical ventilation or also anesthesia of a patient. An sEMG sensor 6 on the skin of the patient detects the electromyographic muscle activity of the breathing muscles of the patient.
U.S. Pat. No. 6,411,843 B1 describes a process and a device for obtaining a processed EMG signal (model EMG signal) from a measured signal, which is formed from a superimposition of an EMG signal and of an ECG signal of a patient. An envelope (first envelope signal) is calculated from the measured signal. In addition, heartbeat times are detected in the measured signal. The processed EMG signal is generated from the generated envelope and the detected heartbeat times. A first logical signal, in which the P wave, the QRS complex and the T wave are removed by calculation, and a second logical signal, in which the P wave, the QRS complex or the T wave are contained, are derived from the measured EMG signal. Furthermore, a first envelope is derived from the measured EMG signal. A modeled EMG signal is derived from the first envelope and from the first logical signal, on the one hand, and from a signal that depends on the second logical signal, on the other hand.
DE 10 2012 003 509 A1 describes a ventilation system with a control device and with a patient module. Electrodes of the patient module derive electrode signals from the surface of the chest of a patient. The control device suppresses in the electrode signals ECG components and derives first ECG signals. Data that represent the ECG are sent in the digital form to an ECG output, on the one hand, and are converted, on the other hand, into an analog signal, which is made available for display.
It is proposed in WO 2018001929 A1 that a first undesired signal component be reduced from a physiological signal by subtracting a model of the undesired signal from the physiological signal. A residual signal is obtained thereby. A filter unit reduces in the residual signal a second undesired signal by a band elimination filter (notch filter) generating a filtered signal. A gating is applied to the filtered signal
A basic object of the present invention is to provide a process and a signal processing unit, which determine an estimate or representation for a cardiogenic signal and/or for a respiratory signal from a sum signal, which sum signal is generated by means of measurements of a signal generated in the body of the patient and results from a superimposition of the cardiac activity of the patient to the spontaneous breathing and/or to the mechanical ventilation of the patient, better than do prior-art processes and signal processing units.
The object is accomplished by a process having the process features of the invention and by a signal processing unit having the device and system features of the invention. Advantageous embodiments are described herein. Advantageous embodiments of the process according to the present invention are also advantageous embodiments of the signal processing unit according to the present invention and vice versa.
According to the present invention, an estimated (representative) cardiogenic signal and/or an estimated (representative) respiratory signal are calculated. The calculated respiratory signal is correlated with the spontaneous breathing and/or with a mechanical ventilation and especially with the flow of breathing air relative to the lungs of the patient. This flow of breathing air may be brought about exclusively by the spontaneous breathing of the patient, exclusively mechanically by mechanical ventilation by means of a ventilator (e.g., the patient is fully sedated) or by the spontaneous breathing assisted by the mechanical ventilation. The calculated respiratory signal also contains a component, which is caused by the cardiac activity. This component is, however, as a rule, smaller than the component in the sum signal generated on the basis of the measurements. The determined cardiogenic signal is correspondingly an indicator for the cardiac activity of the patient. The cardiogenic signal contains a component, which is caused by the breathing or ventilation, and this component is smaller than the respiratory component in the sum signal.
The process according to the present invention comprises a training phase and a subsequent use phase and is carried out automatically with the use of the signal processing unit according to the present invention.
During the training phase the signal processing unit receives measured values from a sum signal sensor device comprising at least one sum signal sensor. The sum signal sensor device measures a signal, which is generated in the body of the patient. Optionally the signal processing unit receives measured values from the sum signal sensor device during the use phase as well.
The signal processing unit generates a sum signal during the training phase. This generated sum signal comprises a superimposition of the cardiac activity to the spontaneous breathing and/or to the mechanical ventilation of the patient. In order to generate the sum signal, the signal processing unit uses the respective time curve (temporal course) of measured values, which have been provided by the sum signal sensor device. Optionally the signal processing unit also generates the sum signal during the use phase.
The signal processing unit detects during the training phase a plurality of heartbeats, which the patient has performed during the training phase, preferably each heartbeat. The signal processing unit generates a sample, which comprises a plurality of sample elements. Each sample element of the sample pertains to a respective heartbeat detected during the training phase.
To generate a sample element for a heartbeat, the signal processing unit carries out the following steps:
The signal processing unit generates in the training phase a signal estimating unit (signal representation unit). The generated signal estimating unit yields the shape parameter or each shape parameter (at least one shape parameter) as a function of the first transmission channel parameter and optionally as a function of at least one additional transmission channel parameter. The signal processing unit uses the sample with the sample elements for the generation.
In the use phase, the signal processing unit detects at least one heartbeat, which the patient carries out in the course of the use phase. The signal processing unit preferably detects each heartbeat in the use phase or at least in a time period of the use phase.
The signal processing unit carries out the following steps during the use phase for at least one detected heartbeat and preferably for each detected heartbeat:
In a first alternative of the present invention, the signal processing unit determines in the use phase the estimated cardiogenic signal. For doing so, it combines the estimated cardiogenic signal segments for the heartbeats detected during the use phase to the estimated cardiogenic signal. In a second alternative of the present invention, the signal processing unit determines during the use phase the estimated respiratory signal. For doing so, it combines the estimated respiratory signal segments for the heartbeats detected during the use phase to the estimated respiratory signal. In a third alternative of the present invention, the signal processing unit likewise determines during the use phase the estimated respiratory signal, but, contrary to the second alternative, it does so by a compensation by calculation. This determination by compensation comprises the following steps:
Thanks to the present invention, it is not necessary to generate the respiratory signal or the cardiogenic signal by direct measurement. This is, as a rule, not possible at all, or even though it is possible, it is not desired, e.g., because a sensor and/or a maneuver needed therefor would stress the patient too strong during the operation of the ventilator. A sum signal is rather generated according to the present invention from the measured values of the sum signal sensor device, and the respiratory signal and/or the cardiogenic signal is determined by calculation using this sum signal.
According to the present invention, a signal estimating unit is generated automatically and a sample with a plurality of sample elements, which was generated during the training phase, is used for this generation. Since a sample is generated empirically and then used, no analytical model is needed; in particular, no model that analytically describes the influence of the cardiac activity or of the breathing/ventilation is needed. Such a model often cannot be set up at all or it can be set up and validated and adapted to a patient at an unacceptably great effort only. However, the present invention can be used in a plurality of configurations combined with an analytical model.
According to the present invention, this sample is generated with the use of measured values that are measured during the training phase at that patient for whom the steps of the subsequent use phase are carried out as well. The present invention therefore avoids errors that would appear, as a rule, if measurements were carried out during the training phase on at least one patient and the results of the training phase were applied in the use phase to another patient. Such errors would often also occur if measurements were carried out during the training phase on a plurality of patients and averaging was performed over the measurements.
Since according to the present invention a plurality of sample elements are used in order to generate the signal estimating unit, the influence of freak values is removed to a certain extent by averaging.
The same sum signal sensor device can be used in both the training phase and in the use phase. The use of different sensor devices during the two phases, which is avoided according to the present invention, could cause additional errors. The present invention avoids this possible source of error.
