Abstract:
The present invention describes a method and apparatus for detecting and quantifying intrinsic positive end-expiratory pressure (PEEPi) of a respiratory patient breathing with the assistance of a ventilator. A processing device receives respiratory airway data from one or more sensors adapted to non-invasively monitor a respiratory patient, calculates from the respiratory airway data two or more parameters that are indicative of or quantify intrinsic positive end-expiratory pressure of the patient, and generates an indication intrinsic positive end-expiratory pressure (PEEPi).

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     This application is a continuation-in-part of U.S. application Ser. No. 12/497,733 filed Jul. 6, 2009, now U.S. Pat. No. 8,544,466, which is a continuation of U.S. application Ser. No. 11/165,617 filed Jun. 23, 2005, now U.S. Pat. No. 7,562,657, which claims priority to the Jun. 24, 2004 filing date of U.S. provisional patent application No. 60/582,409, all of which are incorporated herein by reference. 
    
    
     FIELD OF INVENTION 
     The present invention relates generally to the field of respiratory therapy, physiology, and critical care medicine including ventilator management and respiratory monitor technology, and, more particularly, to a method and apparatus for non-invasive prediction of intrinsic positive end-expiratory pressure (PEEP i ) using certain markers. 
     BACKGROUND OF INVENTION 
     Mechanical ventilatory support is widely accepted as an effective means for supporting and treating patients with respiratory failure. Mechanical ventilators are simply machines designed to assist with inspiration and expiration. Often, a primary objective of ventilatory support is for the ventilator to provide some or all of a patient&#39;s work of breathing (WOB). This goal is often not achieved due in part to an inability to accurately measure and titrate a patient&#39;s WOB. Ventilators must be highly reliable, durable and precise. Most modern ventilators are electronically controlled and most are designed to allow many small, but different, fine-tuning manipulations by the operator. Ideally, the operator uses these controls to match the pressure and flow output characteristics of the ventilator to meet each individual patient&#39;s needs. Optimized ventilator settings also serve to make ventilatory support more tolerable for the patient. 
     The first generation of mechanical ventilators (prior to the mid-1960s), were designed only to support alveolar ventilation and to provide supplemental oxygen for those patients unable to breathe themselves (generally for reasons such as neuromuscular disease or paralysis). These early ventilators provided 100% of the work required to breathe if the supported patient did not breathe on his or her own. If the patient attempted to breathe spontaneously, a complete lack of a response from the ventilator created agitation—also termed “fighting the ventilator”. Those patients that tried to breathe spontaneously became so agitated, it was (and remains in some institutions) common practice to heavily sedate or paralyze the patient to ensure proper synchrony with the ventilator. Sedation or drug-induced muscle paralysis, may create more problems than they solve. For instance, heavily sedated or paralyzed patients simply can not breathe spontaneously; as a result, if they become accidentally disconnected from the ventilator they will quickly die of asphyxia. This required an increased vigilance and a very elevated level of monitoring. More importantly, paralysis and sedation also lead to a rapid deterioration of the patient&#39;s respiratory muscles, or what has become known as: “disuse atrophy”. Without strong, conditioned respiratory muscles, clinicians find it difficult or, on occasion, nearly impossible to liberate their patients from the ventilator, even when their original pulmonary problems have resolved. To overcome these potential problems, modern mechanical ventilators have become far more sophisticated in response to our increasing understanding of lung patho-physiology. In an effort to improve patient tolerance of mechanical ventilation, while simultaneously maintaining adequate respiratory muscle function, many new modes have been developed. Many of these new modes allowed spontaneous breathing, patient-triggered and pressure-assisted breaths, or even patient-triggered mandated breaths. These pressure assisted breaths, when properly adjusted by the clinician, allow the ventilator to share the WOB with the patient. By the early 1970s, a new mode of ventilation allowed patients to breathe at their own pace and magnitude, in the time period between mandated (machine delivered and controlled) breaths that were programmed for delivery at precise intervals each minute; this mode was known as Intermittent Mandatory Ventilation (IMV). Pressure Support Ventilation (PSV), came along a few years later and could be used alone or in combination with IMV; this mode provided a variable pressure-assist (at an operator chosen level of pressure) for each patient-initiated spontaneous inhalation effort. Varieties of “alternative” ventilation modes, addressing the needs of severely impaired patients, continue to be developed. 
     Patients receiving ventilator support need different levels of assistance; some require complete control of their ventilation, while others require varying levels of support depending upon their ability to sustain breathing on their own. Matching the ventilator support provided, to that required by the patient, remains to this day, an imposing challenge; too much support predisposes to disuse atrophy, while too little support often leads to a cycle of fatigue, followed by respiratory failure. At the present time, there is no readily available, easy to use, or reliable method or apparatus for estimating an appropriate support level. Confounding the matter further, a patient&#39;s requisite level of support may vary widely throughout the day, for a variety of reasons. 
     In those instances in which a patient requires mechanical ventilation due to respiratory failure, sleep apnea, post-operative care, or other situations, a wide variety of mechanical ventilators are available. Most modern ventilators allow the clinician to select and use several modes of inhalation either individually or in combination via the ventilator setting controls that are common to the ventilators. Some ventilators, such as those designed for noninvasive ventilation (NIV), primarily utilize mask interfaces and can be very simple to operate with very few choices of settings, typically some baseline pressure level with or without a higher level of pressure support during inhalation. 
