Patent Document

BACKGROUND OF THE INVENTION 
     1. Field of the Invention 
     The invention relates to a target driven inspiratory assist ventilation system. 
     2. Brief Description of the Prior Art 
     The physiological mechanisms which generate myoelectrical activity when a muscle contracts have been known and understood for a long time. In particular, how to record signals from the muscles is one of the most extensively, theoretically described topics in physiology. Although the theoretical understanding is impressive, the bio-physiological application of these theories is, in practice, still deficient. As an example, no standardized analysis procedure has been developed for recording signals produced by activation of several, different motor units, the so called interference wave pattern. The interference wave pattern signal (EMG signal) contains an immense quantity of bio-physiological information about the given neuro-muscular function. However, as this EMG signal is very low in amplitude, it is sensitive to numerous artifacts. The influence of these artifacts varies in relation to the configuration of recording electrodes, the digitizing rate of the signal, and the type of recording technique. 
     Prior art analysis of interference wave pattern signals usually comprises a time consuming, tedious manual determination of the quality of the signal through visual inspection of this signal in the time domain. This determination is performed by a “subjective” investigator. Most of the prior art references describe how to calculate comparison estimates, but present very few comments on the signal quality. It is therefore not surprising to find that, in this technical field, independent studies evaluating the same questions have lead to different or even contradictory results. 
     Also in the prior art, the patient&#39;s inspiratory flow and volume has been used to control inspiratory proportional pressure assist ventilation. Proper adjustment of the relative contribution of flow and volume support during the inspiration requires knowledge of the elastic and viscous properties of the patient&#39;s respiratory system. Since the elastic and viscous properties may change, these measurements must be repeated at regular intervals. Correct and repeated measurements of elastance and resistance are difficult to set up in an intensive care unit. Moreover, in the presence of intrinsic positive end-expiratory pressure, the flow-volume controlled proportional assist ventilation may fail to trigger during whole breaths, and will definitively fail to trigger during at least the initial part of the inspiration which precedes the onset of flow; this period can last up to 300 ms in the case of a patient suffering from obstructive pulmonary disease. Finally leakage in the system will influence and may disturb the performance of the flow controlled proportional assist ventilation. 
     Traditionally, the goal of mechanical ventilation has been to maintain an optimal minute ventilation and respiratory load, and therefore, has included specific measurements of inspiratory flow and tidal volume. New concepts in mechanical ventilation allow patients to take over the control of ventilatory support delivered, both in terms of magnitude and duration. New technology has also incorporated new methods of applying ventilatory assist for example, mask ventilation, uncuffed endotracheal tubes, and miniature endotracheal tubes. These devices frequently cause leakage of gases such that measurement of flow and volume become erroneous. 
     Current technology is therefore often limited in its ability to detect and correct for these gas leaks and patients are at risk of becoming hyper- or hypo-ventilated. 
     OBJECTS OF THE INVENTION 
     An object of the present invention is therefore to overcome the above described drawbacks of the prior art. 
     Another object of the present invention is to provide a method and a device capable of adjusting the degree of inspiratory assist in relation to the real need of the patient, i.e. only to compensate for the degree of incapacity of the patient. 
     A further object of the present invention is to provide a method and a device for controlling inspiratory proportional pressure assist ventilation which requires no knowledge of the elastic and viscous properties of the patient&#39;s respiratory system, is not influenced by intrinsic positive end-expiratory pressure, altered muscle function, and is not influenced by air leakage of the lung ventilator unless the leakage exceeds the pumping capacity of the ventilator. 
     SUMMARY OF THE INVENTION 
     More specifically, in a preferred embodiment of the invention, there is provided a gain controller for adjusting, in relation to a target drive signal, the value of a gain applied to a signal representative of respiratory drive output of a patient during inspiration, to produce an amplified respiratory drive output representative signal for controlling a lung ventilator assisting inspiration of the patient. The gain controller comprises:
         a first input for receiving the signal representative of respiratory drive output having a first amplitude;   a second input for receiving the target drive signal of a second amplitude;   a comparator for determining whether the amplitude of the signal representative of respiratory drive output is higher or lower than the amplitude of the target drive signal; and   a gain adjustment unit for increasing the value of the gain when the amplitude of the signal representative of respiratory drive output is higher than the amplitude the target drive signal and for decreasing the value of the gain when the amplitude of the signal representative of respiratory drive output is lower than the amplitude of the target drive signal.       