According to the present invention, the signal estimating unit, which is generated during the training phase, yields a respective estimated signal segment for at least one and preferably for each heartbeat detected during the use phase. The provided estimated signal segment may differ from one heartbeat to the next. The present invention takes into consideration the following circumstance: The anthropological variable, especially the spontaneous breathing and/or the mechanical ventilation of the patient, influences the respective transmission channel guiding from nerves and/or muscles, which elicit the cardiac activity and/or the spontaneous breathing, to the sum signal sensor device. The spontaneous breathing therefore acts additionally as a disturbance signal on the cardiogenic signal and hence also on the sum signal. The influence of the spontaneous breathing varies, as a rule, from one heartbeat to the next. A mechanical ventilation of the patient also influences such a transmission channel, and this influence may vary from one heartbeat to the next. The first transmission channel parameter, which is taken into account according to the present invention, correlates with the effect that the spontaneous breathing or mechanical ventilation or another anthropological variable has on the transmission channel to the sum signal sensor device, and can be measured. This transmission channel is located completely or at least partially in the body of the patient. Due to the first transmission channel parameter being measured and due to the measured transmission parameter value being analyzed, the effect of breathing and/or ventilation or another anthropological variable on the transmission channel and hence on the cardiogenic signal can at least approximately be taken into account.
In other applications, the anthropological variable may correlate with the cardiac activity of the patient and act as a disturbance signal on the respiratory signal and hence on the sum signal. The present invention makes it possible to compensate the influence of this disturbance variable by calculation in these applications as well.
The signal estimating unit provides in the use phase a respective estimated signal segment each for at least one and preferably for each detected heartbeat. This estimated signal segment pertains to the time period of an individual detected heartbeat. According to the present invention, the estimated signal segment for a heartbeat depends on the value or on at least one value for the first transmission channel parameter that was measured during this heartbeat. The estimated signal segment calculated by the signal estimating unit does consequently take into consideration at least approximately the effect of the anthropological variable, especially the effect of the spontaneous breathing or mechanical ventilation or also of the cardiac activity on the transmission channel during this heartbeat.
If, by contrast, the same estimated or predefined signal segment, for example, a predefined standard signal segment (e.g., a so-called ECG template), were used for each heartbeat, the varying influence especially of breathing or ventilation on the or one transmission channel could not be taken into account at all or it would be able to be taken into account at least only to a markedly lesser extent.
The estimated signal segment for a heartbeat, which is calculated during the use phase, depends on the value or on at least one value for the first transmission channel parameter that was measured in the use phase during this heartbeat. This estimated signal segment is therefore adapted to the anthropological variable, more precisely, adapted to the influence that the anthropological variable in the body of the patient has on the transmission channel from at least one muscle or from another signal source in the body of the patient to the sum signal sensor device used during this heartbeat. The signal source, in which a transmission channel to a sum signal sensor device begins, is, e.g., a heart muscle or a muscle of the respiratory system.
According to the present invention, the signal estimating unit yields during the use phase an estimated signal segment for the time period of this heartbeat. This signal estimating unit is generated during the training phase automatically by means of a sample, wherein each sample element of this sample comprises at least one transmission channel parameter value and at least one associated shape parameter value.
In a first alternative of the present invention the estimated signal segment describes a segment of the estimated cardiogenic signal in the course of this heartbeat. The signal processing unit combines these segments during the use phase to the cardiogenic signal. In another alternative, the signal processing unit uses at least one segment for a heartbeat (estimated cardiogenic signal) in order to compensate the influence of the cardiac activity during this heartbeat on the sum signal by calculation, especially by the segment being subtracted from the sum signal. In yet another alternative, the estimated signal segment describes a segment of the estimated respiratory signal in the course of this heartbeat. In one embodiment, the signal processing combines these segments to the respiratory signal.
In one embodiment, the entirety of the shape parameter values of a sample element for a heartbeat specifies a segment of the respiratory or cardiogenic signal to be determined in the course of this heartbeat. The shape parameter values are, for example, support points of a reference signal segment in the course of a heartbeat.
In another embodiment, a standard reference signal segment is predefined, which segment is valid for each heartbeat and preferably for each patient and depends on at least one shape parameter and preferably describes the cardiac activity. The shape parameter value or the entirety of the shape parameter values together with this standard reference signal segment specify the segment of the sum signal in the time course of the heartbeat during the training phase and the estimated signal segment during the use phase. Preferably the estimated signal segment describes in this application a segment of the cardiogenic signal.
In one embodiment, a change rule is predefined, which vile depends on the shape parameter or on at least one shape parameter. During the use phase the signal estimating unit yields at least one respective shape parameter value for each segment and uses this value for the change rule. The signal processing unit determines for each heartbeat the segment of the sum signal, which segment belongs to this heartbeat. The signal estimating unit applies the change rule parameterized for this heartbeat to a determined sum signal segment. As a result, the signal estimating unit yields the estimated signal segment for this heartbeat. The estimated signal segment can describe a segment of the respiratory signal or of the cardiogenic signal to be determined.
In another embodiment, the shape parameter value or the entirety of the shape parameter values specify a calculation rule wherein the rule is adapted to calculate the estimated signal segment for the heartbeat in question from the segment of the sum signal for a heartbeat, which segment was determined during the use phase.
In one embodiment, a computer-accessible library with a plurality of reference signal segments is generated during the training phase, and each reference signal segment pertains to a class of possible values of the first transmission channel parameter and optionally of at least one additional transmission channel parameter and describes a segment of the estimated cardiogenic or respiratory signal during a heartbeat. Each reference signal segment is generated with the use of at least one and preferably a plurality of sample elements, whose parameter values fall within this class, wherein corresponding segments of the sum signal are combined in a suitable manner to the reference signal segment. The value or at least one value of the first transmission channel parameter and optionally a respective value of each additional transmission channel parameter is measured for a detected heartbeat during this heartbeat during the use phase. Depending on the measured parameter value, at least one reference signal segment is selected in the library, and the estimated signal segment is generated depending on the reference signal segment or each selected reference signal segment.
For example, the two reference signal segments that belong to the two parameter values that are located adjacent to the measured value of the first transmission channel parameter are selected, and the estimated signal segment is generated as a weighted mean value over these two selected reference signal segments. The weighting factors are calculated, e.g., such that the estimated signal segment is an interpolation between the two reference signals.
The embodiment with the library causes an estimated signal segment to be calculated rapidly during the use phase for a detected heartbeat and relatively little storage space to be needed.
In one embodiment, a plurality of classes pertain to partial areas of a regular area of the first transmission channel parameter or of an additional transmission channel parameter and to at least one additional class of “freak values,” which occur in unusual situations, for example, when the patient is coughing or has a muscle spasm or exerts himself excessively or the patient's heartbeat shows a spontaneous irregularity.
In one embodiment, a plurality of transmission channel parameters are predefined and taken into consideration, and these transmission channel parameters influence the transmission channel or at least one respective transmission channel to the sum signal sensor device. The signal estimating unit is generated during the training phase such that the signal estimating unit yields an estimated signal segment in the course of a heartbeat as a function of a plurality of transmission channel parameters. This embodiment makes it possible to take into consideration simultaneously a plurality of different influential factors on the same transmission channel. It is possible, but thanks to the present invention not necessary to determine such parameters, which are dependent on one another in advance by calculation. This may require a large amount of computing time and/or be time-consuming.
The training phase comprises a plurality of heartbeats, preferably between 20 heartbeats and 60 heartbeats. The use phase preferably begins immediately after the end of the training phase.
In one embodiment, the training phase ends after a predefined number of heartbeats and/or after a predefined time period. It is also possible that the training phase ends as soon as a sufficient number of different values have been measured for the first transmission channel parameter or for each transmission channel parameter.