     Gas inadvertently trapped in a patient&#39;s lungs, or PEEP i , interferes with the clinicians best attempts to estimate an appropriate support level for patients. For years, many clinicians were completely unaware of its existence. Hence when it was first reported, it was termed “intrinsic” or even “occult” because it is hidden from view (using conventional monitoring techniques). The extra (un-exhaled or trapped) gas, remaining in a patient&#39;s lungs at the onset of the next inhalation, creates an inspiratory threshold load (a positive pressure level above ambient) that the patient&#39;s inspiratory muscles must overcome before fresh gas can enter the lungs. Furthermore, since the inspiratory muscles are displaced from their normal resting position (by the hyperinflation) they are mechanically disadvantaged; that is, the direction of the respiratory muscles are pulling, no longer generates the largest possible change in volume/unit of force. All of this simply means that patients with dynamic pulmonary hyperinflation (DPH) or PEEP i , must work significantly harder to breathe. In addition, patients receiving ventilator assistance or, those patients that are generating, at least, a portion of the work of breathing, must generate another additional effort to breathe; that is, they must overcome the trigger pressure set on the ventilator before they receive any assistance from the ventilator. A trigger pressure must be used to synchronize the patient&#39;s efforts to the ventilator&#39;s response (otherwise, the ventilator would randomly initiate breaths, some of which might conflict with, or even negate, the patient&#39;s own efforts). The combination of an undetectable and difficult to quantify PEEP i  level and the ventilator&#39;s trigger setting, frequently produce an intolerable additional workload; the additional work is often high enough to produce inspiratory muscle fatigue, particularly in patients with poor respiratory muscle function. For those patients not trying to breathe on their own, undetected PEEP i  is just as problematic. The additional pressure it produces in the patient&#39;s chest can reduce venous blood return into the chest, which in turn reduces cardiac output and, ultimately, can dramatically reduce a patient&#39;s blood pressure. Too much trapped gas in the lungs also predisposes to over-inflation and structural damage, even rupture of the lungs. Research has shown that PEEP i  occurs far more frequently than is commonly believed. It has also been shown that in ventilator-dependent COPD patients, PEEP i  accounts for a large percentage of the patient&#39;s total ventilatory workload. Reducing or eliminating PEEP i  could then, have a major clinical impact for patients with an acute exacerbation of COPD. Clearly, detecting and accurately measuring PEEP i  represents an extremely important tool in managing affected patients. 
     In today&#39;s intensive care unit (ICU), most modern ventilators can measure a patient&#39;s PEEP i  but only when they do not breathe on their own—any movement or spontaneous efforts during measurement will invalidate that measurement. Ventilator-supported patients that breathe spontaneously virtually never have their PEEP i  accurately measured; to do so, requires the placement of an esophageal balloon or use of thoracic impedance measuring equipment. Both approaches are expensive, very technique oriented and extremely time-consuming. As a rule, these approaches are only used by researchers. The esophageal balloon is the most commonly employed approach because it is the least expensive and balloons tend to interfere with fewer other ongoing and required monitors or therapies. The concept involves using a properly inflated and positioned, balloon tipped catheter that is inserted into the patient&#39;s esophagus. When positioned properly, it is used to measure esophageal pressure (P es ). It has been shown that in certain esophageal locations (but not all locations), pressure changes within the esophagus are of the same magnitude as those occurring in the pleural space (although the absolute pressure values will likely NOT be the same). The change in esophageal pressure (from resting, or baseline) needed to abruptly bring expiratory flow to the point where it just crosses the zero flow axis (the instant just prior to the onset of flow into the lungs) represents PEEP i  also called dynamic PEEP i  (PEEP i,dyn ). 
     The esophageal balloon technique has never been popular with clinicians. The balloons are difficult to place, can interfere with important equipment like feeding tubes, and must be positioned and inflated properly to prevent inaccurate and misleading measurements. Further complicating the esophageal balloon procedure is signal quality. Frequent swallowing or inadvertent esophageal spasms can be difficult to discern and yet render the signal temporarily useless. Additionally, the esophagus is located just anterior of the heart and the signal, as a result, is often difficult to interpret without using a “heavy filter” to remove the unwanted pressure fluctuations secondary to the beating heart. Thoracic impedance equipment is not only expensive, difficult to use, and may not be consistently repeatable (according to researchers), it often interferes with absolutely indispensable electrocardiography monitoring leads, crucial intravenous catheters, and other important equipment. For these reasons, it seems likely that the esophageal balloon and thoracic impedance devices will remain investigative tools. 
     For those patients not breathing spontaneously, PEEP i  is measured using an appropriately timed, end-expiratory, airway occlusion maneuver. This involves a sudden occlusion of the expiratory valve (a blocking of the path normally taken by the patient&#39;s exhaled gases). Occlusion of the expiratory valve traps any additional gas that might be still leaving the patient&#39;s lungs, in the breathing circuit (which is attached to the patient&#39;s ventilator) and the patient&#39;s lungs. When the pressure (measured in the breathing circuit) stabilizes, if it is above the normal baseline pressure (the pressure at end-exhalation), the patient is said to have PEEP i . The occlusion technique could, potentially, be used for spontaneously breathing patients; but, to do so would require the patient to hold their breath, making absolutely no attempt to breathe, or even move, until the pressure in the breathing circuit reaches equilibrium. Unfortunately, this level of cooperation is almost nonexistent when patients are extremely sick, comatose, heavily sedated or, struggling to breathe against the additional workload imposed by PEEP i . 
     U.S. Pat. No. 6,588,422 pertains to the field of ventilatory support for respiratory failure, particularly due to lung disease, and in particular to automatically providing sufficient end expiratory pressure to unload PEEP i . The &#39;422 patent seeks to provide continuous and automatic adjustment of the expiratory pressure during ventilatory support, so as to substantially prevent dynamic airway compression and unload PEEP i  with the smallest amount of external expiratory pressure. The object of this invention involves varying the external pressure exerted by the ventilator during the exhalation phase and does not measure or quantify PEEP i  in any manner. Nor is it obvious how one could measure PEEPi using this patent. 
     Additionally, U.S. Pat. No. 6,240,920 discloses a method for determining at least one parameter related to a patient&#39;s spontaneous attempts at inspiration and/or the patient&#39;s respiratory effort in spontaneous attempts at inspiration. What is disclosed in the &#39;920 patent is again a potential method to counteract and manage PEEP i ; it provides no methodology for measuring or quantifying the patients actual PEEP i  level. 
     Developing a non-invasive measure of PEEP i  (that works in spontaneously breathing patients receiving ventilatory support) is complicated by a number of factors. First, the gas trapped at the alveolar level is almost always non-homogeneously distributed across both lungs thereby, making it difficult to measure the true PEEP i . This means that any measurement of PEEP i , is, at best, an average of the two lungs, considered as though they were identical. Alternatively, the excess pressure (trapped air) maybe hidden behind prematurely collapsed airways—making it virtually undetectable. And lastly, patient efforts to breathe create artifacts that vary on a breath-to-breath basis. Such artifacts may be difficult to separate from the ideal passive pressure waveforms. 
