WO2011117861A1

WO2011117861A1 - Differential lung functionality assessment

Info

Publication number: WO2011117861A1
Application number: PCT/IL2011/000247
Authority: WO
Inventors: Merav Gat; Adi Eldar; Konstantin Goulitski
Original assignee: Merav Gat; Adi Eldar; Konstantin Goulitski
Priority date: 2010-03-25
Filing date: 2011-03-15
Publication date: 2011-09-29

Abstract

A non-invasive, radiation-free system for providing differential pulmonary functionality comprising: a plurality of sound transducers adapted to be applied to a planar region of the chest or back skin of an individual to produce analog voltage acoustic signals indicative of pressure waves at each transducer location, an analog to digital converter connected with the transducers for converting the analog acoustic signals into digital form, an electronic processor connected with the analog to digital converter and a spirometry system connected with the processor.

Description

DIFFERENTIAL LUNG FUNCTIONALITY ASSESSMENT

FIELD OF THE INVENTION

This invention relates to medical devices and methods, and more particularly to such devices and methods for analyzing lung functionality.

CROSS-REFERENCE TO RELATED PATENT APPLICATIONS

This patent application claims priority from and is related to U.S. Provisional Patent Application Serial Number 61/282749, filed 25 March 2010, this U.S. Provisional Patent Application incorporated by reference in its entirety herein.

BACKGROUND OF THE INVENTION

Respiratory problems ail infants and adults alike. Some common lung diseases or conditions include asthma, Chronic Obstructive Pulmonary Disease (COPD), regional collapse (atelectasis), consolidation (e.g. pneumonia), interstitial edema, focal lung disease (e.g. tumor) and global lung disease (e.g. emphysema).

Regional assessment of lung physiology has been carried out using radionucleotide perfusion also known as the "VQ scan". In this technique, radioactive particles are either injected into the subject's blood system or the subject is allowed to inhale suspended radioactive particles. SPECT images of the lungs are obtained and one or both of the lungs in the image is divided into two or more regions. A separate analysis of each lung region is then performed. In most regional lung assessments, each of the two lung images is divided into three parts (top, middle and bottom), and an assessment of lung function or physiology in each region is obtained. Typically, regional assessment involves determining the fraction of the radioactivity detected in each region out of the total radioactivity from both lungs. The amount of radioactivity detected in each part may be con-elated with the lung condition in each part.

Body sounds are routinely used by physicians in the diagnosis of various disorders. A physician may place a stethoscope on a person's chest or back and monitor the patient's breathing in order to detect adventitious (i.e. abnormal or unexpected) lung sounds. The identification and classification of adventitious lung sounds often provides important information about pulmonary abnormalities.

The process of fixing one or more microphones onto a subject's chest or back to record lung sounds is an established procedure. U.S. Patent No. 6,139,505 discloses a system in which multiple microphones are placed on a patient's chest. The recordings of the microphones during inhalation and expiration are displayed on a screen, or printed on paper. The recordings are then visually examined by a physician in order to detect a pulmonary disorder in the patient.

Kompis et al. (Chest, 120(4), 2001) disclose a system in which M microphones are placed on a patient's chest, and lung sounds are recorded. The recordings generate M linear equations that are solved using a least-squares fit. The solution of the system is used to determine the location in the lungs of the source of a sound detected in the recordings.

US Patent No. 6,887,208 to Kushnir et al, provides a system and method for recording and analyzing sounds produced by the respiratory tract. Respiratory tract sounds are recorded at a plurality of locations over an individual's thorax and the recorded sounds are processed to produce an image of the respiratory tract. The processing involves determining from the recorded signals an average acoustic energy at a plurality of locations over the thorax over a time interval from t to t2.