According to the present invention, a respective characteristic time and/or a time period of a heartbeat is measured during the use phase. In one embodiment, the sum signal sensor device, comprising at least one sum signal sensor, yields an electrical signal, and the fact is utilized that an electrical signal, which is caused by an individual heartbeat, typically has a curve that comprises a P wave, a QRS wave and a T wave. These waves and the corresponding peaks can also be determined in the sum signal, because the percentage of the cardiogenic signal between the P wave and the T wave is several times greater than the percentage of the respiratory signal. This designation P through T has become established in the literature. The Q peak, the R peak or the S peak of this heartbeat, especially the R peak, is preferably used as the characteristic time of a heartbeat. In another embodiment, a respective heartbeat time period is determined during the use phase for each heartbeat, and the heartbeat takes place during this time period and/or the determined time period comprises this heartbeat. The heartbeat time period extends, for example, from the P wave to the T wave. The heartbeat time period is determined, for example, by an analysis of the sum signal. At least when the sum signal is determined by means of measured values of electrical sensors, the influence of the heartbeats in the sum signal is many times greater than the influence of the breathing activity. The time in this determined heartbeat time period at which the sum signal assumes a maximum or a minimum is preferably detected as the characteristic time of the heartbeat.
In an alternative of the present invention, the influence of at least one detected heartbeat, preferably of each detected heartbeat, on the sum signal is compensated by calculation when a respiratory signal shall be determined. Different embodiments of how this compensation can be carried out are possible. In one embodiment a heartbeat time period is determined for a detected heartbeat. The sum signal and the characteristic heartbeat time period are used for this. For example, the heartbeat time period covers the P wave, the QRS wave and the T wave. The estimated signal segment for the heartbeat is subtracted from the sum signal in the heartbeat time period, or the estimated signal segment multiplied by a factor and/or shifted by a time delay is subtracted.
According to the present invention, at least one value, which the first transmission channel parameter assumes at this heartbeat, is measured for each heartbeat. Optionally a value of at least one additional transmission channel parameter is measured at this heartbeat. The term “value” may designate a single number, i.e., a scalar, or also a vector. For example, the position of a sum signal sensor relative to the heart or to another reference point in the body of the patient may be used as a transmission channel parameter. This relative position depends on the current lung filling level. Each value of this transmission channel parameter is preferably a vector with three components, for example, in a three-dimensional cartesian coordinate system.
In one embodiment the sum signal can be generated by means of passively operating measuring electrodes, which electrodes are positioned at or in the body of the patient and provide a respective electrical measured value each (especially surface electromyogram or electromyogram in the body, e.g., in the esophagus or in the stomach). Each measured electrical value depends on the current activity of the diaphragmatic muscles as well as on the activity of the auxiliary muscles and optionally on the mechanical ventilation of the patient. The measured values of the measuring electrode can be analyzed, which leads to an electrical sum signal.
In many cases, a prediction can be made about the patient on the basis of a respiratory signal, which was determined according to the present invention from an electrical sum signal, better than with other methods. For example, the heart rate can better be predicted based on a cardiogenic signal, which was determined according to the present invention from the electrical sum signal.
In one embodiment, an electrical impedance tomography belt (EIT belt) is used as a sum signal sensor device and/or as a sensor for a transmission channel parameter. Such an EIT belt is applied on the skin of the patient and it comprises a plurality of signal units, which can be operated alternatingly as signal source or signal receiver. At each time exactly one signal unit is a signal source, and the other signal units are signal receivers. The signal source generates a high-frequency signal, preferably in the range of several kHz, which is harmless for the patient and penetrates into the body of the patient. The EIT belt measures the respective electrical impedance in the body of the patient between the signal source and a signal receiver. The electrical impedance in a body part filled with air, especially in the lungs, is several times higher than the electrical impedance in a tissue, which is filled with a salt-containing and therefore electrically conductive solution.
The EIT belt thus generates an image of the lungs in the body of the patient, which is variable over time. If the lung filling level of the patient is the transmission channel parameter or a transmission channel parameter, the signal processing unit is capable of determining the current lung filling level from the image of the lungs, e.g., by image processing. It is also possible that the signal processing unit uses the image of the lungs, which is variable over time, as a sum signal.
In one embodiment, the image of the lungs is divided into a plurality of areas, wherein each area shows a region of the lungs. For example, the image is divided into four quadrants or into a plurality of pixels. Each image area is used as a respective sum signal. By analyzing this sum signal or these sum signals, the signal processing unit is capable of detecting the heartbeats. It is also possible that the signal processing unit receives measured values from another sensor, which detects the heartbeat times and/or heartbeat time periods.
The measured electrical value, which is brought about by the heart muscle, is several times higher than the measured value that is brought about by the breathing muscles. The cardiac activity causes voltages in the mV range, and the breathing activity causes voltages in the microV range. The higher voltages of the cardiac activity occur essentially only from time to time, namely, from time to time in the course of a heartbeat, and not during the rest of the time nor between the heartbeats. It is especially because of this that it is possible to obtain the respiratory signal from the sum signal.
According to the present invention, the respective value, which the or each shape parameter assumes at the heartbeat in question, is determined for each heartbeat. The segment of the sum signal belonging to this heartbeat is preferably used for this determination. Especially if the sum signal is an electrical signal, the sum signal in a heartbeat is essentially equal to the cardiogenic signal. The influence of the respiratory signal is often averaged over a plurality of sample elements when the signal estimating unit is generated.
It is also possible to obtain the sum signal by means of at least one pneumatic sensor, in which case the pneumatic sensor measures, for example, an indicator for the flow of gas into or out of the lungs of the patient and/or the airway pressure. This flow is measured, e.g., at a ventilator which is connected to the patient, or at the mouth of the patient. For example, the volume flow and the achieved ventilation pressure are measured in a fluid connection between the patient and the ventilator. A time delay between the lungs of the patient and the connection ventilator is predefined or estimated, and the time delay is used to correct measurements that are carried out at the ventilator in terms of time and to compensate the delay by calculation in the process.
In one embodiment, the sum signal sensor device comprises a probe or a balloon or a catheter, which is inserted into the body of the patient, for example, into the esophagus, and an electrical or pneumatic transducer. It is also possible to measure the breathing muscles by means of a sensor for a mechanomyogram or vibromyogram. In one embodiment, at least one catheter, which measures the esophageal pressure or the gastric pressure, is used as a sum signal sensor of the sum signal sensor device.
In another embodiment the sum signal sensor device comprises an image recording device, which is directed towards the patient. An imaging method is applied to the signals from the image recording device. This embodiment eliminates the need to position the sum signal sensor at or even in the patient. Rather, a spatial distance between the patient and the sum signal sensor remains. This embodiment leads to a greater tolerance in case of deviations between a desired position and an actual position of a sum signal sensor relative to the patient.
It is also possible to combine different types of sensors with one another. The sum signal is generated in this embodiment from measured values of different sensors.
According to the present invention, the signal processing unit receives measured values from the sum signal sensor or from at least one sum signal sensor. The measured values are preferably processed, for example, amplified and/or smoothed, and/or disturbing influences are filtered out of the measured values. In addition, analog measured values are converted into digital measured values. If a measuring electrode is positioned on the skin of the patient and is used as the or one sum signal sensor, electrochemical effects, which develop due to the contact between the measuring electrode and the skin, especially between the silver of the electrode and sweat on the skin, are preferably compensated by calculation (baseline removing, baseline filtering), and other potential differences are compensated. The signal processing unit generates the sum signal from the measured values processed in this manner and uses especially the processed measured values as the sum signal.
According to the present invention, at least one value per heartbeat of the first transmission channel parameter is measured. This measured transmission channel parameter is correlated with at least one anthropological variable, which influences a transmission channel from a signal source in the body of the patient to the sum signal sensor or to at least one sum signal sensor. In one embodiment, the anthropological variable or an anthropological variable is the current geometry of the body of the patient. This body geometry depends in many cases on the current lung filling level of the patient. The first transmission channel parameter is thus correlated with the lung filling level of the patient.