     In some situations, developing a non-invasive indication of PEEPi, possibly including different levels of severity, is also advantageous. Many clinicians do not need to know the exact value of PEEPi, but instead just need to know whether the patient falls into one of several categories of intrinsic PEEP. For example, without restriction, a patient may have either: a) no PEEPi, b) mild PEEPi, c) moderate PEEPi, or d) severe PEEPi. This information may then be used to adjust the ventilator settings. 
     SUMMARY OF THE INVENTION 
     There is a need in the art for accurately detecting and quantifying PEEP i  non-invasively, especially in spontaneously breathing patients. Non-invasive determination of a patient&#39;s intrinsic PEEP will 1) enable clinicians to be better informed of patient status, enabling better management of patient respiratory therapy, 2) improve patient analysis and physiologic monitoring/modeling, and 3) minimize patient discomfort (non-invasive) and clinician intervention (no need for occlusion) from invasive PEEP i  measurement. The measure will provide clinicians with accurate, continuous information that will enable them to make better decisions on patient&#39;s respiratory therapy. The present invention is designed to address this need. 
     Broadly speaking, the present invention provides a method and apparatus for non-invasively predicting (estimating) PEEP i . The presence of PEEP i  which often remains undetected, creates an inspiratory threshold load (a positive pressure level above ambient) that the patient&#39;s inspiratory muscles must overcome before fresh gas can enter the lungs. In addition, excessive levels of PEEP i  can lead to impaired cardiac function, an increased risk of barotrauma (structural lung damage secondary to excessive lung volumes), reduced inspiratory muscle pressure-generating capacity, and abnormally increased work of breathing. 
     The inventors have innovatively developed multiple indicators to detect and quantify PEEP i  non-invasively. Although some of these markers can be used alone to predict PEEP i , it is observed that some markers are good primarily for detection, while others are optimal for quantification. Furthermore, it is also noticed that certain markers work best on a specific group of subjects. Therefore, in an aspect of the invention, using several of these markers simultaneously has been demonstrated to improve the prediction of PEEP i  and broadens its operability to a wider range of patients. 
     The primary advantage is that the method is non-invasive, breath-to-breath, and in real-time. Therefore, there is no need to insert a catheter into the patient. The method requires only that measurements be made at the subject&#39;s airway, including one or more of the following: pressure at the airway, end tidal CO 2 , and airflow, measured at the mouth. These airway measurements are routinely monitored in most ICUs—by using any of the common respiratory equipment currently used to monitor and maintain patients while they receive ventilatory support. 
     Although the occlusion method is essentially non-invasive, it nevertheless requires a potentially irritating “intervention” and does not work for spontaneously breathing patients. This method does not suffer from either of these difficulties. There is no need to occlude the patient&#39;s airway and, it can be determined continuously without intervention. 
     The approach includes the measurement of several different PEEP i  markers which in turn are fed into a specialized mathematical model (for example, a neural network) that then predicts a PEEP i  value, based on these values. The model can be a linear model (e.g., multiple regression) or a nonlinear model (e.g., a neural network, fuzzy logic, etc.). 
     The PEEP i  markers or indicators may include: 
     Flow/volume trajectory—the concept of plotting a flow/volume waveform has been experimentally proven to be a useful quantification maker that works in the many patients. 
     CO 2  flow/volume trajectory—this marker has been experimentally proven to be an indicator of PEEPi. It is particularly useful for identifying patients with elevated levels of PEEP, and for those suffering from premature airway collapse or expiratory flow limitation. 
     CO 2 /volume ratio—this marker has been experimentally proven to be an indicator for detecting PEEP i  in patients exhibiting severe expiratory flow limitation patients. 
     Flow at the onset of inspiratory effort—this marker has been experimentally proven to useful for quantifying PEEPi. 
     Modeling the patient&#39;s expiratory waveform—this can be accomplished using two different methods: 1) least squares analysis and, 2) exponential modeling. Both methods have been experimentally proven to be useful in quantifying PEEP i . 
     Peak to Mid-Exhalation Flow Ratio: flow limitation patients may have high peak flows that decay extremely rapidly. When the mid-exhalation flow is low, this parameter is high, indicating likely flow limitation and a resultant PEEP i . 
     Duration of low exhaled flow: flow limitation patients may have an abnormally high percentage of their exhalation occur at low flows. This parameter determines the percentage of the exhalation (by volume or time) that occurs at low flows. 
     Capnograph waveform shape: flow limitation patients may have a identifiably unique CO 2  waveform versus normal patients. In particular, a Phase III slope (the slope after the rapid rise time in CO 2 ) is not horizontal and rises throughout exhalation, indicating a continuing low-flow exhalation. 
     Negative expiratory pressure or an increased expiratory gradient: a patient&#39;s expiratory flow rate should be directly proportional to pressure gradient (difference in pressure between the lungs and the pressure in the breathing circuit) during exhalation. If there is no pressure (in the breathing circuit) during exhalation, the gradient can best be increased by briefly applying a negative pressure, during the exhalation phase, in the breathing circuit. For patients receiving additional pressure during exhalation, the gradient can be increased by briefly removing the added pressure. If the expired gas flow does not increase in response to the increased gradient, the patient suffers from expiratory flow limitation and is exceptionally prone to developing PEEP i . 
     In addition to the above mentioned markers, the concept of using two or more of these and/or potentially other markers for PEEP i  as inputs to a mathematical model (linear or non-linear, such as a neural network) for predicting PEEP i  is new and unique. The combined information provided by using two or more markers (those mentioned above and others) may result in a more accurate estimate of PEEP i . 
     In one aspect of the invention, the method comprises creating a mathematical model of the patient&#39;s PEEP i  using predetermined parameters that are collected non-invasively, such as those collected with standard respiratory monitors. The respiratory monitors typically contain airway pressure and airway flow sensors that measure the flow and pressure of gases going into and out of the patient as a function of time. From these time versus flow or pressure waveforms, a variety of parameters are selectively derived that are used in characterizing different aspects of the patient&#39;s breathing and/or the patient&#39;s interaction with the ventilator. These parameters contain information that is extracted to accurately estimate the intrinsic PEEP. 