The term "acoustic energy" at a location is used herein to refer to a parameter indicative of or approximating the product of the pressure and the mass propagation velocity at that location. The image may be used to analyze respiratory tract physiology and to detect pathological conditions. Additionally, a time interval can be divided into a plurality of sub- intervals, and an average acoustic energy determined over the thorax for two or more of the sub-intervals. An image of each of these sub intervals may then be determined and displayed sequentially on a display monitor. This generates a movie showing dynamic changes occurring in the acoustic energy in the respiratory tract over the time interval. WO2007060663 to Kushnir et al, provides a system and method for regional assessment of lung functioning. In accordance with the invention, microphones are affixed to the body surface at a plurality of locations over the thorax, and signals indicative of lung sounds are recorded. The signals are analyzed in order to produce a value of a predetermined parameter at each of two or more locations on the body surface over the lungs. The two or more locations at which the parameter was determined are clustered into groups, where each group consists of locations on the body surface overlying a particular region of the lungs. The regions may correspond to anatomical regions of the lungs, or may be determined independently of the lung anatomy. For each group of locations, a regional assessment of the underlying lung region is obtained based upon the values of the parameter in the group. The regional assessment may be, for example, the sum of the values of the parameter at the locations in the group, the maximum value, the minimum value or an average value. Alternatively, the regional assessment may be the sum of the values of the parameter at the locations in the group divided by the sum of the values of the parameter in all of the groups. In one embodiment, each lung is divided into three regions (top, middle and bottom), and a regional assessment is obtained as explained above for each of the six regions. In another embodiment, the lungs are divided into regions so that each region has the same number of overlying microphones. The regional assessment may be presented in the form of a table.

US20080221467 to Papyan et al, provides a method and system for regional assessment in two or more regions of an individual's lungs. The system includes a plurality of transducers configured to be fixed over the thorax. Each transducer generates a signal P(xi,t) indicative of pressure waves at the location of the transducer. The transducers are divided into subsets, where each subset overlies a specific region of the two or more regions. An energy assessment signal is calculated from each of the signals P(xi,t). For each region, an assessment of the region is calculated from the energy assessment signals of the region.

A spirometer is a device used to monitor a person's ability to breathe out air. It measures the airflow through the bronchi and thus the degree of obstruction in the airways.

A spirometer (i.e. Spirometer. United States Patent 4462410 and Portable spirometer. United States Patent 5277195) is a device that measures how much (volume) and how fast (flow) air is moved into and out of the lungs. A computerized sensor (which is part of the spirometer) calculates and graphs the results. The results demonstrate a person's air flow rates or the volume forced out within the first second. This is the Forced Expiratory Volume in the first second (FEV1). This indicates whether or not there is airway obstruction. Spirometry also records the total volume of air forced out of the lungs. This is the Forced Vital Capacity (FVC).

In assessing lung function/mechanics, the physiology of the lungs should be taken into consideration. Generally, information is obtained by the use of a forced expiratory maneuver (with a Spirometer) that generates a maximum expiratory flow-volume curve. However, the essence of the flow volume curve is that in healthy individuals, the lung empties homogeneously during a maximally forced deflation. Nonetheless, such behavior would appear to be implausible if for no other reason than that airway structure is known to be substantially heterogeneous among parallel pathways of gas conduction. Furthermore, bronchial tree asymmetry is responsible for a non-gravitational component of regional volume variability. Reductions in flow from obstructed regions appear to be compensated by increases in flow from unobstructed regions and thus mask upstream non-uniformities. These mechanisms may explain in part why the maximal expiratory flow-volume curve has been a relatively insensitive tool for the detection of early non-uniform airway disease.

It would be advantageous to provide a differential spirometry system which may indicate regional lung functionality. For example:

In pulmonary emphysema, it is often observed that alveolar destruction occurs non- homogeneously throughout the lung and that some lung regions are severely diseased, whereas other regions are relatively spared. This non-uniform process results in a lung in which mechanical properties are variable between regions.

Air-trapping in COPD: differential spirometry could demonstrate and quantify regional air trapping in asthma or COPD that will not be shown in a full forced vital capacity maneuver. Response to bronchodilator: differential FVC (Forced Expiratory Vital Capacity) can detect reasons for paradoxical response to bronchodilator.

Cystic Fibrosis: a quantitative measurement of localized airways obstruction and structural damage in the early stages of CF may be obtained.