In one embodiment, a mechanical or pneumatic or optical sensor measures an indicator for the body geometry, for example, the flow of breathing air into the lungs or out of the lungs or the body circumference of the patient in such a measurement position in which the body circumference varies with the lung filling level. An optical sensor comprises especially an image recording device and an image analysis unit, which employs an imaging method. The variable body geometry influences the transmission channel from the heart or from a part of the breathing muscles to the sum signal sensor or to at least one sum signal sensor, for example, because the distance varies.
In one embodiment, the current posture or body position of the patient is used as the transmission channel parameter or as a transmission channel parameter, for example, the position of the patient in a bed or whether the upper body of the patient is upright or curved. The posture also influences the transmission channel.
In one embodiment, the anthropological variable causes the time interval between two consecutive heartbeats to vary and has, for example, a periodicity covering at least two heartbeats or is irregular. This interval is an indicator indicating how rapidly the heart muscle recovers after a heartbeat. Or else the time interval between two peaks of the sum signal is influenced by the anthropological variable, and the peaks are reached in the course of a heartbeat. The anthropological variable is, for example, the posture of the patient or even an irregularity in the cardiac activity. The time interval between two consecutive heartbeats or the time interval between two peaks in the course of the same heartbeat, e.g., the amplitude of this heartbeat, is used as the first transmission channel parameter or as an additional transmission channel parameter. The signal estimating unit yields the estimated signal segment as a function of the heartbeat distance. This embodiment requires no additional sensor for the first transmission channel parameter. The measured values of the sum signal sensor or the measured values of the array of sum signal sensors rather yield both the sum signal and the values of the first transmission channel parameter. Or else the value of the first transmission channel parameter is calculated by analyzing the sum signal.
In one embodiment, existing knowledge about the signal being sought in the course of a heartbeat is used. This existing knowledge was gained, for example, by means of a plurality of samples on a plurality of patients. In one embodiment the existing knowledge, which is caused by the cardiac activity in the course of a heartbeat and which depends on the shape parameter or at least one shape parameter, is predefined in the form of a standard reference signal segment for the process according to the present invention. During the training phase the signal processing unit generates a signal estimating unit, which yields the shape parameter or each shape parameter of the standard reference signal segment in the form of the used parameter or each used transmission channel parameter. During the use phase the signal processing unit applies for each detected heartbeat the signal estimating unit to the respective value of the measured transmission channel parameter or of each measured transmission channel parameter, which yields a respective value of each shape parameter. Using these shape parameter values, the signal processing unit adapts the predefined standard reference signal segment anew for each detected heartbeat, for example, by the signal processing unit introducing the shape parameter values into the standard reference signal segment. The standard reference segment adapted in this manner acts as the estimated signal segment for this heartbeat, or the estimated signal segment depends on the adapted standard reference signal segment in another manner. The shape parameter or a shape parameter may be, e.g., a time shift, a compression factor/stretching factor along the time axis or a signal amplification factor. The shape parameter or a shape parameter may influence the entire standard reference signal segment or also only at least one defined segment of the standard reference signal segment, e.g., segments with a great slope or segments with a small slope.
This embodiment with the standard reference signal segment, which is valid for each heartbeat and is parameterized, saves computing time and/or storage space in many cases. In order to specify a segment of a sum signal, substantially more points are needed, as a rule, than there are shape parameters. A maximum of five, at times even only three shape parameters are often sufficient.
In one embodiment a single such standard reference signal segment is used. In another embodiment, the value range of the first transmission channel parameter and/or of an additional transmission channel parameter is divided into classes in advance. A respective standard reference signal segment, which depends on the shape parameter or a shape parameter, is assigned to each class. During the training phase the signal processing unit generates a respective signal estimating unit each for each class and hence for each standard reference signal segment. During the use phase the signal processing unit decides for the detected heartbeat into which class the value of the first transmission channel parameter or of an additional transmission channel parameter, which value was measured during this heartbeat, falls. It selects the associated standard reference signal segment as well as the adapted signal estimating unit and fits the selected standard reference signal segment by applying the selected signal estimating unit.
In one embodiment, the signal processing unit carries out all process steps in the time range. In another embodiment, the signal processing unit transforms during the training phase for each heartbeat a segment of the sum signal belonging to this heartbeat from the time range into the frequency range. The generated signal estimating unit yields an estimated signal segment in the frequency range as a function of the first transmission channel parameter and in one embodiment additionally as a function of another transmission channel parameter. During the use phase the signal processing unit calculates for at least one detected heartbeat an estimated signal segment in the frequency range, transforms this segment into an estimated signal segment in the time range and uses the estimated signal segment in the time range in the manner according to the present invention. It is also possible that a respiratory or cardiogenic signal is generated and used in the frequency range from the sum signal generated in the time range by applying the process according to the present invention in the use phase.
The embodiment of transforming a segment of the sum signal during the training phase makes it possible to apply defined signal processing processes in the frequency range, for example, in order to remove disturbance signals with defined frequencies and to generate sample elements from cleaned-up segments of the sum signal in the frequency range. For example, the signal processing unit applies a low-pass filter, a high-pass filter and/or another band pass filter, removes frequencies in certain ranges, for example, in the range of the line voltage (50 Hz in Germany), or applies wavelet denoising or empirical-mode decomposition-based denoising. In one embodiment, at least one first frequency range is predefined, and in a preferred embodiment a plurality of preferably disjunct first frequency ranges are predefined. The signal processing unit generates an overall sum signal. The signal processing unit determines for each predefined first frequency range a respective signal component each, which is in this first frequency range. The signal processing unit determines, furthermore, a respective respiratory signal component and/or a cardiogenic signal component for the first frequency range or for each first frequency range and/or a cardiogenic signal component. The signal processing unit applies here the process according to the present invention for the first frequency range or for each first frequency range repeatedly, and the signal processing unit uses the signal component in this first frequency range as the sum signal. The signal processing unit subsequently determines the respiratory signal and uses for this determination the respiratory signal component or each respiratory signal component that is in the first frequency range or in a first frequency range and was determined by the use of the process according to the present invention. For example, it adds these respiratory signal components. Or else, the signal processing unit determines the cardiogenic signal and uses for this the cardiogenic signal component or each cardiogenic signal component in the first frequency range or in a first frequency range.
According to the present invention, the value which the respective transmission channel parameter assumes at a heartbeat is measured for the first transmission channel parameter and optionally for at least one additional transmission channel parameter. The signal processing unit receives these transmission channel parameter values. In one embodiment, a respective value each is measured for each transmission channel parameter and for each heartbeat. In another embodiment a split-up of the heartbeat period into at least two heartbeat time period phases, which split-up is valid for each heartbeat, is predefined. For example, a split-up into a first phase with the P-wave or P peak, into a second phase with the QRS-wave or QRS peak and into a third phase with the T-wave or T peak is predefined.
In this embodiment the signal processing unit receives for each detected heartbeat and for each transmission channel parameter a respective value per heartbeat time period phase of this heartbeat. During the training phase the signal processing unit generates a respective sample element each for each heartbeat time period phase of each detected heartbeat. If the training phase comprises, for example, 50 heartbeats and three heartbeat time period phases are predefined, the signal processing unit generates 50 sample elements per phase, i.e., a total of 50×3=150 sample elements.