     In another embodiment of the invention, there is provided a method for adjusting, in relation to a target drive signal, the value of a gain applied to a signal representative of respiratory drive output of a patient during inspiration, to produce an amplified respiratory drive output representative signal for controlling a lung ventilator assisting inspiration of the patient. The method comprises:
         receiving the signal representative of respiratory drive output having a first amplitude;   receiving the target drive signal of a second amplitude;   determining whether the amplitude of the signal representative of respiratory drive output is higher or lower than the amplitude of the target drive signal; and   increasing the value of the gain when the amplitude of the signal representative of respiratory drive output is higher than the amplitude of said target drive signal; and decreasing the value of the gain when the amplitude of the signal representative of respiratory drive output is lower than the amplitude of the target drive signal.       

     Target drive ventilation is based on the assumption that the patient&#39;s respiratory centers are intact and the patient is able to control minute ventilation as long as he/she has sufficient respiratory muscle. In a preferred embodiment of the invention, determination of respiratory drive is made by measuring the electrical activation of the diaphragm during an inspiration. Of course, any other signal representative of respiratory drive output may be used in other embodiments of the invention. Electrical activity of the diaphragm has previously been demonstrated to reflect global respiratory drive. The inspiratory electrical activation of the diaphragm can be quantified as the mean, median, total, peak, etc. and the trend of the previous breaths is used to adjust ventilatory assist for the present breath. 
     The invention is aimed to control ventilatory assist levels in order to maintain the respiratory drive (determined by diaphragm electric activation) at a sustainable target level. The lung ventilator can use a pressure/flow/volume generating device with a control unit which operates to maintain the mean (could also be median/peak/total, etc.) pressure/flow/volume in the ventilatory line sufficient for maintaining a constant target diaphragm electrical activity. The diaphragm electrical activity during a breath will be calculated in order to determine the mean (could also be median/peak/total, etc.) neural drive to the diaphragm for that particular breath. The trend for respiratory drive can be obtained from diaphragm electrical activity of previous breaths such that one can determine whether respiratory drive increases, decreases, or remains constant. A trend for a change in diaphragm electrical activity indicating an increase in respiratory drive will result in a progressive increase ventilatory assist until diaphragm electrical activity, i.e., respiratory drive has returned to its target level. Similarly, the decrease in diaphragm electrical activity, indicating reduced respiratory drive, will produce a progressive decrease in ventilatory assist until diaphragm electrical activity i.e. respiratory drive has returned to its target level. 
     Target Drive Ventilation would be more efficiently used in combination with Neurally Adjusted Proportional Pressure Assist (U.S. Pat. No. 5,820,560 to Sinderby et al., 1998), where ventilatory assistance will be proportional to the patient&#39;s respiratory drive throughout the breath and the average respiratory drive would remain constant over time. For proportional assist ventilation or other modes which deliver varying levels of support, the increasing or decreasing levels of mean/total ventilatory assist will be adjusted by increasing or decreasing the gain factor applied in the respective functions. 
     Target drive ventilation can also be applied with other modes of ventilatory assist. For use with ventilatory support modes that provide constant levels of support, for example pressure support, the increasing or decreasing levels of mean/total ventilatory assist will be achieved by relative increases or decreases of the pressure support, the increasing or decreasing levels of mean/total ventilatory assist will be achieved by relative increases or decreases of the pressure support level. Extreme pressure support levels will be avoided by introducing safety limits. 
     The advantage of Target drive ventilation is that this mode of ventilation does not depend on flow or volume measurements. A leaky ventilatory line will introduce a change in respiratory drive which will change the ventilatory assist in order to return the respiratory drive to its target level. Also, changes in the patient&#39;s metabolic or patho-physiological status which result in altered respiratory drive will be compensated. In contrast with present methods of controlling mechanical ventilators, an increase in respiratory assistance using a signal representative of respiratory drive output (e.g., an EMG signal) does not affect the efficiency with which these signals reliably control the ventilator (unless of course the disease affects the neuro-muscular function). 