     More specifically, the method of the invention comprises a method of estimating the actual PEEP i  using a combination of multiple parameters derived from the aforementioned sensors, as well as others attached to the patient and/or ventilator. The PEEP i  parameter can be any parameter that represents the amount of excess pressure in the lungs, including but not limited to flow/volume trajectory, CO 2  flow/volume trajectory, CO 2 /volume ratio, flow onset, modeling on expiratory waveform, and peak to mid-exhalation flow ratio. 
     This method includes using a linear combination of parameters or a nonlinear combination of parameters, including but not limited to a neural network, fuzzy logic, mixture of experts, or polynomial model. Moreover, multiple different models can be used to estimate the PEEP i  of different subsets of patients. These subsets can be determined by various means, including but not limited to patient condition (pathophysiology), patient physiologic parameters (lung resistance and compliance), or other parameters. 
     In an aspect of the invention, the method for estimating PEEP i  in a patient comprises use of a neural network, wherein the neural network provides PEEP i  information for the patient based upon input data, wherein the input data includes at least one of the following parameters: flow/volume trajectory, CO 2  flow/volume trajectory, CO 2 /volume ratio, flow onset, modeling on expiratory waveform, peak to mid-exhalation flow ratio, duration of low flow exhalation, and capnograph waveform shape, wherein the intrinsic PEEP information is provided as an output variable. 
     In the above-noted method, the neural network is trained by collecting clinical data from a test population of patients used to obtain teaching data, the teaching data comprising the above-noted input information, inputting the teaching data into the neural network, whereby the neural network is trained to provide an output variable corresponding to the intrinsic PEEP. 
     As a system for estimating intrinsic PEEP in a patient, the system comprises a neural network which first receives as input primary teaching data obtained from clinical testing of a test population of patients, whereby the neural network learns the teaching data and is trained to provide an output variable for intrinsic PEEP, such that when said neural network receives patient input data in the form of the above-noted parameters obtained from a patient, the neural network provides the output variable for estimating intrinsic PEEP for that patient. 
     The invention can be implemented in numerous ways, including: as a system (including a computer processing or database system), as a method (including a computerized method of collecting and processing input data) and, as a method for evaluating such data to provide an output(s), an apparatus, a computer readable medium, a computer program product, or a data structure tangibly fixed in a computer readable memory. Several embodiments of the invention are discussed below. 
     As a system, an embodiment of the invention includes a processor unit having input and output devices. The processor unit operates to receive input parameters, process the input and provide an output corresponding to PEEPi. This output can be then used to control external devices, such as a ventilator. The processing of the data can be accomplished by various means such as neural networks, parallel distributed processing systems, neuromorphic systems, or the like. 
     As a method of predicting PEEP i , the method includes processing predetermined input variables (parameters), preferably through the use of a neural network. 
     As a computer readable media containing program instructions, an embodiment of the invention includes: computer readable code devices for receiving input variables, processing the input, and providing an output indicative of PEEP i . In a preferred embodiment, processing comprises utilizing a neural network. The method may further include controlling a ventilator in response to the output obtained. 
     The methods of the present invention may be implemented as a computer program product with a computer-readable medium having code thereon. The program product includes a program and a signal bearing media bearing the program. 
     As an apparatus, the present invention may include at least one processor, a memory coupled to the processor, and a program residing in the memory which implements the methods of the present invention. 
     Other aspects and advantages of the invention will become apparent from the following detailed description taken in conjunction with the accompanying drawings, illustrating, by way of example, the principles of the invention. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
       In order that the manner in which the above-recited and other advantages and objects of the invention are obtained, a more particular description of the invention briefly described above will be rendered by reference to specific embodiments thereof which are illustrated in the appended drawings. Understanding that these drawings depict only typical embodiments of the invention and are not therefore to be considered to be limiting of its scope, the invention will be described and explained with additional specificity and detail through the use of the accompanying drawings in which: 
         FIG. 1  depicts a method of one aspect of the invention for a patient on a ventilator. 
         FIG. 2 a    depicts a graph of a flow/volume loop in subjects without PEEP i . 
         FIG. 2 b    depicts a graph of a flow/volume loop in subjects with PEEP i . 
         FIG. 3  depicts a temporal plot of rising CO 2  during exhalation. 
         FIG. 4 a    depicts a graph of flow CO 2 /volume loop in subjects without PEEP i . 
         FIG. 4 b    depicts a graph of flow CO 2 /volume loop in subjects with PEEP i . 
         FIG. 5  depicts a graph of isolation of pressure drop due to airway. 
         FIG. 6  depicts a simplified electrical circuit representing a patient&#39;s respiratory system when receiving support provided by a ventilator. 
         FIG. 7  depicts an exemplary neural network architecture 
         FIG. 8  depicts inputs and outputs of an adaptive system having backpropagation. 
         FIGS. 9-12  show graphs of experimental data derived for multiple patients showing the results of exemplary systems using both linear and nonlinear models (neural network), and using multiple patients, some experiencing flow limitations and some experiencing no flow limitations, and a subset of patients experiencing no flow limitations. 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     Referring now to the drawings, the preferred embodiment of the present invention will be described. 