SUMMARY OF THE INVENTION

According to a first aspect of the invention there is provided a non-invasive, radiation-free system for providing differential pulmonary functionality comprising: a plurality of sound transducers adapted to be applied to a planar region of the chest or back skin of an individual to produce analog voltage acoustic signals indicative of pressure waves at each transducer location; an analog to digital converter connected with said transducers for converting said analog acoustic signals into digital form; an electronic processor connected with said analog to digital converter; and a spirometry system connected with said processor.

According to a second aspect of the invention there is provided a method of providing differential pulmonary functionality, comprising: attaching a plurality of sound transducers to a planar region of the chest or back skin of an individual; using a spirometer, measuring the total volume of air inhaled and exhaled by the individual during at least one breathing cycle; acquiring analog voltage acoustic signals indicative of pressure waves produced at each transducer location during the same at least one breathing cycle; converting said analog acoustic signals into digital form; and using said total air volume and said analog acoustic signals to calculate the relative air flow in each area of the lungs.

BRIEF DESCRIPTION OF THE DRAWINGS For a better understanding of the invention and to show how the same may be carried into effect, reference will now be made, purely by way of example, to the accompanying drawings.

With specific reference now to the drawings in detail, it is stressed that the particulars shown are by way of example and for purposes of illustrative discussion of the preferred embodiments of the present invention only, and are presented in the cause of providing what is believed to be the most useful and readily understood description of the principles and conceptual aspects of the invention. In this regard, no attempt is made to show structural details of the invention in more detail than is necessary for a fundamental understanding of the invention, the description taken with the drawings making apparent to those skilled in the art how the several forms of the invention may be embodied in practice. In the accompanying drawings:

Fig. 1 is a schematic block diagram outlining the main components of the system according to the present invention;

Fig. 2 shows the correlation between total flow (x-axis) and acoustic envelope;

Fig. 3 is a flowchart describing an example algorithm for computing separated airflow of the lungs; and

Fig. 4 is a flowchart showing an example envelope calculation algorithm, based on recognition of peaks.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

The present invention discloses a system and method for providing a non-invasive, radiation-free imaging system that provides differential pulmonary function (split-lung) based on quantitative lung sound information in different regions of the lungs, such as disclosed in WO2007060663 combined with lung function assessed by spirometry.

Fig. 1 is a schematic block diagram outlining the main components of the system according to the present invention. The system 100 comprises a plurality of N sound transducers 105, of which four are shown, applied to a planar region of the chest or back skin of individual 110 (not shown in the drawing). The transducers 105 may be applied to the subject by any means known in the art, for example using an adhesive, suction, or fastening straps. Each transducer 105 produces an analog voltage signal 115 indicative of pressure waves arriving to the transducer. The analog signals 115 are digitized by a multichannel analog to digital converter 120.

The digital data signals P(x, , t) 125, represent the pressure wave at the location xi of the ith transducer (i= 1 to N) at time t. The data signals 125 are input to a memory 130. Data input to the memory 130 are accessed by a processor 135 configured to process the data signals 125. The signals 125 may be de-noised by filtering components. For example frequencies above or below the known range of lung sounds can be filtered out, as well as vibrations originating from external noise (due to speaking etc.). Each signal 125 may also be subject to band pass filtering so that analysis is done only on frequency components within a range of interest.

An input device, such as a computer keyboard 140 or mouse 145, is used to input relevant information relating to the examination such as personal details of the individual 110. The input device 140 may also be used to input values of one or more times t1 and t2 that specify times at which the signals P(Xj,t) are to be analyzed or that specify one or more time intervals over which no signals are to be analyzed. The processor 135 calculates the value of a parameter at a plurality of locations over the lungs at the specified times or over the specified time intervals.

The locations at which the parameter is calculated are divided into groups, where each group overlies a region of the lungs. The processor 135 is further configured to perform a regional assessment of the lungs. The regional assessment comprises for each of the groups determining the value of one or more regional parameters where each regional parameter is obtained in a calculation involving the parameter values calculated at the various locations in the region. For example, a regional parameter may be the sum of the parameters in the region, the maximum of the parameter value, the minimum or the average. The regional parameter values may be normalized by dividing the regional parameter by the sum of the regional parameter values.