     A combination of Target Drive Ventilation and Neurally Adjusted Proportional Pressure Assist (U.S. Pat. No. 5,820,560), would provide partial correction for leaks within breaths and compensation for leaks over long periods of time. The use of neural triggers would also overcome issues related to intrinsic PEEP. 
     The objects, advantages and other features of the present invention will become more apparent upon reading of the following non restrictive description of a preferred embodiment thereof, given by way of example only with reference to the accompanying drawings. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
       In the appended drawings: 
         FIG. 1  is a schematic representation of a set-up of an EMG analysis system; 
         FIG. 2  is a section of oesophageal catheter on which an array of electrodes of the EMG analysis system of  FIG. 1  is mounted; 
         FIG. 3  illustrates a section of oesophageal catheter on which a second embodiment of the array of electrodes is mounted; 
         FIG. 4  is a graph showing a set of EMGdi signals of the diaphragm detected by pairs of successive electrodes of the array of  FIG. 2 ; 
         FIG. 5A  is a flow chart showing a method for conducting double subtraction technique of the EMGdi signals; 
         FIG. 5B  is a flow chart showing a method for controlling a gain value in accordance with an embodiment of the invention; 
         FIG. 5C  is a flow chart showing an illustrative embodiment of a method according to the invention for controlling the level of a ventilatory support control signal for application to and control of a lung ventilator; 
         FIG. 6  is a graph showing the distribution of correlation coefficients calculated for determining the position of the center of the depolarizing region of the diaphragm along the array of electrodes of  FIG. 2 ; 
         FIG. 7  is a schematic diagram illustrating in the time domain a double subtraction technique for improving the signal-to-noise ratio and to reduce an electrode-position-induced filter effect; 
         FIG. 8   a  is a graph showing the power density spectrum of electrode motion artifacts, the power density spectrum of ECG, and the power density spectrum of EMGdi signals; 
         FIG. 8   b  is a graph showing an example of transfer function for a filter to be used for filtering out the electrode motion artifacts, ECG, and the 50 or 60 Hz disturbance from electrical mains; 
         FIG. 9  is a schematic diagram illustrating in the frequency domain stabilization by the double subtraction technique of the center frequency upon displacement of the center of the depolarizing region of the diaphragm along the array of electrodes of  FIG. 2 ; 
         FIG. 10  is a schematic block diagram of a lung ventilator showing control of inspiratory proportional pressure assist ventilation by means of an EMG signal obtained with the above mentioned double subtraction technique; and 
         FIG. 11  is a schematic block diagram showing a structure for implementing the steps of the method described in  FIG. 5B . 
     
    
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT 
     To measure EMG activity of the diaphragm  11  (EMGdi) of a human patient  14 , an array of electrodes such as  12  ( FIGS. 1 and 2 ) are mounted on the free end section  15  of an oesophageal catheter  13 , with a constant inter-electrode distance d ( FIG. 2 ). As shown in  FIG. 1 , the catheter  13  is introduced into the patient&#39;s oesophagus through one nostril or the mouth until the array of electrodes  12  are situated at the level of the gastro esophageal junction. The diaphragm  11  and/or the oesophagus slightly move during breathing of the patient  14  whereby the array of electrodes  12  also slightly moves about the diaphragm  11 . As will be explained in the following description, automatic compensation for this displacement is provided. 
     To mount an electrode  12  on the free end section  15  of the catheter  13 , stainless steel wire (not shown) may be wound around the catheter  13 . The wound stainless steel wire presents a rough surface smoothed out by solder, which in turn is electroplated with nickel, copper and then gold or silver. Of course, EMG signals from other muscles and other constructions of electrodes can be implemented. 
     Electric wires (not shown) interconnect each pair of successive electrodes such as  1 – 7  ( FIG. 2 ) with a respective one of a group of differential amplifiers  16 . This defines an overlap array. Obviously, these electric wires follow the catheter  13  from the respective electrodes  12  to the corresponding amplifiers  16 , and are preferably integrated to the catheter  13 . Preferably, the electric wires transmitting the EMGdi signals collected by the various pairs  1 – 7  of electrodes  12  are shielded to reduce the influence of external noise, in particular disturbance from the 50 or 60 Hz current and voltage of the electrical mains. 
     The group of differential amplifiers  16  amplifies (first subtraction step of the double subtraction technique) and band-pass filters each EMGdi signal. This first subtraction step may also be carried out in the personal computer  19  when the amplifiers  16  are single-ended or equivalently designed amplifiers (monopolar readings). 
