Patent Publication Number: US-9421338-B2

Title: Ventilator leak compensation

Description:
RELATED PATENT APPLICATION 
     This application is a continuation application of U.S. patent application Ser. No. 12/414,419 (now U.S. Pat. No. 8,434,480), entitled “VENTILATOR LEAK COMPENSATION,” filed on Mar. 30, 2009, which application claims priority from U.S. Provisional Application Ser. No. 61/041,070, which was filed on Mar. 31, 2008, the complete disclosures of which are hereby incorporated by reference in their entirety. 
    
    
     BACKGROUND 
     The present description pertains to ventilator devices used to provide breathing assistance. Modern ventilator technologies commonly employ positive pressure to assist patient ventilation. For example, after determining a patient-initiated or timed trigger, the ventilator delivers a specified gas mixture into an inhalation airway connected to the patient to track a specified desired pressure or flow trajectory, causing or assisting the patient&#39;s lungs to fill. Upon reaching the end of the inspiration, the added support is removed and the patient is allowed to passively exhale and the ventilator controls the gas flow through the system to maintain a designated airway pressure level (PEEP) during the exhalation phase. Other types of ventilators are non-triggered, and mandate a specified breathing pattern regardless of patient effort. 
     Modern ventilators typically include microprocessors or other controllers that employ various control schemes. These control schemes are used to command a pneumatic system (e.g., valves) that regulates the flow rates of breathing gases to and from the patient. Closed-loop control is often employed, using data from pressure/flow sensors. 
     Many therapeutic settings involve the potential for leaks occurring at various locations on the ventilator device. The magnitude of these leaks can vary from setting to setting, and/or dynamically within a particular setting, dependent upon a host of variables. Leaks can impair triggering (transition into inhalation phase) and cycling (transition into exhalation phase) of the ventilator; and thus cause problems with patient-device synchrony; undesirably increase patient breathing work; degrade advisory information available to treatment providers; and/or otherwise compromise the desired respiratory therapy. 
     Accordingly, attempts have been made in existing control systems to compensate for leaks in ventilator components. Though some benefits have been achieved, prior compensation mechanisms typically are predicated on simplified assumptions or limited information, which limits the ability to accurately and dynamically account for system leaks in general and instantaneous leak rates in particular. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a schematic depiction of a ventilator. 
         FIG. 2  schematically depicts control systems and methods that may be employed with the ventilator of  FIG. 1 . 
         FIGS. 3A and 3B  depict exemplary tidal breathing in a patient, and examples of pressure/flow waveforms observed in a ventilator under pressure support with and without leak condition. Under leak condition, the inhalation flow is the total delivered flow including the leak flow and the exhalation flow is the output flow rate measured by the ventilator and excludes the exhaled flow exhausted through the leak. 
         FIGS. 4A and 4B  depict an example embodiment of the patient interface shown in  FIG. 1 . 
         FIG. 5  depicts an exemplary method for controlling the ventilator of  FIG. 1 , including a method for compensating for leaks in ventilator components. 
     
    
    