System 100 additionally comprises a spirometer 160 connected to a mouthpiece 170. The spirometer 160 calculated parameters (flow, volume, time) are input to memory 130. The spirometer may comprise a spirometer.

System 100 may additionally comprise display means 150 for graphical or alpha-numerical display of the test results.

The main goal of system is to measure non-invasively the general air-flow of breathing, separated at least into two parts: left and right, i.e. the part that flows into the left lung and the part that flows into the right lung.

Air-flow in the air-ways of a lung produces vibrations which may be explained by turbulent flow. The source of these vibrations can be classified by its frequency response, which corresponds to the air-way's size: big air-ways excite vibrations at low frequency range while small air-ways excite vibrations at high frequency range. The vibration is caused by a kind of dissipation process having two main mechanisms operating when air flows into a lung's air-way. The first one is the characteristics of the response of the elastic walls of a lung's air-way to air flow which occurs due to the fact that the walls have non regular structure. In this case the non regular wall structure leads to local changes of the flow velocity and hence to local differences of pressure which apply force on the elastic walls of the air-ways at neighboring points with different cross section area. Due to the elasticity of the air-way's walls they react to the pressure, leading to instability and to walls vibration. This phenomenon corresponds mostly to the big size air-ways, where the irregular structure is strongly marked and it is decreased in small size air-ways. The second mechanism is the standard development of acoustic waves in a tube. As gas propagating through the air-way channel (the tube) reaches any obstacle (for example, bifurcation of the bronchial tree) it creates a reflected wave in the opposite direction. Thus, the sum of propagating and reflected waves leads to the appearance of standing waves in this channel. These standing waves generate vibration with specific frequency corresponding to the size of the air-way. This phenomenon generally appears at inspiration flow, because at expiration there are no obstacles in the air-flow way (air ways are just merged on the way out via the trachea).

These two mechanisms appear at breathing and are responsible for the frequency response of the vibration which is recorded from a human body: the maximum intensity at 150-250Hz which generally corresponds to trachea and principal bronchi, the linear decreasing of intensity toward the high frequencies, and the higher intensity of the inspiration phase relative to the expiration one.

In addition to those mechanisms, vibrations are also caused (or changed) due to local changes of the flow (that change the local static pressure on the airway walls). Therefore, in general, when the flow is increased the vibration is increased too.

The standard method to measure air flow is by a spirometer which is widely used in Spirometry. Intensity of vibration may be presented by an amplitude modulation envelope of a signal recorded by microphones which are placed on the chest or back of a subject during breathing. The algorithm of amplitude envelope calculation might differ according to the problem to be solved, but the main concept of this algorithm is to generate a smoothed curve which presents the average variations of the vibration intensity along time, without the effects of transient sounds and other noises which are not related to the air flow.

The following experiment was done to prove that intensity of vibration is correlated to the flow intensity: three healthy subjects were recorded by DeepBreeze supine V-array with a spirometer. The spirometer recorded general flow of air during breathing of the subject. Example of the correlation between the general flow and the intensity of the recorded vibration is presented in Fig. 2, in which correlation between total flow (x-axis) and acoustic envelope (y-axis) is presented for the average of three recordings which were done by supine V-array of one healthy subject. Each curve corresponds to an inspiration phase of one breathing cycle, recorded by one microphone.

Despite small distortions, it may generally be concluded from this experiment that the flow- vibration correlation exists and it is very similar at different microphone locations and different healthy subjects. A difference between correlations at different microphone locations might be explained by local changes of the lung; a difference between correlations of different healthy subjects might be explained by different body structure.

Consequently, the ratio between the average vibration envelopes of the left and right lungs can give the ratio between the flow of the left and right lungs providing that the total airways cross sections for the left and the right lung are similar.

To summarize, the differential spirometer of the present invention is based on the following observations:

1. The breath sound vibration envelope is correlated to the air flow.

2. In order to measure the separated air-flow, the general air-flow can be divided according to the ratio between the sum or average values of the left and the right envelopes.