     In the example illustrated in  FIGS. 1 and 2 , the free end section  15  of the catheter  13  is provided with an array of eight electrodes  12  defining seven pairs  1 ,  2 ,  3 ,  4 ,  5 ,  6  and  7  of successive electrodes  12  respectively collecting seven different EMGdi signals. Although it has been found that EMG activity of the diaphragm (EMGdi) can be measured accurately with an oesophageal catheter  13  provided on the free end section  15  thereof with an array of eight electrodes  12 , a different number and/or configuration of pairs of electrodes  12  can be contemplated depending on the patient&#39;s anatomy and movement of the diaphragm. Also, the pairs  1 – 7  do not need to be pairs of successive electrodes;  FIG. 3  illustrates an array of nine electrodes to form seven overlapping pairs of electrodes  1 – 7 . 
     A major problem in recording EMGdi signals is to maintain the noise level as low and as constant as possible. Since the electric wires transmitting the EMGdi signals from the electrodes  12  to the differential amplifiers  16  act as an antenna, it is crucial, as indicated in the foregoing description, to shield these electric wires to thereby protect the EMGdi signals from additional artifactual noise. Also, the package enclosing the differential amplifiers  16  is preferably made as small as possible (miniaturized) and is positioned in close proximity to the patient&#39;s nose to decrease as much as possible the distance between the electrodes  12  and the amplifiers  16 . 
     The amplified EMGdi signals are supplied to a personal computer  19  through respective isolation amplifiers of a unit  18 . Unit  18  supplies electric power to the various electronic components of the differential and isolation amplifiers while ensuring adequate isolation of the patient&#39;s body from such power supply. The unit  18  also incorporates bandpass filters included in the respective EMGdi signal channels to eliminate the effects of aliasing. The EMGdi signals are then digitally processed into the personal computer  19  after analog-to-digital conversion thereof. This analog-to-digital conversion is conveniently carried out by an analog-to-digital converter implemented in the personal computer  19 . The personal computer  19  includes a monitor  40  and a keyboard  31 . 
     It is believed to be within the capacity of those of ordinary skill in the art to construct suitable differential amplifiers  16  and an adequate isolation amplifiers and power supply unit  18 . Accordingly, the amplifiers  16  and the unit  18  will not be further described in the present specification. 
     An example of the seven EMGdi signals collected by the pairs  1 – 7  of successive electrodes  12  ( FIGS. 1 and 2 ) and supplied to the computer  19  is illustrated in  FIG. 4 . 
     As the diaphragm is generally perpendicular to the longitudinal axis of the oesophageal catheter  13  equipped with an array of electrodes  12 , only a portion of the electrodes  12  are situated in the vicinity of the diaphragm. It is therefore important to determine the position of the diaphragm with respect to the oesophageal electrode array. 
     The portion of the crural diaphragm  11  which forms the muscular tunnel through which the oesophageal catheter  13  is passed is referred to the “diaphragm depolarizing region” (DDR). The thickness of the DDR is 20–30 mm. It can be assumed that, within the DDR, the distribution of active muscle fibers has a center from which the majority of the EMGdi signals originate, i.e. the “diaphragm depolarizing region center” (DDR center). Therefore, EMGdi signals detected on opposite sides of the DDR center will be reversed in polarity with no phase shift; in other words, EMGdi signals obtained along the electrode array are reversing in polarity at the DDR center. 
     Moving centrally from the boundaries of the DDR, EMGdi power spectrums progressively attenuate and enhance in frequency. Reversal of signal polarity on either side of the electrode pair  4  with the most attenuated power spectrum confirms the position from which the EMGdi signals originate, the DDR center. 
     Referring to  FIG. 5A , the first task of the computer  19  is to determine the center of the DDR. The center of the DDR is repeatedly determined at predetermined time intervals. 
     For that purpose, slow trend is first removed from each EMGdi signal (step  500 ). To carry out such trend removal, the processing conducted by the computer  19  on each EMGdi signal is equivalent to high-pass filtering each EMGdi signal at a transition frequency of about 20 Hz. In particular, step  500  will remove the direct current component of the EMGdi signals to enable the computer  19  to evaluate the polarities of the EMGdi signals relative to each other. 