During the use phase the signal processing unit calculates a respective shape parameter or—in case of a plurality of shape parameters—a set of shape parameter values for one and preferably for each detected heartbeat and for each heartbeat time period phase of this heartbeat. In case of ten shape parameters and three predefined heartbeat time period phases, these equal 10×3=30 shape parameter values per detected heartbeat. The signal processing unit calculates the estimated signal segment for this heartbeat with the use of the shape parameter values for the heartbeat time period phases, i.e., for example, of the 30 shape parameter values.
Preferably the signal processing unit generates in the training phase a respective signal phase estimating unit for each heartbeat time period phase. This signal phase estimating unit is valid for this heartbeat time period phase and yields, just like the signal estimating unit, the shape parameter or each shape parameter as a function of the transmission channel parameter or each transmission channel parameter. To generate this signal phase estimating unit. the signal processing unit uses the sample elements that belong to this heartbeat time period phase.
In this embodiment, the signal processing unit applies for each detected heartbeat the signal phase estimating unit for a heartbeat time period phase to the transmission channel parameter or to each transmission channel parameter that was obtained in this heartbeat time period phase of this heartbeat. As a result, a signal segment that describes the respiratory or cardiogenic signal in this heartbeat time period phase of this heartbeat is used. The signal processing unit generates the estimated signal segment for the heartbeat with the use of all signal segments for the phases of this heartbeat. For example, the signal processing unit combines the signal segments for the heartbeat time period phases to the estimated signal segment.
In one embodiment, the signal estimating unit, which the signal processing unit has generated during the training phase, is used unchanged during the entire use phase. In a preferred embodiment, the signal estimating unit is, by contrast, adapted during the use phase at least once, preferably continually, to the measured values obtained so far during the use phase. In this preferred embodiment, the signal processing unit generates the sum signal also during the use phase. In addition, the signal processing unit generates during the use phase at least one additional sample element, which pertains to a heartbeat detected in the use phase, preferably one additional sample element each for each heartbeat detected during the use phase. The signal estimating unit generated during the training phase is adapted during the use phase at least once with the use of the sample element or of another sample element. The signal estimating unit is preferably adapted continually to all further sample elements generated up to that time during the use phase. For example, a signal estimating unit is generated again repeatedly, doing so by means of a sample from the sample elements of the training phase and from the sample elements or at least some sample elements generated up to that point during the use phase.
In other words, the training phase acts in this embodiment as a starting phase for the generation of the signal estimating unit, and the use phase overlaps an improvement phase or adaptation phase for the signal estimating unit.
In particular, this embodiment makes it possible to take into consideration at least approximately the following influential factors in the course of the use phase:
In one embodiment, a respiratory signal is determined during the use phase. This respiratory signal can be used, for example, for the following applications:
The cardiogenic signal generated according to the present invention can be used instead of a conventionally determined ECG signal, and the same measuring electrodes can continue to be used. The cardiogenic signal approximately compensates the influence of the anthropological variable or at least one anthropological variable, especially that of the breathing activity on the measured signal.
The present invention will be described below on the basis of an exemplary embodiment. The various features of novelty which characterize the invention are pointed out with particularity in the claims annexed to and forming a part of this disclosure. For a better understanding of the invention, its operating advantages and specific objects attained by its uses, reference is made to the accompanying drawings and descriptive matter in which preferred embodiments of the invention are illustrated.
In the drawings:
Referring to the drawings, the process according to the present invention is used in one application to automatically control a mechanical ventilator. This ventilator assists the spontaneous breathing of a patient or replaces same completely if the patient is sedated. This work of the ventilator, especially the times and amplitudes of the ventilation strokes, shall be synchronized—if present—with the spontaneous breathing of the patient.
A signal processing unit 5, which preferably belongs to the ventilator 1, generates a sum signal SigSum by processing measured values of the sensors 2.1 and 2.2 and/or of the pneumatic sensor 3 and/or of the optical sensor 4. This sum signal SigSum results from a superimposition of a respiratory signal Sigres and a cardiogenic signal Sigkar. The respiratory signal Sigres describes in this application the intrinsic breathing activity of the patient P. This respiratory signal Sigres is used for controlling the ventilator 1 and is the wanted signal. The cardiogenic signal Sigkar is caused by the cardiac activity of the patient P and is in this application an unwanted signal. The spontaneous breathing of the patient P, which is described by the respiratory signal Sigres, as well as the mechanical ventilation by the ventilator 1, generate an overall breathing and ventilation of the patient P, which is described by an overall signal Sigges.
The breathing muscles AM of the patient P generate a breathing activity. The heart muscle HM generates a cardiac activity. The intrinsic breathing activity is transmitted in the body of the patient P via a transmission channel Tss to a summation point Σ, where—stated in a simplified manner—the respiratory signal Sigres appears behind the transmission channel Tss. The cardiogenic signal Sigkar is transmitted via a transmission channel Tns to the summation point Σ, and the cardiogenic signal Sigkar appears behind the transmission channel Tns. The transmission channels Tss and Tns thus influence the measured breathing activity and the measured cardiac activity. The signals Sigres and Sigkar are superimposed to one another—simply speaking—in this summation point Σ. In addition, a transmission channel Tnn is shown. The subscript s designates the wanted signal, and the subscript n (noise) the unwanted signal.
The sensors 2.1 and 2.2 generate respective electrical measured values, as a rule, electrical voltages. A signal processor 13 with an amplifier and with an analog-digital converter processes these electrical measured values. The signal processor 13 preferably carries out, in addition, a baseline filtering, especially in order to compensate electrochemical processes in the measuring electrodes 2.1 and 2.2 on the skin of the patient P and other low-frequency potential differences by calculation. These processed measured values act in the exemplary embodiment as the sum signal or a sum signal Sigsum. The sensors 2.1 and 2.2 are therefore sum signal sensors forming a sum signal sensor device in the sense of the present invention. The pneumatic sensor 3 and the optical sensor 4 also yield measured values, from which a sum signal is generated in variants of the present invention and another parameter value is generated in other variants.
The signal processing unit 5, which preferably forms a part of the ventilator 1, determines from this sum signal Sigsum the estimate Sigres.est for the respiratory signal Sigres being sought. The signal processing unit 5 determines for this an estimate Sigkar,est for the cardiogenic signal Sigkar, which acts in this application as an unwanted signal. In other applications, the estimated cardiogenic signal Sigkar,est is used as a wanted signal, and the respiratory signal Sigres is an unwanted signal. Or else both signals Sigres and Sigkar are wanted signals.
The cardiogenic signal segment SigHz,kar describes an estimated segment of the cardiogenic signal in the course of a single heartbeat. A heartbeat time detector 7 detects the respective time H_Zp(n) of the nth detected heartbeat (n=1, 2, 3, . . . ). This heartbeat time detector 7 detects, for example, the so-called R peak or also the QRS curve in the sum signal SigSum or also in a signal that is obtained exclusively from measured values of the set 2.1 of measuring electrodes that is placed close to the heart, cf.
The respiratory signal Sigres and/or the cardiogenic signal Sigkar are influenced by at least one respective anthropological variable in the body of the patient P. A measurable parameter, which acts on at least one above-described transmission channel Tss, Tns and is therefore called transmission channel parameter, correlates with the anthropological variable or with at least one anthropological variable. This influence is not taken into consideration in
An additional transmission channel parameter, namely the position Pos of a measuring electrode 2.1 or 2.2 relative to the signal source for the cardiogenic signal Sigkar, is optionally taken into account in the example shown in
A third transmission channel parameter, which does not require an additional sensor, especially the length of a heartbeat or also the time period between the two characteristic times H_Zp(x), H_Zp(x+1) of two consecutive heartbeats or the time interval between two signal peaks, e.g., the P peak and the T peak, of the segment Abs.x, Abs.y, . . . of the sum signal SigSum, which occurs in the course of a single heartbeat, is optionally taken into consideration. This time period may remain the same over time or vary over time. A heartbeat time period detector 11 analyzes the sum signal SigSum and the detected heartbeat times H_Zp(x), H_Zp(x+1), . . . and calculates the time interval between two consecutive heartbeat times.
