in the embodiment shown in fig1 , a ventilation system 1 has a driven fan 3 . the fan 3 is driven by a motor or some other power source . the fan 3 is connected to a control valve 4 , which has a control unit 6 and transforms a control voltage to an associated valve position . in the illustrated embodiment , a pressure sensor 8 and a volume flow sensor 9 are connected to a ventilator 7 , which consists essentially of the fan 3 and the control valve 4 . the volume flow sensor 9 is typically designed as a differential pressure sensor , whose signal is converted to an associated volume flow . the ventilator 7 is connected by a respiratory gas hose 10 and an expiration valve 11 to a ventilation mask 12 , which can be positioned over the face of a patient 13 . the sensors 8 , 9 are connected to an evaluation unit 14 , which in turn is connected to the control unit 6 . fig2 shows an equivalent electrical circuit diagram , which reproduces the function of the lung of the patient 13 and is used in the evaluation unit 14 as a model for performing the computations . a volume flow generated by a volume flow source 15 is supplied to the parallel connection of two flow branches . one of the branches contains the series connection of the resistance 16 and the compliance 17 and a source of interference 18 , which generates an additional volume flow and represents possible activity of the patient 13 himself . the second flow branch contains a leakage resistance 19 . the equivalent electrical circuit diagram of the respiratory tract according to fig2 shows that the pressure p mus of a patient produced by his active breathing activity contributes to the airway pressure p aw and thus to the flow into the lung v ′ aw . accordingly , leak identification must also take p mus into account if the leakage flow v ′ l is to be adequately compensated . assuming that the respiratory effort of the patient between inspirations varies by only an insignificant amount , the respiratory effort component can be eliminated by subtraction . the following relationship in the complex variable domain of the laplace transform is obtained from the equivalent circuit diagram according to fig2 for the airway pressure p aw and the inspiratory flow v ′ insp : p aw ( s )=[ p mus ( s )+ v ′ insp ( s )( 1 + src )/ sc ]× sr l c /( 1 s ( r + r l ) c ) the differences of the airway pressure δp aw and the inspiratory flow δv ′ insp are computed as follows : δ p aw ( s )= p aw , k ( s )− p aw , k1 ( s ) δ v ′ insp ( s )= v ′ insp , k ( s )− v ′ insp , k 1 ( s ) assuming that the respiratory effort with respect to time is the same in successive inspirations , i . e ., assuming that a simpler relationship is obtained between the pressure difference δp aw and the flow difference δv ′ insp in the complex variable domain of the laplace transform : δ p aw ( s )= δ v ′ insp ( s )( 1 + src ) r l /( 1 + s ( r + r l ) c ) to reconstruct the behavior of p mus , the inversion of the first equation must be used . to this end , the identified estimated values are used as parameters , and the measured values are accounted for . p ^ mus ( s )=( 1 + s ( r ^+ r ^ l ) c ^)/ sr ^ l c ^ p fs ( s )−( 1 + sr ^ c ^ )/ sc ^ v ′ fs ( s ) a reconstruction of this type allows an evaluation of the spontaneous respiration and thus indirectly an evaluation of the quality of ventilation . in the simplest case , it can be decided by integration of the reconstructed values whether the patient is breathing with or against . for the first time , more complicated methods of classification can derive from this a decided evaluation for the volume - controlled ventilation , for example , with respect to synchronization or the level of respiratory effort . fig3 shows the behavior with respect to time of the difference of the measured quantities between two respiratory cycles , one being an inspiratory volume flow difference 20 , and the other being a pressure difference 21 in the area of the ventilation mask 12 . fig4 shows , with respect to a first pressure curve 22 for a first respiratory cycle and a second pressure curve 23 for a second respiratory cycle , the determined reconstruction curve 24 for spontaneous respiration of the patient 13 , taking into account the volume flow difference 20 and the pressure difference 21 between two inspirations with different inspiratory volume flows . fig5 shows the device illustrated in fig1 in the form of a functional block diagram . the ventilator 7 is connected to a respiratory gas hose 10 , and the volume flow sensor 9 and the pressure sensor 8 are arranged in the area between the ventilator 7 and the respiratory gas hose 10 . the respiratory gas hose 10 is connected to the ventilation mask 12 , and an additional volume flow sensor 25 and an additional pressure sensor 26 are positioned in the area between the respiratory gas hose 10 and the ventilation mask 12 . as an alternative to the illustration in fig5 , the sensors can also be arranged only in the area between the ventilator 7 and the respiratory gas hose 10 or only in the area between the respiratory gas hose 10 and the ventilation mask 12 . naturally , a setup with a mass flowmeter is also possible ; and in this case , a conversion to volume flow is made . in accordance with the embodiment shown in fig6 , an expiration valve 11 is arranged in the area between the respiratory gas hose 10 and the ventilation mask 12 , and a discharge system 27 is arranged between the ventilator 7 and the respiratory gas hose 10 . it is also possible to arrange only the expiration valve 11 in the area between the respiratory gas hose 10 and the ventilation mask 12 or to arrange only the discharge system 27 in the area between the ventilator 7 and the respiratory gas hose 10 . furthermore , it is possible to replace the expiration valve 11 with a discharge system 27 and the discharge system 27 with an expiration valve 11 . as an alternative to the ventilation mask 12 shown in the specific embodiments , other noninvasive devices can be used to provide a connection with the patient 13 . for example , the use of masks or headpieces is possible . a connection between the ventilation system 1 and the patient 13 can also be