     In step  501 , the EMGdi signals are cross-correlated in pairs. As well known to those of ordinary skill in the art, cross-correlation is a statistical determination of the phase relationship between two signals and essentially calculates the similarity between two signals in terms of a correlation coefficient r (step  502 ). A negative correlation coefficient r indicates that the cross-correlated signals are of opposite polarities. 
       FIG. 6  shows curves of the value of the correlation coefficient r versus the midpoint between the pairs of electrodes from which the correlated EMGdi signals originate. In this example, the inter-electrode distance is 10 mm. Curves are drawn for distances between the correlated pairs of electrodes  12  of 5 mm (curve  20 ), 10 mm (curve  21 ), 15 mm (curve  22 ) and 20 mm (curve  23 ). One can appreciate from  FIG. 5A  that negative correlation coefficients r are obtained when EMGdi signals from respective electrode pairs situated on opposite sides of the electrode pair  4  are cross-correlated. It therefore appears that the change in polarity occur in the region of electrode pair  4 , which is confirmed by the curves of  FIG. 4 . Accordingly, it can be assumed that the center of the DDR is situated substantially midway between the electrodes  12  forming pair  4 . 
     For example, the center of the DDR can be precisely determined by interpolation (step  503  of  FIG. 5A ) using a square law based fit of the three most negative correlation coefficients of curve  21  obtained by successive cross-correlation of the EMGdi signals from each electrode pair to the EMGdi signals from the second next electrode pair. Association of the center of the DDR to a pair of electrodes  12  provides a “reference position” from which to obtain EMGdi signals within the DDR. Such control is essential in overcoming the artifactual influence on the EMGdi power spectrum. 
     It has been experimentally demonstrated that EMGdi signals recorded in the oesophagus are satisfactory as long as they are obtained from electrode pairs (with an inter-electrode distance situated between 5 and 20 mm) positioned at a distance situated between 5 and 30 mm on the opposite sides of the DDR center (the inter-pair distance being therefore situated between 5 and 30 mm). Although EMGdi signals obtained from these positions offers a clear improvement in acceptance rate, the signal-to-noise ratio during quiet breathing still tends to remain unsatisfactorily low. 
     In another embodiment of the invention, step  500  can be eliminated and steps  501 ,  502  and  503  could be implemented immediately after step  505 . 
     For example, in  FIG. 4 , the EMGdi signals originating from the electrode pairs  3  and  5  situated respectively 10 mm below and 10 mm above the DDR are strongly inversely correlated at zero time delay. In contrast to the inversely correlated EMGdi signals, the noise components for electrode pairs  3  and  5  are likely to be positively correlated. Hence, as illustrated in  FIG. 7 , subtraction of the EMGdi signals  24  and  25  from electrode pairs  3  and  5  will result into an addition of the corresponding EMGdi signals (signal  26  of  FIG. 6 ) and into a subtraction, that is an elimination of the common noise components. This technique will be referred to as “the double subtraction technique” (step  504  of  FIG. 5A ). 
     Subtraction step  504  (second subtraction step of the double subtraction technique) can be carried out either in the time domain, or after conversion of signals  24  and  25  in the frequency domain. Double subtraction technique can be performed by subtracting other combinations of signals, for example by subtracting the EMGdi signal from electrode pair  2  from the EMGdi signal from electrode pair  5  ( FIG. 4 ), by subtracting signal from electrode pair  6  from the signal from electrode pair  3  and by adding these differences, etc. Other means for reducing the effect of electrode filtering can be applied. 
     The double subtraction technique is carried out in step  504  on the pair of EMGdi signals (for example the signals from electrode pairs  3  and  5  shown in  FIG. 4 ) identified in step  503 , after appropriate filtering of these EMGdi signals in step  505 . Filtering step  505  will remove from each EMGdi signal the motion artifacts, the electrocardiogram (ECG) component, and the disturbance from the electrical mains. Motion artifacts are induced by motion of the electrodes. More generally, motion artifacts are defined as a low frequency fluctuation of the EMGdi signals&#39; DC level induced by mechanical alterations of the electrode metal to electrolyte interface i.e. changes in electrode contact area and/or changes in pressure that the tissue exerts on the electrode. 