A heartbeat time detector 7 detects, in turn, the respective time H_Zp(n) (n=1, 2, . . . ) of each heartbeat. A signal estimating unit (a signal representation unit) 6 receives the measured values from the two sensors 3 and 10 and calculates from these the respective current value, which the transmission channel parameter LF or Pos assumes at the heartbeat time H_Zp(x).
Depending on the measured value for the lung filling level, LF, and optionally on the measured value for the relative position Pos during a heartbeat, the signal estimating unit 6 calculates during the use phase Np a respective estimated signal segment SigHz,kar.LF of the cardiogenic signal Sigkar in the course of this heartbeat for each heartbeat, wherein the estimated signal segment SigHz,kar.LF depends on the lung filling level, LF, at this heartbeat as well as optionally on the position Pos of the measuring electrode 2.1 or 2.2 and/or on the time interval RR between two heartbeats. This signal segment SigHz,kar.LF estimated as a function of at least one transmission channel parameter, varies, as a rule, from one heartbeat to the next. The estimated signal segments SigHz,kar.LF are combined into the estimated cardiogenic signal Sigkar,est with the use of the heartbeat times.
In one embodiment, each estimated signal segment SigHz,kar.LF has the same length. The intermediate space in the estimated signal Sigkar,est is bridged over by a smoothing. In another embodiment, the respective time period H_Zr(x), H_Zr(x+1), . . . of each heartbeat is measured during the use phase Np, and the estimated signal segment SigHz,kar.LF is adapted to this heartbeat time period by stretching or compression.
The signal estimating unit 6 has in one embodiment reading access to a predefined standard reference signal segment SigHz,Ref, which is stored in a library. This segment describes an average segment of the cardiogenic signal Sigkar in the course of a single heartbeat. This standard reference signal segment SigHz was generated, for example, in advance by measurements on different patients. It contains at least one and preferably a plurality of shape parameters, which change the geometric shape of the reference signal segment SigHz,Ref. The influence of a transmission channel parameter is taken into account indirectly by at least one shape parameter, which will be described farther below.
Examples of shape parameters are, cf.
Due to a respective shape parameter value each being introduced into the shape parameter or each shape parameter of this standard reference signal segment SigHz,Ref, a parameterized cardiogenic estimated signal segment SigHz,kar.LF is generated, which describes the estimated cardiac activity in the course of an individual heartbeat and depends in this example on the lung filling level, LF, and optionally on the position Pos. This parameterized standard reference signal segment SigHz,kar.LF is shown in
These shape parameter values depend in the example according to
In one embodiment, the signal estimating unit 6 calculates for each shape parameter of the standard reference signal segment SigHz,Ref and for each detected heartbeat a respective shape parameter value each, which the shape parameter assumes at the heartbeat time H_Zp(x) or during the heartbeat time period H_Zr(x). Using these shape parameter values, the signal processing unit 5 generates during the use phase from the standard reference signal segment SigHz,Ref for each heartbeat an estimated signal segment SigHz,kar.LF, which is adapted to the current value of the lung filling level, LF, and optionally to the current position Pos and/or other transmission channel parameter, and which describes the expected or estimated cardiogenic signal Sigres in the course of this heartbeat. This is carried out for each heartbeat detected during the use phase Np.
In another embodiment, the signal estimating unit 6 determines in a library 12 a stored reference signal segment SigHz,kar,LF.1 or . . . SigHz,kar,LF.4, which segment is associated with this lung filling level, LF.1, . . . , LF.4 and optionally to this position Pos. The signal estimating unit 6 yields the estimated signal segment SigHz,kar,LF for a heartbeat as a function of the determined reference signal segment or of each determined reference signal segment. No standard reference signal segment SigHz,ref is needed after the end of the training phase Tp in this embodiment. The reconstructing unit 8 combines, in both embodiments during the useful segment Np, the estimated cardiogenic signal segments SigHz,kar,LF in the course of a respective heartbeat each into an estimated cardiogenic signal Sigkar,est and uses for this the heartbeat times H_Zp(x), H_Zp(x+1), . . . , which the time detector 7 has detected. According to the embodiment of the present invention, the reconstructing unit 8 combines the estimated signal segments SigHz,kar,LF adapted to the current filling levels of the lungs, LF, into the reconstructed cardiogenic signal Sigkar,est. This is preferably repeated continually as soon as a new heartbeat is detected.
A plurality of variants of the process according to the present invention will be described below, as it is illustrated by
The R peak is used in one embodiment as a characteristic time H_Zp(n) of a heartbeat. In addition, the following geometric parameters are illustrated:
The R-R interval RR correlates with the pulse of the patient P.
The adaptation phase Ap overlaps the use phase Np. The time is plotted from left to right on the respective x axis of each signal. The time curves of the following signals are shown:
In this variant, the sum signal SigSum is generated by analyzing electrical measured values of the measuring electrodes 2.1 and 2.2. The volume flow Vol′ is measured, for example, by means of the pneumatic sensor 3, and the current lung filling level, LF, is derived from the respective volume flow Vol′ at a plurality of times. Four classes of filling levels of the lungs, namely, LF.1 (lungs almost empty, lung filling level below a first limit), LF.4 (lungs almost full, lung filling level above a second limit) and two filling levels of the lungs in between, LF.2 and LF.3, are distinguished in the example shown. A different number of classes of filling levels of the lungs, LF, and of other transmission channel parameters may, of course, be distinguished as well. The signal with the time curve that indicates the class to which the current lung filling level, LF, belongs, is designated by LF_cl in
In an embodiment of the example according to
The class to which the lung filling level, LF(t), belongs at the time t=H_Zp (n) is determined during the use phase Np in one embodiment at the characteristic heartbeat time H_Zp(n) of the nth heartbeat. In one embodiment, the respective reference signal segment which is assigned to this class in the library 12, is used as the estimated signal segment SigHz,kar,LF(n). It describes the segment of the cardiogenic signal in the course of this heartbeat. For example, the reference signal segment Sigliz,kar,LF.3 for the lung filling level LF.3 is selected for the time H_Zp(y) and is used as an estimated signal segment SigHz,kar,LF(y), and the reference signal segment SigHz,kar,LF.4 for the lung filling level LF.4 is used for the time H_Zp(z) as the estimated signal segment SigHz,kar,LF(z).
In another embodiment, the signal processing unit 5 calculates for each class of lung filling levels, in addition to the reference signal segment, a respective reference parameter value each, for example, as a weighted mean value or as a center or median of the transmission channel parameter values (here: lung filling levels) of this class. For example, the relative frequencies of transmission channel parameter values are used as weighting factors. The signal processing unit 5 determines during the use phase Np for each heartbeat the two reference parameter values that are closest to the transmission channel parameter value of this heartbeat and calculates the estimated signal segment for this heartbeat by smoothing, for example, an interpolation or regression.
The signal estimating unit 6 consequently yields for each heartbeat time H_Zp(y) an estimated signal segment SigHz,kar,LF(y), which depends on the four possible reference signal segments SigHz,kar,LF.1, . . . SigHz,kar,LF.4. In one embodiment, each estimated signal segment SigHz,kar,LF(y) of the cardiogenic signal is equal to a reference signal segment SigHz,kar,LF.1, . . . , SigHz,kar,LF.4 in the library 12. The estimated signal segment provided depends on which of the four classes LF.1, . . . , LF.4 the lung filling level LF belongs to during this heartbeat.