     DETAILED DESCRIPTION 
       FIG. 1  depicts a ventilator  20  according to the present description. As will be described in detail, the various ventilator system and method embodiments described herein may be provided with control schemes that provide improved leak estimation and/or compensation. These control schemes typically model leaks based upon factors that are not accounted for in prior ventilators, such as elastic properties and/or size variations of leak-susceptible components. The present discussion will focus on specific example embodiments, though it should be appreciated that the present systems and methods are applicable to a wide variety of ventilator devices. 
     Referring now specifically to  FIG. 1 , ventilator  20  includes a pneumatic system  22  for circulating breathing gases to and from patient  24  via airway  26 , which couples the patient to the pneumatic system via physical patient interface  28  and breathing circuit  30 . Breathing circuit  30  could be a two-limb or one-limb circuit for carrying gas to and from the patient. A wye fitting  36  may be provided as shown to couple the patient interface to the breathing circuit. 
     The present systems and methods have proved particularly advantageous in non-invasive settings, such as with facial breathing masks, as those settings typically are more susceptible to leaks. However, leaks do occur in a variety of settings, and the present description contemplates that the patient interface may be invasive or non-invasive, and of any configuration suitable for communicating a flow of breathing gas from the patient circuit to an airway of the patient. Examples of suitable patient interface devices include a nasal mask, nasal/oral mask (which is shown in  FIG. 1 ), nasal prong, full-face mask, tracheal tube, endotracheal tube, nasal pillow, etc. 
     Pneumatic system  22  may be configured in a variety of ways. In the present example, system  22  includes an expiratory module  40  coupled with an expiratory limb  34  and an inspiratory module  42  coupled with an inspiratory limb  32 . Compressor  44  is coupled with inspiratory module  42  to provide a gas source for ventilatory support via inspiratory limb  32 . 
     The pneumatic system may include a variety of other components, including sources for pressurized air and/or oxygen, mixing modules, valves, sensors, tubing, accumulators, filters, etc. Controller  50  is operatively coupled with pneumatic system  22 , signal measurement and acquisition systems, and an operator interface  52  may be provided to enable an operator to interact with the ventilator (e.g., change ventilator settings, select operational modes, view monitored parameters, etc.). Controller  50  may include memory  54 , one or more processors  56 , storage  58 , and/or other components of the type commonly found in command and control computing devices. As described in more detail below, controller  50  issues commands to pneumatic system  22  in order to control the breathing assistance provided to the patient by the ventilator. The specific commands may be based on inputs received from patient  24 , pneumatic system  22  and sensors, operator interface  52  and/or other components of the ventilator. In the depicted example, operator interface includes a display  59  that is touch-sensitive, enabling the display to serve both as an input and output device. 
       FIG. 2  schematically depicts exemplary systems and methods of ventilator control. As shown, controller  50  issues control commands  60  to drive pneumatic system  22  and thereby circulate breathing gas to and from patient  24 . The depicted schematic interaction between pneumatic system  22  and patient  24  may be viewed in terms of pressure and/or flow “signals.” For example, signal  62  may be an increased pressure which is applied to the patient via inspiratory limb  32 . Control commands  60  are based upon inputs received at controller  50  which may include, among other things, inputs from operator interface  52 , and feedback from pneumatic system  22  (e.g., from pressure/flow sensors) and/or sensed from patient  24 . 
     In many cases, it may be desirable to establish a baseline pressure and/or flow trajectory for a given respiratory therapy session. The volume of breathing gas delivered to the patient&#39;s lung and the volume of the gas exhaled by the patient are measured or determined, and the measured or predicted/estimated leaks are accounted for to ensure accurate delivery and data reporting and monitoring. Accordingly, the more accurate the leak estimation, the better the baseline calculation of delivered and exhaled volume as well as event detection (triggering and cycling phase transitions). 
       FIGS. 2, 3A and 3B  may be used to illustrate and understand leak effects and errors. As discussed above, therapy goals may include generating a desired time-controlled pressure within the lungs of patient  24 , and in patient-triggered and -cycled modes, achieve a high level of patient-device synchrony. 