     The graph of  FIG. 8   a  shows the power density spectrum of the above defined electrode motion artifacts, the power density spectrum of ECG, and the power density spectrum of EMGdi signals. The graph of  FIG. 8   b  shows an example of transfer function for a filter (the dashed line showing the optimal transfer function, and the solid line the transfer function implemented by the inventors) to be used in step  505  for filtering out the electrode motion artifacts, ECG, and the 50 or 60 Hz disturbance from the electrical mains. Processing of the EMGdi signals by the computer  19  to follow as closely as possible the optimal transfer function of  FIG. 8   b  will conduct adequately filtering step  505 . 
     Referring back to  FIG. 5A , step  506  calculates the RMS (Root-mean-square) value of the double-subtracted signal produced in step  504 . The increase in amplitude obtained with the double subtraction technique is associated with a twofold increase in RMS values. RMS values obtained with the double subtraction technique are closely and linearly related to the original signals. The RMS value can be replaced by any other value representative of the strength of the double-subtracted signal, for example mean, median, peak or total signal amplitudes. 
     The double subtraction technique compensates for the changes in signal strength and frequency caused by movement of the diaphragm  11  ( FIG. 1 ) and/or the oesophagus during breathing of the patient  14  causing movement of the array of electrodes  12  with respect to the diaphragm  11 . Referring to  FIG. 9 , off center of the array of electrodes  12  (electrode-position-induced filter effect) causes a variation of center frequency values (see curves  27  and  28 ) for the EMGdi signals from the electrode pairs  3  and  5 . The double subtraction technique eliminates such variation of center frequency values as indicated by curve  29  as well as variation of signal strength. Therefore, the reciprocal influence of the position of the DDR center on the EMGdi signal frequency content is eliminated by the double subtraction technique. 
     It has been found that the double subtraction technique may improve the signal-to-noise ratio by more than 2 dB ratio and reduce an electrode-position-induced filter effect. Double subtraction technique is also responsible for a relative increase in acceptance rate by more than 30%. 
     Cross-talk signals from adjacent muscles are strongly correlated at zero time delay and equal in polarity between all pairs of electrodes  12 . Hence, these cross-talk signals appear as a common mode signal for all electrode pairs and therefore, are eliminated by the double subtraction technique. 
     Referring to  FIG. 5B , a target drive signal is set by an operator at step  601 . The output of block  601  is therefore the target drive signal  602 . The value of the target drive signal  602  is determined by a person skilled in the art. The target drive signal  602  can be any signal which is representative of respiratory drive output. In a preferred embodiment of the invention, this signal can be any signal representative of the electrical activity of a muscle (i.e., electromyographic signal) that reflects the global respiratory drive. The target drive signal  602  can therefore be quantified as, among others, the mean, the median, the total or the peak of the EMGdi signal. The target drive signal  602  is then compared to the RMS value on line  508  in block  604 . If the present breath RMS value on line  508  is greater than the target drive signal  602 , this indicates that there is a trend for a change in diaphragm electrical activity (EMGdi) indicating an increase in respiratory drive and requiring a progressive increase in ventilatory assistance. The result of the decision block  604  will be positive until the EMGdi  508  returns to the target level  602 . In this case, the stored gain (k) will be increased in block  606 . The gain (k)  610  will then be recorded in block  611  and outputted as signal  613 . 
     Returning now to block  604 , if the present breath RMS value on line  508  is not greater than the target drive signal  602 , it may be equal or smaller. In block  605 , it is determined if the present breath RMS value  508  is equal to the target drive signal  602 . If it is equal, the gain value (k) does not change. 
     If the present breath RMS value  508  is not equal to the target drive signal  602 , then it is smaller than the target drive signal  602 , and this indicates a decrease in diaphragm electrical activity (EMGdi) resulting in reduced respiratory drive. It will therefore be necessary to progressively decrease the ventilatory assist until the diaphragm electrical activity (EMGdi)  508  returns to the target level  602 . In this case, the result from decision block  604  will be negative resulting in a decrease in the stored gain (k) in block  608 . The gain (k)  609  will then be stored at step  611  and outputted as signal  613 . 