This is carried out after the training phase Tp, i.e., when the signal estimating unit 6 is generated. A respective standard reference signal segment predefined in advance for each detected heartbeat is preferably used for each heartbeat time before the end of the training phase Tp.
These estimated signal segments SigHz,kar,LF are combined by the reconstructing unit 8 into the estimated cardiogenic signal Sigkar,est. This estimated cardiogenic signal Sigkar,est as well as the estimated respiratory signal Sigres,est are shown below the curve LF_cl in
In a possible embodiment, this leads with four classes LF.1, . . . , LF.4 for the lung filling level LF and for two classes for the exhalation time (yes and no, i.e., breathing begins and breathing does not begin shortly before the Q wave) to a total of 2×4=8 different classes. By contrast, only four classes are used in the embodiment shown. The possible values for the lung filling level LF are grouped into three classes LF.a, LF.b, LF.c. In connection with the event that the exhalation time is not close to Q, this leads to three classes LQ.a, LQ.b, LQ.c. In addition, a fourth class Q.d is introduced, namely, that the exhalation time is close to Q, regardless of how the lung filling level LF is. Furthermore, the time curve of the belonging to one of these four classes LF.a, LF.b, LF.c, Q.d is shown in
The sum signal SigSum is a pressure signal in this variant, which is measured in or in front of the esophagus Sp (esophagus) of the patient P, for example, with a probe or with a balloon in the esophagus Sp. The pressure signal could also be the pressure Paw at the transition from a hose of the ventilator 1 to the mouth of the patient P, which is measured by the sensor 3. This pneumatic sum signal SigSum results from a superimposition of the pneumatic respiratory signal Sigres brought about by the breathing activity to a pneumatic cardiogenic signal Sigkar brought about by the cardiac activity.
In the variant shown, the signal processing unit 5 therefore additionally carries out a detrending in the training phase Tp and therefore in the training path Tpf. As a result, the risk that different trends would distort the combination of the sum signal segments arranged in correct time into a reference signal segment is reduced. Both the sum signal SigSum and the processed sum signal SigSum,DT generated by the detrending are shown in
An embodiment of generating the detrending is the following: The signal processing unit determines for each heartbeat the sum signal segment Abs.w, Abs.x belonging to this heartbeat. It calculates a fitted curve, especially a fitted curve through this sum signal segment Abs.w, Abs.x. This fitted curve is generated, for example, by interpolation or as a straight line from the chronologically first to the chronologically last signal value of the sum signal segment Abs.w, Abs.x. The respective fitted curve is subtracted for each heartbeat from the sum signal segment Abs.w, Abs.x. The remaining residue, i.e., the difference, forms the processed sum signal segment Abs_DT.w, Abs_DT.x generated by the detrending. Each sample element comprises such a processed sum signal segment. These segments yield the estimated signal segments SigHz,kar,LD(y), SigHz,kar,LF(z) which are combined into the processed sum signal SigSum,DT.
The signal estimating unit 6 yields during the use phase a respective processed sum signal segment Abs_DT.w, Abs_DT.x for each detected heartbeat.
The signal estimating unit 6 yields a respective estimated signal segment SigHz,kar,LQ, which is selected among four possible reference segments SigSigHz,kar,LQ.a, . . . , SigSigHz,kar,Q.d of the cardiogenic signal Sigkar, during the use phase Np for each heartbeat in one embodiment in the variant according to
In the variant according to
In the variant according to
The sum signal SigSum is broken down in the example shown into a respective signal component per predefined frequency range in both the training phase Tp and the use phase Np. For example, a wavelet transformation or a band filter or a low-pass filter or a high-pass filter is employed.
The respective maximum and the respective minimum are detected in the course of a heartbeat in the signal component SigSum,high during the training phase Tp. For example, two maxima Max.1 and Max.8 are shown. The same is carried out for the minima. For example, a minimum Min.1 is shown. These maxima are divided into four classes of maxima in the respective heartbeat depending on the particular lung filling level LF.1, . . . , LF.4.
The signal estimating unit 6 determines the sum signal segment Abs.x, the higher-frequency signal component segment and the respective lung filling level for each heartbeat in the use phase. The signal estimating unit 6 determines a respective averaged maximum as well as a respective averaged minimum, for which the signal estimating unit 6 uses the measured lung filling level LF at this heartbeat as well as the maxima and minima determined in the library 12. The signal estimating unit 6 cuts off by calculation the components that are above the averaged maxima or below the averaged minima in the segment of the higher-frequency signal component SigSum,high that belongs to this heartbeat. These components originate with certainty essentially from the cardiogenic signal Sigkar and contain no respiratory components that is to be taken into account. The cutting off is illustrated in
Averaged maxima and averaged minima are used in this example as shape parameter values of a class of transmission channel parameter values (here: lung filling level end LF.1, . . . , LF.4). These shape parameter values are used in this variant to parameterize a predefined change rule. The parameterized change rule changes a respective segment Abs.x, Abs.y of the sum signal SigSum—in this variant: a segment of the higher-frequency signal component SigSum,high. The change comprises in this variant the step of cutting off signal components above the maxima and below the minima.
It is also possible to use additional or other arithmetic shape parameters and hence other change rules, e.g., averaged first and/or second derivatives. It is also possible to use weighting factors and/or a “soft threshold.” A segment of the sum signal SigSum, which segment belongs to a heartbeat, or of a signal component in the segments in which the slope of the sum signal SigSum is below a predefined limit, is stretched in another embodiment. Due to the variant shown in
In a preferred application of the variant that is shown by
In the embodiments just described, the signal processing unit 5 receives a plurality of measured values from at least one sensor, wherein this sensor is not a sum signal sensor 1, 2.1, 2.2, 3, 4, and it generates by signal processing the transmission channel parameter value or each transmission channel parameter value from these measured values. It is also possible that the signal processing unit 5 calculates the value of at least one transmission channel parameter and measures it by the calculation by the signal processing unit 5 analyzing the sum signal SigSum. Another sensor for the transmission channel parameter is thus unnecessary for this transmission channel parameter.
Possible transmission channel parameters, which can be measured by calculation and without a separate physical sensor, are shown in
An embodiment of this variant will be explained below with reference to
These standardized signal segments Abs_std.x, Abs_std.y, . . . are arranged in a matrix M. Each row of this matrix represents a heartbeat and each column a scanning time. The signal processing unit applies to the set of these standardized signal segments in a first part Tpf.1 of the training path Tpf a singular value decomposition (SVD) or also a principal component analysis (PCA). This step yields a plurality of reference curves in a decreasing order, wherein the order depends decreasingly on an agreement value. The first reference curve V.1 agrees with the standardized signal segments most strongly, etc. The three most important reference curves V.1 through V.3 are shown in
The reference curves V.1, V.2 are predefined in an alternative embodiment.