     In the embodiment depicted in  FIG. 1 , a patient  10  requiring respiratory support is connected to a ventilator  12  via a ventilator circuit  16  to communicate a flow of gas with an airway of a patient. Sensors  14  are provided such as an airway flow and pressure sensor, along with possibly a carbon dioxide detector attached to the circuit such as at the y-piece of the standard ventilator circuit or at the breathing tube of a single limb system such as a typical NIV ventilator. These sensors measure the flow, pressure, and partial pressure of carbon dioxide in the gases that pass to and from the patient. These raw signals  18  may be preprocessed in a signal processor  20  using analog and digital signal processing to clean the signal, remove sensor biases and offsets, etc. These signals can then be processed in a parameter extraction module  22  to calculate a variety of other parameters from the flow, pressure, and CO 2  data and identify indicators, or markers indicative of PEEP i . In an aspect of the invention, a neural network  24  may be provided to model the parameters so that a ventilator may be controlled through controller  26 . The interface appliance connecting the ventilator to the patient may be configured to engage the airway of subject without an intervening appliance and may include one or more of a nasal mask, a nasal/oral mask, a full face mask, a total face mask, a partial rebreathing mask, or other interface appliances that communicate a flow of gas with an airway of a subject. Sensors may be in fluid communication with the breathing circuit and/or the interface appliance. 
     The approach to measuring PEEP i  relies on monitoring several different patient parameters in real-time. The concept entails measuring the “markers” that indicate the presence of PEEP i , and feeding those qualified markers into a neural network, linear regression model, or the like. A value for PEEP i  is then predicted by using all of the different markers detected using methods best described as akin to a neural network, linear multiple regression modeling, or nonlinear multiple regression modeling. The models may be “pre-trained” using actual clinical data collected from patients suffering with varying degrees of PEEP i -levels that have been measured as accurately as possible using an esophageal balloon or other gold standard such as an esophageal catheter with electronic pressure transducers. PEEP i  measured via the esophageal pressure technique (PEEP i,pes ) is considered a reference or, true PEEP i . The PEEP i,pes  is used for training of the model as well as for validation of the approach. The model is trained to predict the actual PEEP i  using the PEEPi markers as input parameters. 
     In certain embodiments, particularly those where input parameters/markers are collected via a non-invasive patient interface such as a mask or canula where circuit leaks are significant, compensation for leak flow is done to derive accurate pressure, flow, volume, and/or CO2 waveforms. When the leak is significant, the flow in and out of the patient is much different than that measured by the sensor since much of the air is lost through the leak. The goal of leak compensation is to accurately estimate the true values of the waveforms moving in and out of the patient. There are numerous methods for correcting for leak. Examples of correcting flow and volume waveforms include, but are not limited to, (a) equally redistributing the measured lost volume over the breath and subtracting the bias flow from flow waveform, (b) calculating parabolic conductance and using the parabolic conductance to correct a measured flow waveform, (c) calculating the missing volume between inhalation and exhalation and allocating this volume throughout the breath, and (d) adjusting the flow waveform to minimize the average flow. Similarly, the CO2 content in the exhaled gas may be washed out by leaks. Similar methods can be used to determine the amount of CO2 that was exhaled based on the CO2 levels measured and the estimated leak flow. Leak correction modifies the pressure, flow, volume, and/or CO2 waveforms on a sample by sample basis based on the parameters (e.g. parabolic conductance) calculated from the input waveforms. These parameters can be calculated and adjusted continuously on a sample by sample basis, or calculated periodically based on a windowing approach that may include overlap in the windows. Once the waveforms are corrected, the PEEPi markers or parameters can be calculated accurately and utilized to determine non-invasive PEEPi. 
     Examples of PEEP i  markers include: Sudden flow reversal marking end-exhalation; Frequent volume channel “resets” at end-exhalation. This occurs because modern volume measuring equipment starts from zero volume at each breath. Spikes at onset of expiratory flow accompanied by at least two distinct expiratory flow decay patterns. Continuous increase in end-tidal carbon dioxide (ETCO 2 ) regardless of the expiratory time. High total respiratory system resistance and compliance, along with high breathing rates or elevated tidal volumes; since the product of total resistance and compliance equals the time constant for the lungs (60% of the volume above the V r  will be exhaled in the interval of one time constant), the greater the time constant, the greater the chance the patient will exhibit PEEP i , particularly when breathing rapidly or when taking large breaths. 
     Additional markers have been discovered, forming the backbone of the invention, that provide information related to the magnitude and type of PEEP i . These markers estimate PEEP i  based on flow/volume trajectory, carbon dioxide (CO 2 ) flow/volume trajectory, CO 2 /volume ratio, expiratory flow at onset of inhalation, and modeling on expiratory waveform. Many of these markers are unique and by themselves can be used to measure PEEPi. However, a combination of two or more of the markers may provide a more robust and accurate measure. 
     Now, referring to  FIGS. 2 a  and 2 b   , it follows that the flow/volume trajectory relies on analyzing the expiratory flow and volume of each breath. Under normal conditions, a plot of expiratory flow versus volume results in a nearly straight line  28  that intersects the volume axis  29  at approximately zero ( FIG. 2 a   ). That is, at end-exhalation the volume of gas coming out of the lungs is zero. The slope of the flow/volume trajectory is related to the average time constant of the lungs. A typical flow/volume loop for patients with PEEP i  is shown in  FIG. 2 b   . In this particular case, the trajectory line  30  intersects the volume axis  29  at well below zero; this indicates that had the exhalation phase continued, an additional 0.24 L (y-axis intercept  27 ) of gas would have been expelled from the lungs. Dividing the additional gas volume by the patient&#39;s respiratory compliance yields a quantifiable inference of the PEEP i  pressure. 
     CO 2  flow/volume trajectory is similar to the flow/volume trajectory, except CO 2  flow  36  is plotted as shown in  FIG. 3  instead of exhaled flow. CO 2  flow is obtained by multiplying the exhaled CO 2  and exhaled flow. In some PEEP i  patients, the exhaled CO 2  tends to continue to rise when there is very minimal exhalation flow (FIG.  3 ). The CO 2  flow parameter captures this rising CO 2  trend  38 , and when plotted against volume as shown in  FIG. 4 b   , results in a trajectory  34  that often parallels the volume axis  33  for PEEP i  patients.  FIGS. 4 a  and 4 b    illustrate a comparison between patients that have and do not have PEEP i , respectively. In non-PEEP i  patients as shown in  FIG. 4 a   , the trajectory  32  eventually intersects near the point where the volume  33  and flow  35  axes meet. The slope of the trajectories  32 ,  34  can provide an indication to the severity of PEEP i , where steep slopes such as shown in  FIG. 4 b    indicate severe PEEP i  and shallow slopes such as shown in  FIG. 4 a    indicate low levels of PEEP i . 
     CO 2 /volume ratio, another PEEP i  marker, is a fractional value of exhaled CO2 divided by exhaled volume. The maximum exhaled CO2 value and the change in volume during exhalation are computed for each breath. The ratio is given by:
 