       FIG. 3A  shows several cycles of flow/pressure waveforms spontaneous breathing under Pressure Support mode with and without leak condition. As discussed above, a patient may have difficulty achieving normal tidal breathing, due to illness or other factors. 
     Regardless of the particular cause or nature of the underlying condition, ventilator  20  typically provides breathing assistance during inspiration and exhalation.  FIG. 3B  shows an example of flow waveform under Pressure Support in presence of no leak as well as leak conditions. During inspiration more flow is required (depending on the leak size and circuit pressure) to achieve the same pressure level compared to no leak condition. During exhalation, a portion of the volume exhaled by the patient would exit through the leak and be missed by the ventilator exhalation flow measurement subsystem. In many cases, the goal of the control system is to deliver a controlled pressure or flow profile or trajectory (e.g., pressure or flow as a function of time) during the inspiratory phase of the breathing cycle. In other words, control is performed to achieve a desired time-varying pressure or flow output  62  from pneumatic system  22 , with an eye toward causing or aiding the desired tidal breathing shown in  FIG. 3A . 
     Improper leak accounting can compromise the timing and magnitude of the control signals applied from controller  50  to pneumatic system  22  especially during volume delivery. Also, lack or inaccurate leak compensation can jeopardize spirometry and patient data monitoring and reporting calculations. As shown at schematic leak source L 1 , the pressure applied from the pneumatic system  22  to patient interface  28  may cause leakage of breathing gas to atmosphere. This leakage to atmosphere may occur, for example, at some point on inspiratory limb  32  or expiratory limb  34 , or at where breathing circuit  30  couples to patient interface  28  or pneumatic system  22 . 
     In the case of non-invasive ventilation, it is typical for some amount of breathing gas to escape via the opening defined between the patient interface (e.g., facial breathing mask) and the surface of the patient&#39;s face. In facial masks, this opening can occur at a variety of locations around the edge of the mask, and the size and deformability of the mask can create significant leak variations. As one example, as shown in  FIG. 4A  and  FIG. 4B , the facial breathing mask may be formed of a deformable plastic material with elastic characteristics. Under varying pressures, during inspiration and expiration the mask may deform, altering the size of the leak orifice  61 . Furthermore, the patient may shift (e.g., talk or otherwise move facial muscles), altering the size of leak orifice  61 . Due to the elastic nature of the mask and the movement of the patient a leak compensation strategy assuming a constant size leak orifice may be inadequate. 
     Accurately accounting for the magnitude of leak L 1  may provide significant advantages. In order for controller  50  to command pneumatic system  22  to deliver the desired amount of volume/pressure to the patient at the desired time and measure/estimate the accurate amount of gas volume exhaled by the patient, the controller must have knowledge of how large leak L 1  is during operation of the ventilator. The fact that the leak magnitude changes dynamically during operation of the ventilator introduces additional complexity to the problem of leak modeling. 
     Triggering and cycling (patient-ventilator) synchrony may also be compromised by sub-optimal leak estimation. In devices with patient-triggered and patient-cycled modalities that support spontaneous breathing efforts by the patient, it can be important to accurately detect when the patient wishes to inhale and exhale. Detection commonly occurs by using accurate pressure and/or lung flow (flow rates into or out of the patient lung) variations. Leak source L 2  represents a leak in the airway that causes an error in the signals to the sensors of pneumatic system  22 . This error may impede the ability of ventilator to detect the start of an inspiratory effort, which in turn compromises the ability of controller  50  to drive the pneumatic system in a fashion that is synchronous with the patient&#39;s spontaneous breathing cycles. 
     Improved leak estimation may be achieved in the present examples through provision of a control scheme that more fully accounts for factors affecting the time-varying magnitude of leaks under interface and airway pressure variations. The present example may include, in part, a constant-size leak model consisting of a single parameter (orifice resistance, leak conductance, or leak factor) utilized in conjunction with the pneumatic flow equation through a rigid orifice, namely,
 