3. The right/left flow separation will be more accurate providing that the air ways cross sections for the left and the right lung similar.

Fig. 3 is a flowchart describing an example algorithm for computing separated airflow of the lungs.

In step 300, the algorithm receives initial raw data from a matrix of N (e.g. 40) microphones placed on a subject's back or chest, which record vibrations during breathing and raw data from a spirometer, which measures total air flow rate of the subject's breathing. The location of each microphone is determined by a row number r and a column number c. The raw data of each microphone may be presented as Sig(r,c,t), where t denotes the time scale. The raw data from both microphones and spirometer may be provided to memory 130 with time stamps for synchronizing the two devices. Alternatively, the system may attach time stamps to the incoming signals automatically.

In step 310, the raw data from each microphone is filtered by a band pass filter, for example between 100 and 1000 Hz. Generally this filtering is done to remove energy of signals which correspond to heart sounds (low frequencies) or to noises (high frequencies).

In step 320, the start and end points of an inspiration phase of each recorded breath cycle are identified on the time scale of a record, by processing the raw data of the spirometer. Generally, the algorithm finds points of the flow data where it crosses the zero value. We shall refer to these points as where / is the cycle number and j equals 1 for a starting point of the inspiration phase and 2 for an end point of the inspiration phase.

In step 330, the signal envelopes are calculated. This calculation may have different variants. The envelope may be calculated by computing a standard deviation value in a sliding window, by computing the Hilbert transform of the signal, by implementing a median value in a sliding window or by a number of other ways.

The envelope may be calculated for the whole signal or for each cycle or even for each phase of a breathing cycle separately.

Fig. 4 is a flowchart showing an exemplary envelope calculation algorithm, based on peaks recognition. The envelope is calculated for each phase of the breathing cycle separately, but may also be calculated for entire breathing cycles or for a series of breathing cycles. In step 400, the algorithm calculates the absolute value of each signal in the time intervals that corresponds to the predetermined phase of a breathing cycle (i.e. the union of the inspiration and expiration intervals).

In step 410, all points corresponding to the maximal values (peaks) are found for each phase.

In step 420, the algorithm calculates a median filter for the vector of these peaks, with a sliding window whose length is equal to 10% (or any other percentage) of the length of the peaks vector.

In step 430, a running average is calculated for the resultant median peaks vector, with a sliding window whose length is equal to 10% (or any other percentage) of the length of the vector and uniform weighting coefficients equaling 1 over the window length.

Finally in step 440 the resulting (heavily filtered) peaks vector is interpolated (using cubic interpolation or similar) in order to generate the envelope vector Env(t) in the original time resolution

Returning to fig. 3, in step 340, summary or average envelopes are computed at each phase (inspiration and expiration) of the recorded breathing, corresponding to the left and right lungs. For this partition we define two subsets of the n microphones, Ni of Π; microphones on the left side and N_rof n_r microphones on the right side. Accordingly: AvgEnvLefl =— Env (t)

I jeN,

AvgEnvRig t =— ^ £n _y (i)

Where j - is the index of the microphone, which is running over ni microphones for the left side and over n_r for the right side.

In step 350 the flow/volume curves (that were retrieved from the spirometer) which correspond to a specific phase (either inspiration or expiration) for all the breathing cycles are averaged. For this purpose all curves are normalized for a [0,1] time interval with constant step, using a standard interpolation function (in this context 0 means start time of the phase, 1 means end time of the same phase). Subsequently all curves are summarized and divided by the number of cycles in the recorded signal. Then the averaged curve is interpolated again back to the original time scale:

m is the number of breathing cycles; j is the index of each cycle tnorm is the normalized time scale

Fj(t_mm)) is the time normalized air flow function in the specific phase in cycle j

A similar procedure of computation of average curves of envelopes is implemented for each side and breathing phase (of the m cycles):

AvgEnvLefi^,nsp(t) =— _jAvgEnvLeft_J ^nsp (t)

Finally, in step 360 the separated flow is calculated.

There are a number of techniques to calculate the separated flow, we shall describe two examples.