The signal processing unit 5 classifies next the standardized sum signal segments Abs_std.x, Abs_std.y, i.e., in a second part Tpf.2 of the training path Tpf. Only the two most important reference curves V.1 and V.2 are used for this in the example being shown. It is also possible to use more than two reference curves. The signal processing unit 5 calculates for each sum signal segment Abs_std.x, Abs_std.y, . . . a respective value each for the agreement between this standardized sum signal segment and the reference curve V.1, V.2 used. For example, it calculates the scalar product between the standardized sum signal segment Abs_std.x, Abs_std.y, . . . and the reference curve V.1, V.2. The time course ÜM.1 of the individual agreement value for the first reference curve V.1 and the time course ÜM.2 of the individual agreement value for the second reference curve V.2 are shown in
In a number of the versions just shown each sample element comprises a respective sum signal segment or a processed sum signal segment. Depending on the calculated value or calculated values of the transmission channel parameters used, the signal processing unit 5 combines in the training phase Tp the sample elements into classes. The signal processing unit 5 generates for each class a respective reference signal segment, e.g., the four reference signal segments SigHz,kar,LF.1, . . . , SigHz,kar,LF.4 or SigHz,kar,ÜM.a, . . . , SigSigHz,kar,ÜM.d. Different processes are possible for combining the sum signal segments of a class of sample elements into a reference signal segment, which will then be stored in the library 12.
Time, more precisely, a plurality of relative scanning times, are plotted om the x axis. “Relative” means relative to the beginning of the signal segment. The transmission channel parameter used or a transmission channel parameter used, in this example the R-R interval RR between the R peaks of two consecutive heartbeats, is plotted on the y axis. This process can just as well be used for other transmission channel parameters with two numbers as the parameter values and also for a plurality of transmission channel parameters. The value range of the transmission channel parameter plotted on the y axis is divided in this example into more than ten classes, and in the extreme case up to the accuracy of the machine, i.e., one class per number that can be displayed on the signal processing unit 5 used. The signal value, i.e., the value of the sum signal at this scanning time and at this transmission channel parameter value is plotted on the z axis. The sum signal segments of the sample elements were standardized in advance, so that the standardized sum signal segments Abs_std.x, Abs_std.y have all the same length and the R peaks have the same relative scanning time. These sum signal segments are represented one on top of another with the correct time in the view shown in
The signal processing unit 5 calculates in the training phase Tp a fitted curve, which extends in the y-z plane, for each scanning time (x axis) by smoothing. This is illustrated in
The time curve of the input signal E_SigSum, which is generated from electrical measured values of the measuring electrodes 2.1 and 2.2 and results from a superimposition of the heartbeat activity and the breathing activity of the patient P, is shown in the topmost row in
In the row H_Zp under the input signal E_SigSum, the respective beginning, on the one hand, as well as the respective QRS segment of each heartbeat, for example, the beginning Anf_Zp(x) and the QRS segment H_Zp(x) of the xth heartbeat, are shown. The respective QRS segment acts in one embodiment as the characteristic heartbeat time.
The sum signal SigSum is subjected to a wavelet transformation, and different frequency ranges are predefined. The wavelet transformation yields a respective signal component for each predefined frequency range. Three signal components A through C are calculated in the example being shown, and more than three signal components are preferably calculated. A respective other process, which will be described below, is carried out for each signal component A through C.
The EMG power (power of the respiratory signal), which is illustrated in
A respective limit each is determined for each class in the training phase, i.e., a total of three limits Max_Pow1 (for EMG_Pow1), Max_Pow2 (for EMG_Pow2) and Max_Pow3 (for EMG_Pow3) are determined. The row shows the application in the use phase. The cardiogenic component in the signal component A shall be determined. Designated by SigSum,A in the signal component A, the values whose respective absolute value is above the respective limit Max_Pow1, Max_Pow2, Max_Pow3 are used as values belonging to the cardiogenic component. Which threshold value it is depends on the current EMG power. The other signal values are set at zero by calculation.
The signal power is calculated from the signal component, e.g., by calculating the effective value (root mean square). This calculation yields a time curve of the signal power. A respective power curve segment is calculated for each heartbeat. Depending on the lung filling level LF.1 or LF.2 or LF.3 at this heartbeat, a power curve segment SigHz,Pow.LF.1 or SigHz,Pow.LF.2 or SigHz,Pow.LF.3 is calculated hereby at this heartbeat.
The power curve segments for a lung filling level class LF.12 or LF.2 or LF.3 are placed one on top of another with the correct time. The segments of one class placed one on top of another are combined, for example, averaged. As a result, a standard power curve segment is formed for each class. The three standard power curve segments SigHz,Pow,LF.1 and SigHz,Pow,LF.2 and SigHz,Pow,LF.3 calculated in this manner are shown in the lower row of
The limit Max_Pow.LF.n is then calculated depending on this median, for example, according to the formula
Max Pow.LF.n=min(α*Median Pow.LF.n,β+γ*Median Pow LF.x/SigHz,Pow,LF.n).
Where, α, β and γ are predefined constants, for example, α=6, β=0.01 and γ=0.05.
These limits Max_Pow.LF.1, Max_Pow.LF.2 and Max_Pow.LF.3 are the result of the training phase Tp in this approach.
Only the values of the signal component C that belong to the cardiogenic signal, which are above the limit for the respective lung filling level class, are used again in the use phase Np.
Depending on the respective lung filling level class LF.1 or LF.2 or LF.3, the respective limit Max_Pow.LF.1 or Max_Pow.LF.2 or Max_Pow.LF.3 is entered.
The respective cardiogenic component in the three signal components A, B and C are combined into an estimated cardiogenic signal Sigkar,est. This estimated cardiogenic signal Sigkar,est is shown in the third row. The difference from the sum signal SigSum and from the estimated cardiogenic signal Sigkar,est yields the estimated respiratory signal Sigres,est, which is shown in the fourth row.
It is possible to use an additional transmission channel parameter, namely, the instantaneous EMG power, as this was explained for the signal component B in reference to
Two characteristic heartbeat times, namely, the maximum value of the P peak and the maximum value of the QRS area, are detected for each heartbeat in the signal component C, likewise designated by SigSum,C. These terms were explained with reference to
Two histograms, namely, a histogram Hist_P for the maximum P values and a histogram Hist_QRS for the maximum QRS values, are calculated from these maximum values. The signal value is shown on the x axis and the percentage frequency on the y axis.
Using these two histograms Hist_P and Hist_QRS, three limits, which are variable over time, are again calculated for the three classes LF.1, LF.2, LF.3. These limits are designated by Max_PQRS.LF.1, Max_PQRS.LF.2 and Max_PQRS.LF.3.
A mean value Mean_QRS.LF.x for the class LF.n is calculated by averaging over all maxima Max_QRS(x) of the QRS segments of all heartbeats, which belong to the class LF.n, arithmetically or in another manner. A mean value Mean_P.LF.x, in which averaging is carried out over all maxima Max_P(x) of the P peaks of all heartbeats, which belong to the class LF.n, is correspondingly calculated for the class LF.n. These six mean values are shown in
A predefined limit is used at the beginning of the use phase Np. As soon as a sufficient number of heartbeats are detected, two different limits are used for each class LF.1, LF.2, LF.3, namely,
α1−β1*Mean_P.LF.x and
α2−β2*Mean_QRS.LF.x
The four predefined constants have, for example, the values α1=0.05, β1=0.5, α2=0.025 and β2=0.05.
While specific embodiments of the invention have been shown and described in detail to illustrate the application of the principles of the invention, it will be understood that the invention may be embodied otherwise without departing from such principles.
S_Q Signal, which describes another transmission channel parameter, namely, whether the exhalation by the patient begins shortly before the Q wave or not
| Number | Date | Country | Kind |
|---|---|---|---|
| 10 2019 006 866.1 | Oct 2019 | DE | national |
This application is a United States National Phase Application of International Application PCT/EP2020/073826, filed Aug. 26, 2020, and claims the benefit of priority under 35 U.S.C. § 119 of German Application 10 2019 006 866.1, filed Oct. 2, 2019, the entire contents of which are incorporated herein by reference.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/EP2020/073826 | 8/26/2020 | WO |