CO 2 /volume ratio =max(ETCO 2 )/volume exhaled  
 
     It has been observed that PEEP i  patients have a larger ratio value versus those patients who do not have PEEP i . 
     The fourth of the additional PEEP i  markers, expiratory flow at the onset of inhalation, attempts to capture the exhaled flow rate at precise moment of end-exhalation by locating the onset of an inhalation effort. If gas is still flowing out of the lungs at the onset of inhalation, it can be reasoned that the only force driving this gas flow, at this instant in time, is PEEP i . 
     PEEP i  at the onset of inhalation (PEEP i,onset ), is estimated using the product of the expiratory airflow at inhalation onset (flow onset  and resistance to airflow produced by the airways of the lungs (R aw ).
 
PEEP i,onset =flow onset   ×R   aw  
 
     Total respiratory resistance (R total ) traditionally determined by programming the patient&#39;s ventilator to produce an end-inspiratory pause (usually the pause lasts 0.5 sec. or more) when delivering any mandated breath—these are breaths where the operator, not the patient, determines the gas flow rate, the gas flow pattern, the tidal volume and the frequency at which they are delivered/min. During each mandated breath, the difference between the peak inflation pressure (PIP) and the plateau pressure (P plat ) is determined. The difference is divided by the airflow measured at the moment the PIP was observed. It is also traditional to perform this measurement using a square flow pattern and with the gas flow programmed at or very near to 60 liters/min (1 liter/sec). This is done because resistance is defined as the pressure drop (measured in cm H 2 O) when gas is flowing at precisely 60 liters/min (1 liter/sec). Symbolically, resistance is determined as follows:
 
 R   aw =(PIP− P   plat )/flow.
 
     As defined above, resistance, computed in this manner, represents the total resistance of the respiratory system, and is acceptable as long as both lungs have similar resistance values. 
     During an end-inspiratory pause, since gas cannot escape from the lungs, it gradually redistributes from the hyper-inflated alveoli to under-inflated alveoli—a process called pendelluft. 
     Resistance to airflow (only) can be isolated from most PIP to P plat  pressure differences, by analyzing the flow versus pressure loop for that specific breath. In the presence of pendelluft, the expiratory side of the loop contains two distinct flow-pressure decay rates or slopes ( FIG. 5 ). The first slope  40 , in which pressure changes very rapidly (slope  1   40  in  FIG. 5 ), is due to the resistance the airways produce; the second slope  42  (slope  2  in  FIG. 5 ), the much slower rate of pressure change, is from the redistribution of gases. The difference in pressure drop between PIP and the pressure obtained from first decay rate (P plat,new ) represents the differential airway pressure used to determine airflow resistance only. R aw  is then obtained by dividing difference between the PIP and P plat,new  by the measured flow at the moment PIP was reached. 
     Another marker is a mathematical modeling of an expiratory waveform that estimates the respiratory system time constant (the product of resistance and compliance) changes during the course of exhalation. The concept is to get a better measure of system dynamics and, to predict PEEP i . Two modeling techniques have been explored: 1) estimate of system time constant and lung compliance using least squares, and 2) modeling resistance using an exponential function. Both methods rely on similar principles. 
     The respiratory system and a patient&#39;s ventilator can be represented using an electrical circuit diagram ( FIG. 6 ). Table 1 below lists the definition of terms used in  FIG. 6 : 
     
       
         
               
               
             
           
               
                   
                 TABLE 1 
               
               
                   
                   
               
             
             
               
                   
                 C—lung compliance 
               
               
                   
                 V l —lung volume 
               
               
                   
                 f—flow 
               
               
                   
                 R aw —airway resistance 
               
               
                   
                 R ex —ventilator exhalation valve resistance 
               
               
                   
                 P l —lung pressure 
               
               
                   
                 P aw —airway pressure 
               
               
                   
                   
               
             
          
         
       
     
     Airway pressure and flow are measured at the patient mouth by a differential pressure transducer. The corresponding gas volume (V nico ) from the measured flow is computed by integration of flow over the exhalation time period. The ventilator exhalation resistance changes based on ventilator PEEP setting to maintain P aw  at PEEP pressure at the end of exhalation. 
     With reference to  FIG. 6 , an estimate of system time constant and lung compliance using least squares as follows:
         During exhalation, flow is defined as:
 
 f ( t )=( P   l ( t )− P   aw ( t ))/ R   aw ,
   the lung pressure can be represented as lung volume divided by lung compliance, so flow can be written as:       

     
       
         
           
             
               
                 f 
                 ⁡ 
                 
                   ( 
                   t 
                   ) 
                 
               
               = 
               
                 
                   ( 
                   
                     
                       
                         
                           V 
                           l 
                         
                         ⁡ 
                         
                           ( 
                           t 
                           ) 
                         
                       
                       C 
                     
                     - 
                     
                       
                         P 
                         aw 
                       
                       ⁡ 
                       
                         ( 
                         t 
                         ) 
                       
                     
                   
                   ) 
                 
                 / 
                 
                   R 
                   aw 
                 
               
             
             , 
           
         
       
         
         
           
             and rearranged to represent lung volume as:
 
 V   l ( t )=τ* f ( t )+ C*P   aw ( t ),
 
             where τ is the time constant and defined as:
 
τ= R   aw   C.  
 
             The lung volume above the functional residual capacity (FRC) of the lung can be approximated as the summation of the actual measured gas volume inhaled by patient, volume due to PEEP, and any trapped gas.
 
 V   l ( t )= V   nico ( t )+ V   PEEP   +V   PEEPi  
 
             So, lung volume can now be described as:
 
 V   nico ( t )τ* f ( t )+ C*P   aw ( t )− V   PEEP   −V   PEEPi  
 
             Since volume due to PEEP and PEEP i  are constant, they can be eliminated from the equation by observing only differential changes.
 