established by invasive coupling devices , for example , an intubation tube , a tracheostoma , or a laryngeal mask . the ventilation systems that are used can be designed for carrying out various types of ventilation , for example , control mode ventilation , assist / control mode ventilation , or assist mode ventilation . in addition , the method can be used in periodic breathing and for cpap ventilation or apap ventilation . the determination of resistance and compliance can be made , for example , exclusively during the inspiratory or exclusively during the expiratory time intervals of the respiratory cycles . however , it is also possible to make these determinations during both the inspiratory and expiratory phases . fig7 shows the previously mentioned embodiment , in which the respiratory gas hose 10 is directly coupled with the ventilation mask 12 , and the discharge system 27 and the sensors 8 , 9 are positioned only in the area between the ventilator 7 and the respiratory gas hose 10 . fig8 shows an embodiment that is a modification of the embodiment shown in fig7 , in which the discharge system 27 and the volume flow sensor 9 are positioned between the ventilator 7 and the respiratory gas hose 10 , and the pressure sensor 8 is positioned between the respiratory gas hose 10 and the ventilation mask 12 . in the embodiment in fig9 , the discharge system 27 and the pressure sensor 8 are positioned between the ventilator 7 and the respiratory gas hose 10 , and the volume flow sensor 9 is positioned between the respiratory gas hose 10 and the ventilation mask 12 . in the modification shown in fig1 , only the discharge system 27 is arranged between the ventilator 7 and the respiratory gas hose 10 , and both the volume flow sensor 9 and the pressure sensor 8 are positioned between the respiratory gas hose 10 and the ventilation mask 12 . in the modification shown in fig1 , both the pressure sensor 8 and the volume flow sensor 9 are positioned between the ventilator 7 and the respiratory gas hose 10 , and the expiration valve 11 is positioned between the respiratory gas hose 10 and the ventilation mask 12 . in the embodiment shown in fig1 , only the volume flow sensor 9 is arranged between the ventilator 7 and the respiratory gas hose 10 , and the expiration valve 11 and the pressure sensor 8 are positioned between the respiratory gas hose 10 and the ventilation mask 12 . in the embodiment shown in fig1 , only the pressure sensor 8 is located between the ventilator 7 and the respiratory gas hose 10 , and both the expiration valve 11 and the volume flow sensor 9 are located between the respiratory gas hose 10 and the ventilation mask 12 . finally , fig1 shows an embodiment in which the ventilator 7 is directly coupled with the respiratory gas hose 10 , and the expiration valve 11 and both the pressure sensor 8 and the volume flow sensor 9 are positioned between the respiratory gas hose 10 and the ventilation mask 12 . fig1 shows an equivalent circuit diagram that is supplemented relative to the diagram in fig2 and is used in the case of linear behavior of the system comprising the patient and the respiratory gas hose . in an additional branch , the capacity 28 resulting from the volume of the hose is taken into consideration . for the relationship between the measured data for flow v ′ fs and pressure p fs in the device , the transfer function corresponding to the following equation is obtained from the equivalent circuit diagram in the complex variable domain of the laplace transform , neglecting a peep ( positive end - expiratory pressure ): p fs ( s )/ v ′ fs ( s )=( r l ×( 1 + src ))/( 1 + s ( rc + r l c + r l c s )+ s 2 rcr l c s ) in this regard , the transit time behavior of the pressure inside the hose was disregarded ( p fs = p aw ), and r l was adopted as the leak resistance . after a time discretization of this equation with a zero - order hold ( s −& gt ;( 1 − z − 1 )/ t ), with the discrete - time shift operator z − 1 , and with the sampling time t , the following representation of the relationship between the sequences of values p fs , k and v ′ fs , k is obtained : p fs , k / v ′ fs , k =( b 0 + b 1 z − 1 )/( 1 + a 1 z 1 + a 2 z 2 ) if we now consider the conditions on the patient side of the respiratory gas hose , the following transfer function between v ′ insp and p aw is obtained : p aw ( s )/ v ′ insp ( s )= r l ×( 1 + src )/( 1 + s ( r + r l )+ c ) p aw , k / v ′ insp , k =( b 0 + b 1 z 1 )/( 1 + a 1 z 1 ) if the transit time of the air pressure through the respiratory gas hose is taken into account , then p fs is not equal to p aw . for pressure sensors on the device side , the expansion with p fs ( s )= p aw ( s )/( 1 + st s × e stt ) can be used , in which t s is the integral - action time of the flexible respiratory gas hose with the hose connector . the delay time tt corresponds to the sound transit time through the hose . a higher degree of accuracy of the models in the above equations can be used to take into account such effects as gas compressibility . more precise formulations can also be used in the transition to the discrete - time representation . the parameters r , c , and r l can be estimated by well - known identification and parameter estimation methods for linear as well as nonlinear systems . especially a time - distributed computation of the parameter estimation is advantageous for implementation . in this regard , recursive methods make it possible to obtain results with the greatest time proximity to the last observation period .
the invention relates to a method and a device for detecting leaks in respiratory gas supply systems . both the pressure and the volume flow of the respiratory gas are detected and the relevant values are supplied to an evaluation device . the evaluation device is used to calculate both the respiratory parameter resistance and compliance and the leak for at least two successive breathing cycles . at least one control parameter with different signal amplitudes is pre - determined for the successive breathing cycles . the leak resistance is determined from the resulting differential sequences of pressure and flow for the successive breathing cycles .