     Those skilled in the art will understand that the amount of the change (increase, step  606 , or decrease, step  608 ) in ventilatory assistance is derived from experience, the patient&#39;s condition, the environment, etc. The amount of change can therefore be adjusted on a case by case basis. Also, in a particular embodiment of the invention, the increase or decrease in ventilatory assistance could be a relative value; that is, a fraction or percentage multiplied by, for example, the target drive signal, the signal representative of respiratory drive output, or a difference between the amplitude of the signal representative of respiratory drive output and the amplitude of the target drive signal. 
     Those skilled in the art will also understand that the test at steps  604  and  605  can include a certain range of amplitudes for the target drive signal  602  (whether they are absolute or relative amplitudes); that is, for example, the target drive signal could be X plus or minus a predetermined value. Therefore, at step  604 , the present breath RMS value  508  needs to be greater than X plus the predetermined value in order to proceed to step  606 . In the same way, at step  605 , the present breath RMS value  508  needs to be smaller than X minus the predetermined value in order to proceed to step  608 . 
     Finally, the present RMS value on line  508  will be multiplied by the gain (k)  613  in block  612  to produce a control signal  614 . The control signal  614  will be the input to lung ventilator  54  of  FIG. 10 . 
       FIG. 11  illustrates a possible physical embodiment of steps  604 ,  605 ,  606 ,  608 ,  611 , and  612  of  FIG. 5B . A gain controller  620  and a gain multiplier  628  are provided. A first input to gain controller  620  is the RMS value on line  508 , a second input to gain controller  620  is target drive signal  602 , and the output of gain controller  620  is gain (k)  613  value. The gain (k)  613  value is then inputted to the gain multiplier  628  where it is multiplied by the RMS value on line  508  for the present inspiration (step  612  of  FIG. 5B ) resulting in a control signal  614 . 
     The gain controller  620  further comprises a comparator  624  and a gain adjustment block  626 . The comparator  624  implements steps  604  and  605  and the gain adjustment block  626  implements steps  606 ,  608  and  611  of  FIG. 5B . 
     Referring to  FIG. 5C , the target drive signal is set by an operator in block  601 . The output of block  601  is therefore the target drive signal  602 . The value of the target drive signal  602  is determined by a person skilled in the art. As indicated in the foregoing description, the target drive signal  602  can be any signal representative of respiratory drive output. In an illustrative embodiment, this signal can be any signal representative of the electrical activity of a muscle (i.e. electromyographic signal) that reflects the global respiratory drive. The target drive signal  602  can therefore be quantified as, among others, the mean, the median, the total or the peak of the EMGdi signal. 
     The target drive signal  602  is then compared to the RMS value on line  508  in block  604 . 
     If the present breath RMS value on line  508  is greater than the target drive signal  602 , this indicates that there is a trend for a change in diaphragm electrical activity (EMGdi) indicating an increase in respiratory drive and requiring an increase in ventilatory assistance. Block  701  then commands block  700  to produce an increase of the ventilatory support, for example pressure support, through an increase of the level of the ventilatory support control signal  614 . The function of block  700  is to produce this ventilatory support control signal  614 . In other words block  701  commands block  700  to increase the level of the ventilatory support control signal  614  by a given amount. 
     Returning to block  604 , if the present breath RMS value on line  508  is not greater than the target drive signal  602 , it may be equal or smaller. In block  605 , it is determined if the present breath RMS value  508  is equal to the target drive signal  602 . 
     If it is equal, block  605  commands block  700  to keep the level of the ventilatory support control signal  614  constant. 
     Detection by block  605  that the present breath RMS value  508  is smaller than the target drive signal  602  indicates a decrease in diaphragm electrical activity (EMGdi) resulting in reduced respiratory drive. Block  703  then commands block  700  to produce a decrease of the ventilatory support, for example pressure support, through a decrease of the level of the ventilatory support control signal  614 . In other words block  703  commands block  700  to decrease the level of the ventilatory support control signal  614  by a given amount. 
     Those of ordinary skill in the art will understand that the amount of the change (increase or decrease) in ventilatory assistance is derived from experience, the patient&#39;s condition, the environment, etc. The amount of change can therefore be adjusted on a case by case basis. 
     Those of ordinary skill in the art will also understand that the test at blocks  604  and  605  can include a certain range of amplitudes for the target drive signal  602  (whether they are absolute or relative amplitudes); for example, the target drive signal could be X plus or minus a predetermined value. Therefore, at block  604 , the present breath RMS value  508  needs to be greater than X plus the predetermined value in order to proceed to block  701 . In the same way, at step  605 , the present breath RMS value  508  needs to be smaller than X minus the predetermined value in order to proceed to step  703 . 