 Q   leak =(leak factor/Resistance/Conductance)*√{square root over (Δ P )}  (1)
 
Where ΔP=pressure differential across the leak site. This assumes a fixed size leak (i.e., a constant leak resistance or conductance or factor over at least one breath period).
 
     To provide a more accurate estimate of instantaneous leak, the leak detection system and method may also take into account the elastic properties of one or more components of the ventilator device (e.g., the face mask, tubing used in the breathing circuit, etc.). This more accurate leak accounting enhances patient-ventilator synchrony and effectiveness under time-varying airway pressure conditions in the presence of both rigid orifice constant size leaks as well as pressure-dependent varying-size elastic leak sources. 
     According to the pneumatic equations governing the flow across an orifice, the flow rate is a function of the area and square root of the pressure difference across the orifice as well as gas properties. For derivation of the algorithm carried out by the controller, constant gas properties are assumed and a combination of leak sources comprising of rigid fixed-size orifices (total area=A r =constant) and elastic opening through the patient interface [total area=A e (P)=function of applied pressure]. Therefore,
 
 Q   leak   =K   o *( A   r   +A   e ( P ))*√{square root over (Δ P )}  (2)
         K o =assumed constant       

     For the purposes of this implementation, at low pressure differences, the maximum center deflection for elastic membranes and thin plates are a quasi-linear function of applied pressure as well as dependent on other factors such as radius, thickness, stress, Young&#39;s Modulus of Elasticity, Poisson&#39;s Ratio, etc. Therefore,
 
 A   e ( P )= K   e   *ΔP   (3)
         K e =assumed constant       

     As ΔP is the pressure difference across a leak source to ambient (P ambient =0), then we substitute ΔP by the instantaneous applied pressure P(t) and rearrange equation 1 as follows (K 1  and K 2  are assumed to be constant):
 
 Q   leak   =K   0 ( A   r   +K   e   P ( t )√{square root over ( P ( t ))}  (4)
 
 Q   leak   =K   1   *P ( t ) 1/2   +K   2   *P ( t ) 3/2   (5)
 
     Also, the total volume loss over one breath period=V leak =Delivered Volume−Exhausted Volume; 
     
       
         
           
             
               
                 
                   
                     
                       
                         
                           
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     The general equation of motion for a patient ventilator system during passive exhalation can then be written,
 
 P   aw   +P   m   =R *( Q   leak   +Q   exh   −Q   delivered )+(1 /C )*∫[ Q   leak   +Q   exh   −Q   delivered   ]*dt   (7)
         P aw =airway pressure   P m =muscle pressure   R=resistance   C=Compliance       

     Assuming that when end exhalation conditions are present a constant airway pressure is being delivered (steady PEEP), constant bias flow maintained during exhalation phase Q delivered , constant leak flow (due to no pressure variation), and P m =0 (due to no patient respiratory effort), the equation of motion could be differentiated and reorganized as follows: 
                       ⅆ     P   aw         ⅆ   t       =     0   =       R   *     Q   exh     ⁢   dot     +         Q   leak     +     Q   exh     -     Q   delivered       C                 (   8   )                   Q   leak     =       (       Q   delivered     -     Q   exh       )     -     R   *   C   *     Q   exh     ⁢   dot         ⁢     
     ⁢         Q   exh     ⁢   dot     =     time   ⁢           ⁢   derivative   ⁢           ⁢   of   ⁢           ⁢   exhausted   ⁢           ⁢   flow               (   9   )               
If Q exh dot=0 equation 8 can be reduced to
 
 Q   leak   =Q   delivered   −Q   exh   (10)
 
And subsequently,
 
 Q   leak   =K   1 (PEEP) 1/2   +K   2 (PEEP) 3/2   (11)
 
     Otherwise Q exh dot≠0. In this case, an appropriate duration of time ΔT is taken during passive exhalation period and assuming constant delivered flow, equation can be derived as follows: 
                     R   *   C     =       (         Q   exh     ⁡     (     t   +     Δ   ⁢           ⁢   T       )       -       Q   exh     ⁡     (   t   )             (         Q   exh     ⁢     dot   ⁡     (     t   +     Δ   ⁢           ⁢   T       )         -       Q   exh     ⁢     dot   ⁡     (   t   )                       (   12   )               
And,
 
 Q   leak ( t   i   +ΔT )= K   1 (PEEP) 1/2   +K   2 (PEEP) 3/2   =[Q   delivered ( t   i   +ΔT )− Q   exh ( t   i   +ΔT )]− R*C*Q   exh dot( t   i   +ΔT )  (13)
 