Δ V   nico   =τ*Δf+C*ΔP   aw  
 
             The time constant and compliance are solved by least squares analysis. Volume due to PEEP and PEEP i  can be computed as:
 
 V   PEEP   +V   PEEPi   =τ*f ( t )+ C*P   aw ( t )− V   nico ( t )
 
             PEEP i  pressure can then be easily computed from: 
           
         
       
    
     
       
         
           
             
               PEEP 
               i 
             
             = 
             
               
                 
                   
                     V 
                     PEEP 
                   
                   + 
                   
                     V 
                     PEEPi 
                   
                 
                 C 
               
               - 
               PEEP 
             
           
         
       
     
     Modeling resistance using an exponential function is performed as follows:
         During exhalation, lung pressure can be described as:
 
 P   l ( t )= P   aw ( t )+ R   aw   *f ( t )
   Flow can be described by an exponential decaying waveform during exhalation as:
 
 f ( t )= f   0   *e   −t/R     aw     C  
   and solved for R aw  by:       

     
       
         
           
             
               R 
               aw 
             
             = 
             
               - 
               
                 t 
                 
                   C 
                   * 
                   
                     ln 
                     ⁡ 
                     
                       ( 
                       
                         
                           f 
                           ⁡ 
                           
                             ( 
                             t 
                             ) 
                           
                         
                         / 
                         
                           f 
                           0 
                         
                       
                       ) 
                     
                   
                 
               
             
           
         
       
         
         
           
             The lung pressure can then be described as: 
           
         
       
    
     
       
         
           
             
               
                 P 
                 l 
               
               ⁡ 
               
                 ( 
                 t 
                 ) 
               
             
             = 
             
               
                 
                   P 
                   aw 
                 
                 ⁡ 
                 
                   ( 
                   t 
                   ) 
                 
               
               - 
               
                 
                   f 
                   ⁡ 
                   
                     ( 
                     t 
                     ) 
                   
                 
                 * 
                 
                   t 
                   
                     C 
                     * 
                     
                       ln 
                       ⁡ 
                       
                         ( 
                         
                           
                             f 
                             ⁡ 
                             
                               ( 
                               t 
                               ) 
                             
                           
                           / 
                           
                             f 
                             0 
                           
                         
                         ) 
                       
                     
                   
                 
               
             
           
         
       
         
         
           
             From this equation, the PEEP i  pressure can be estimated by calculating the difference in lung pressure between the inhalation onset (t onset ) and location at zero flow (t end ).
 
PEEP i   =P   l ( t   onset )− P   l ( t   end )
 
             Inhalation onset can be detected as described previously (flow onset marker) or by observing the P i  itself, which goes through a sudden slope change at onset. 
             This PEEP i  estimate was done based on the assumption that 1) lung compliance remains fixed during exhalation, and 2) flow during exhalation decays exponentially. A slight variation of this method assumes that resistance remains fixed during exhalation, instead of lung compliance. In this case, the lung pressure is defined by modeling the lung compliance. 
           
         
       