       FIG. 10  illustrates a lung ventilator  54  capable of being controlled by the multiplied, RMS value  614  of the double-subtracted signal produced in step  612  of  FIG. 5B . Although an air-flow-based pressure ventilator is illustrated as an example in  FIG. 10 , it should be kept in mind that the RMS value of the double subtracted signal can be used for controlling any other lung ventilator. 
     Ventilator  54  shown in  FIG. 10  as an illustrative example only comprises a flow control unit  53 , a flow pump  55 , a patient&#39;s respiratory (inspiratory and expiratory) implement  56  such as a mask, a tracheal tube connector, or any other respiratory implement, a pressure sensor  57 , a pressurizing valve  58 , and a depressurizing valve  59 . 
     The flow pump  55  produces a constant air flow and supply of this air flow to the patient&#39;s respiratory accessory  56  is controlled through the pressurizing valve  58 . The patient is allowed to breathe out through the respiratory accessory  56  and the depressurizing valve  59 . The pressurizing and depressurizing valves  58  and  59  are controlled by the flow control unit  53 . 
     The pressure sensor  57  is connected close to the respiratory implement  56  through a line  60 . The pressure sensor  57  produces a corresponding respiratory pressure representative signal  61  supplied to the flow control unit  53 . Accordingly, the pressure sensor  57  provides feedback of actual respiratory pressure close to the respiratory implement  56 . The flow control unit  53  is also supplied with the multiplied, RMS value  614  of the double-subtracted signal delivered on line  62  ( FIG. 10 ) by step  612  of  FIG. 5B . 
     Those of ordinary skill in the art know that the amplitude of the multiplied, RMS value  614  of the double-subtracted signal delivered on line  62  is a representation of the demand to breathe from the brain. 
     When the RMS value  614  supplied to the flow control unit  53  is higher than the amplitude of the pressure representative signal  61 , this indicates that the demand to breath from the brain is higher than the air actually breathed by the patient. Inspiratory assist is then required and the flow control unit  53  will open pressurizing valve  58  to supply air flow from the pump  55  to the patient&#39;s respiratory accessory (depressurizing valve  59  being closed) until the amplitude of the pressure representative signal  61  is equal to the multiplied, RMS value  614 . The flow control unit  53  will continue to control the position of valve  58  to maintain the amplitude of the pressure representative signal  61  equal to the multiplied, RMS value  614  during all the inspiratory cycle. 
     During the inspiratory cycle, when the multiplied, RMS value  614  falls slightly below the amplitude of the pressure representative signal  61 , depressurizing valve  59  can be opened to correct the situation and maintain the amplitude of the pressure representative signal  61  equal to the multiplied, RMS value  614 . 
     When the multiplied, RMS value  614  drops below a given threshold, this indicates the beginning of an expiratory cycle. Then, the flow control unit  53  closes pressurizing valve  58  and opens depressurizing valve  59  to allow the patient to breath out through the respiratory accessory  56  and the depressurizing valve  59 . 
     In another example embodiment of the invention, in order to obtain correct proportionality between the pressure representative signal  61  and the multiplied, RMS value  614 , a gain adjustment is introduced for example in sensor  57  or on the line  62  to adequately control pressure assist to the respiratory implement  56  in function of the multiplied, RMS value  614 . 
     Accordingly, the subject invention presents a major advantage over the prior art. Indeed, the degree of inspiratory assist is adjusted in relation to the real need of the patient. In other words, assist is proportional to the difference between the pressure representative signal  61  and the multiplied, RMS value  614 . Inspiratory assist is therefore provided only to compensate for the degree of incapacity of the patient. The patient still contributes to inspiration as a function of his capacity to prevent the lung ventilator to further reduce the patient&#39;s inability to breathe. Requiring breathing efforts from the patient usually accelerates recovery of the patient and faster disconnection of the patient from the lung ventilator. 
     Although the present invention has been described herein above with reference to preferred embodiments thereof, these embodiments can be modified at will, within the scope of the appended claims, without departing from the spirit and nature of the subject invention.

Technology Category: 1