     Therefore, equation 6 and equation 10 and equation 13 may be used to solve for K 1  and K 2 . These calculations may be repeated every breath cycle and applied over appropriate time windows (i.e. during exhalation) and breathing conditions to optimize parameter estimation and minimize the total error between estimated total volume loss and actual measured volume loss across the full breath cycle. The constants K 1  and K 2  may be stored and compared to track changes and update various parameters of the system such as the triggering and cycling sensitivities, etc. 
       FIG. 5  shows an exemplary control strategy that may be implemented by the controller  50  to increase the accuracy and timing of the baseline breathing assistance provided by ventilator  20  and pneumatic system  22  for a variety of respiratory therapies. In this example, the method is repeated periodically every breathing cycle. In other examples, the dynamic updating of leak estimation may occur more or less than once per patient breathing cycle. 
     At  512  the routine establishes a baseline level of leak estimation and compensation. This may be a preset value stored in the controller  50  or may be updated taking into account various parameters of the breathing cycle and ventilator  20 , such as the Positive End Expiratory Pressure PEEP, the set inspiratory pressure or flow/volume targets, the volumetric airflow delivered by pneumatic system  22 , and type of the breathing circuit  30 , etc. 
     The routine then proceeds to  514  where the feedback control (e.g., as shown in  FIG. 3 ) is implemented. Various control regimes may be implemented, including pressure, volume and/or flow regulation. Control may also be predicated on inputs received from the patient, such as pressure variations in the breathing circuit which indicate commencement of inspiration. Inputs applied via operator interface  52  may also be used to vary the particular control regime used. For example, the ventilator may be configured to run in various different operator-selectable modes, each employing different control methodologies. 
     The routine advances to  516  where the leak compensation is performed. Various types of leak compensation may be implemented. For example, as shown at  518 , rigid-orifice compensation may be employed using values calculated as discussed above. In particular, holes or other leak sources may be present in non-elastic parts of the breathing circuit, such as the ports of a facial mask (not shown) and/or in the inspiratory and expiratory limbs. Equation 1 may be used to calculate the volumetric airflow through such an orifice, assuming the leak factor/resistance/conductance is constant. 
     Elastic properties of ventilator components may also be accounted for during leak compensation, as shown at  520 , for example using values calculated as described above. Specifically, elastic properties of patient interface  28  and/or breathing circuit  30  may be established (e.g., derived based on material properties such as elastic modulus, Poisson&#39;s ratio, etc.), and employed in connection with calculations such as those discussed above in reference to equations 6, 10, and/or 13, to account for the deformation of orifice  61 , as shown in  FIG. 4B . Using these example calculations, constants K 1  and K 2  may be solved for and updated dynamically to improve the accuracy of leak estimation. In alternate implementations, the method may use any suitable alternate mechanism or models for taking into account the elastic properties of a ventilator component having a leak-susceptible orifice. 
     The routine then proceeds to  522  where appropriate baseline control commands and measurements are adjusted to compensate for the leaks in the ventilator calculated in  616  i.e. adjust appropriate control command and correct and/or compensate applicable measurements. In many settings, it will be desirable to regularly and dynamically update the compensation level (e.g., once every breathing cycle) in order to optimize the control and compensation. 
     It will be appreciated that the embodiments and method implementations disclosed herein are exemplary in nature, and that these specific examples are not to be considered in a limiting sense, because numerous variations are possible. The subject matter of the present disclosure includes all novel and nonobvious combinations and subcombinations of the various configurations and method implementations, and other features, functions, and/or properties disclosed herein. Claims may be presented that particularly point out certain combinations and subcombinations regarded as novel and nonobvious. Such claims may refer to “an” element or “a first” element or the equivalent thereof. Such claims should be understood to include incorporation of one or more such elements, neither requiring nor excluding two or more such elements. Other combinations and subcombinations of the disclosed features, functions, elements, and/or properties may be claimed through amendment of the present claims or through presentation of new claims in this or a related application. Such claims, whether broader, narrower, equal, or different in scope to the original claims, also are regarded as included within the subject matter of the present disclosure.