    
     Peak to Mid-Exhalation Flow Ratio. This marker is calculated by dividing the peak exhalation flow by the flow calculated when about 20% to 30%, and preferably, about 25%, of the tidal volume remains in the lungs (75% has been exhaled). Flow limitation patients have high peak flows that decay very rapidly. This parameter will be large when the exhaled flow decays very quickly, indicating flow limitation. 
     As described above, developing a non-invasive indication of PEEPi, possibly including different levels of severity, is also advantageous. Many clinicians do not need to know the exact value of PEEPi, but instead just need to know whether the patient falls into one of several categories of intrinsic PEEP. For example, without restriction, a patient may have either: a) no PEEPi, b) mild PEEPi, c) moderate PEEPi, or d) severe PEEPi. Since methods to remove PEEPi are often inexact and not easily titrated, clinicians may only choose to treat cases of severe (or possibly) moderate PEEPi. Therefore, the clinical goal, for example, would be to ensure that all patients were in categories (a) or (b). 
     In addition, it may not be possible, particularly in ventilatory scenarios with large leaks, to accurately quantify PEEPi. However, the PEEPi markers will indicate the severity of the PEEPi but not necessarily quantify it. In an example embodiment, the patient is connected to a mask and NIV ventilator. At the output of the mask, the flow/pressure/CO2 sensors are connected to create the waveforms. The pressure, flow, volume, and CO2 waveforms are first leak compensated by first estimating the leak in the circuit and then correcting for it. Several methods of leak compensation are possible. In one embodiment the leak is related to the difference between the inhaled and exhaled volumes and flow is subtracted from the inhaled waveform and added to the exhaled waveform in proportion to the pressure in the circuit, since leak through an orifice is proportional to pressure. In a second method, an equation of the leak is derived that includes a parabolic conductance term. Parabolic conductance is estimated throughout the breath by dividing flow by pressure raised to the nth power, where n is related to the laminar flow and is 0.5 in this example (indicating turbulent flow). The leak corrected flow is then estimated by multiplying the estimated parabolic conductance times the square root of pressure. 
     Next, the PEEPi parameters are calculated, although due to the potential imperfection of the leak compensation, they may not be as accurate as is required to predict the actual PEEPi Value. Next the PEEPi parameters are used to determine the severity of the PEEPi. In the preferred embodiment, the PEEPi parameters are input to a mathematical model that is trained with clinical data to determine the severity of the PEEPi. The model can create discrete outputs with PEEPi categories, or the model can create a continuous output with a post processing step that thresholds the results of the model to create the PEEPi categories. The model can be a regression model, polynomial model, neural network, nonlinear model, mixture of experts, information theoretic model (such as support vector machine), or other model known in the art. 
     Description of Neural Networks 
     Artificial neural networks loosely model the functioning of a biological neural network, such as the human brain. Accordingly, neural networks are typically implemented as computer simulations of a system of interconnected neurons. In particular, neural networks are hierarchical collections of interconnected processing elements (PEs). These elements are typically arranged in layers, where the input layer receives the input data, the hidden layers transform the data, and the output layer produces the desired output. Other embodiments of a neural network can also be used. 
     Each processing element in the neural network receives multiple input signals, or data values, that are processed to compute a single output. The inputs are received from the outputs of PEs in the previous layer or from the input data. The output value of a PE is calculated using a mathematical equation, known in the art as an activation function or a transfer function that specifies the relationship between input data values. As known in the art, the activation function may include a threshold, or a bias element. The outputs of elements at lower network levels are provided as inputs to elements at higher levels. The highest level element, or elements, produces a final system output, or outputs. 
     In the context of the present invention, the neural network is a computer simulation that is used to produce a noninvasive estimate of the quantified intrinsic PEEP described previously. The neural network of the present invention may be constructed by specifying the number, arrangement, and connection of the processing elements which make up the network. A simple embodiment of a neural network consists of a fully connected network of processing elements. As shown in  FIG. 7 , the processing elements of the neural network are grouped into layers: an input layer where the parameters collected and/or derived from the airway pressure and flow sensors are inputted to the network; a hidden layer of processing elements; and an output layer where the resulting prediction of intrinsic PEEP is produced. The number of connections, and consequently the number of connection weights, is fixed by the number of elements in each layer. 
     The most common training methodology for neural networks is based upon iterative improvement of the system parameters (normally called weights) by minimizing the mean squared difference between the desired output and the network output (mean squared error, MSE). The input is applied to the neural network, the neural network passes the data through its hierarchical structure, and an output is created. This network output is compared with the desired output corresponding to that input and an error is calculated. This error is then used to adjust the weights of the system so that the next time that particular input is applied to the system the network output will be closer to the desired output. There are many possible methodologies to adjust the weights, called the training algorithm. As shown in  FIG. 8 , the most common is called backpropagation that involves calculating each weight&#39;s responsibility for the error, and calculating a local gradient from this error in order to use a gradient descent learning rule for each weight. 
     An exemplary system for implementing the invention includes a computing device or a network of computing devices. In a basic configuration, computing device may include any type of stationary computing device or a mobile computing device. Computing device typically includes at least one processing unit and system memory. Depending on the exact configuration and type of computing device, system memory may be volatile (such as RAM), non-volatile (such as ROM, flash memory, and the like) or some combination of the two. System memory typically includes operating system, one or more applications, and may include program data. Computing device may also have additional features or functionality. For example, computing device may also include additional data storage devices (removable and/or non-removable) such as, for example, magnetic disks, optical disks, or tape. Computer storage media may include volatile and non-volatile, removable and non-removable media implemented in any method or technology for storage of information, such as computer readable instructions, data structures, program modules or other data. System memory, removable storage and non-removable storage are all examples of computer storage media. Computer storage media includes, but is not limited to, RAM, ROM, EEPROM, flash memory or other memory technology, CD-ROM, digital versatile disks (DVD) or other optical storage, magnetic cassettes, magnetic tape, magnetic disk storage or other magnetic storage devices, or any other physical medium which can be used to store the desired information and which can be accessed by computing device. Any such computer storage media may be part of device. Computing device may also have input device(s) such as a keyboard, mouse, pen, voice input device, touch input device, etc. Output device(s) such as a display, speakers, printer, etc. may also be included. Computing device also contains communication connection(s) that allow the device to communicate with other computing devices, such as over a network or a wireless network. By way of example, and not limitation, communication connection(s) may include wired media such as a wired network or direct-wired connection, and wireless media such as acoustic, RF, infrared and other wireless media. 
     Computer program code for carrying out operations of the invention described above may be written in a high-level programming language, such as C or C++, for development convenience. In addition, computer program code for carrying out operations of embodiments of the present invention may also be written in other programming languages, such as, but not limited to, interpreted languages. Some modules or routines may be written in assembly language or even micro-code to enhance performance and/or memory usage. It will be further appreciated that the functionality of any or all of the program modules may also be implemented using discrete hardware components, one or more application specific integrated circuits (ASICs), or a programmed digital signal processor or microcontroller. A code in which a program of the present invention is described can be included as a firmware in a RAM, a ROM and a flash memory. Otherwise, the code can be stored in a tangible computer-readable storage medium such as a magnetic tape, a flexible disc, a hard disc, a compact disc, a photo-magnetic disc, a digital versatile disc (DVD). The present invention can be configured for use in a computer or an information processing apparatus which includes a memory, such as a central processing unit (CPU), a RAM and a ROM as well as a storage medium such as a hard disc. 
     The “step-by-step process” for performing the claimed functions herein is a specific algorithm, and may be shown as a mathematical formula, in the text of the specification as prose, and/or in a flow chart. The instructions of the software program create a special purpose machine for carrying out the particular algorithm. Thus, in any means-plus-function claim herein in which the disclosed structure is a computer, or microprocessor, programmed to carry out an algorithm, the disclosed structure is not the general purpose computer, but rather the special purpose computer programmed to perform the disclosed algorithm. 
     A general purpose computer, or microprocessor, may be programmed to carry out the algorithm/steps of the present invention creating a new machine. The general purpose computer becomes a special purpose computer once it is programmed to perform particular functions pursuant to instructions from program software of the present invention. The instructions of the software program that carry out the algorithm/steps electrically change the general purpose computer by creating electrical paths within the device. These electrical paths create a special purpose machine for carrying out the particular algorithm/steps. 
     Unless specifically stated otherwise as apparent from the discussion, it is appreciated that throughout the description, discussions utilizing terms such as “processing” or “computing” or “calculating” or “determining” or “displaying” or the like, refer to the action and processes of a computer system, or similar electronic computing device, that manipulates and transforms data represented as physical (electronic) quantities within the computer system&#39;s registers and memories into other data similarly represented as physical quantities within the computer system memories or registers or other such information storage, transmission or display devices. 
     It should be understood that the examples and embodiments described herein are for illustrative purposes only and that various modifications or changes in light thereof will be suggested to persons skilled in the art and are to be included within the spirit and purview of this application.