The first method is the most intuitive one: at each time point t we split the average flow F(t) by the ratio of the left and right lungs' vibration envelope. Formally:

⁽ AvgEnvRight¹^ (Q

F_emRight^Eip(t) = F(ti

AvgEnvLefi^Eip (t) + AvgEnvRight^ (t)

Where F_envLeft!^nsp(t) is the estimated left lung flow based on the relative acoustic envelope for the inspiration phase and similar formulas for the flow during the expiration phase and for the right lung.

The limitation of this method is that the coupling between the original flow (from the spirometer) and the vibration envelope is very tight and is forced for all time points. Consequently, the final shape of the regional flow volume curves (for each lung) is governed by the shape of the original flow volume curve, and any asymmetry between the lungs (that might be apparent from the vibration envelopes) is diminished.

An alternative method is to loosen that coupling by forcing a match between the vibration envelope and the flow only at the start and end times (t₀, tend) of the breathing phase. In order to be able to take into account all time points in between, we apply this technique on the integrals of the flow from to to nd which is actually the Volume. Formally, the total volume of inhaled/exhaled air is calculated by time integration of the total flow/volume curve at each phase of the recorded breathing.

Where 7)_Π5ρ is the average duration of the inspiration phase (over all cycles) and T_exp is the same average for the expiration phase.

The integral value of average envelope at each side is calculated by a similar integration procedure:

T T

V_emRig t^lnsp = AvgEnvRight^Insp(t) V_emRight^Exp ^ AvgEnvRight** (t)

1=0

Now we assume that the total volume is correlated to the above integral of the acoustic envelope and we force this correlation at t₀ and nd- As a result we can split the total flow to the left and right sides by the relative integral of the acoustic envelope for the left and right sides:

Where F_envLef(^ns (t) is the estimated left lung flow based on the relative acoustic envelope for the inspiration phase and similar formulas for flow during the expiration phase and the right lung.

It is appreciated that certain features of the invention, which are, for clarity, described in the context of separate embodiments, may also be provided in combination in a single embodiment. Conversely, various features of the invention which are, for brevity, described in the context of a single embodiment, may also be provided separately or in any suitable subcombination.

Unless otherwise defined, all technical and scientific terms used herein have the same meanings as are commonly understood by one of ordinary skill in the art to which this invention belongs. Although methods similar or equivalent to those described herein can be used in the practice or testing of the present invention, suitable methods are described herein.

All publications, patent applications, patents, and other references mentioned herein are incorporated by reference in their entirety. In case of conflict, the patent specification, including definitions, will prevail. In addition, the materials, methods, and examples are illustrative only and not intended to be limiting.

It will be appreciated by persons skilled in the art that the present invention is not limited to what has been particularly shown and described hereinabove. Rather the scope of the present invention is defined by the appended claims and includes both combinations and subcombinations of the various features described hereinabove as well as variations and modifications thereof which would occur to persons skilled in the art upon reading the foregoing description.

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Claims

1. A non-invasive, radiation-free system for providing differential pulmonary functionality comprising: a plurality of sound transducers adapted to be applied to a planar region of the chest or back skin of an individual to produce analog voltage acoustic signals indicative of pressure waves at each transducer location; an analog to digital converter connected with said transducers for converting said analog acoustic signals into digital form; an electronic processor connected with said analog to digital converter; and a spirometry system connected with said processor.

2. The system of claim 1 , additionally comprising filtering means connected with said processor.

3. The system of claim 1 , additionally comprising display means.

4. A method of providing differential pulmonary functionality, comprising: attaching a plurality of sound transducers to a planar region of the chest or back skin of an individual; using a spirometer, measuring the total volume of air inhaled and exhaled by the individual during at least one breathing cycle; acquiring analog voltage acoustic signals indicative of pressure waves produced at each transducer location during the same at least one breathing cycle; converting said analog acoustic signals into digital form; and using said total air volume and said analog acoustic signals to calculate the relative air flow in each area of the lungs.

5. The method of claim 4, wherein said area comprises an entire lung.

6. The method of claim 4, wherein said area comprises part of a lung.

7. The method of claim 4, wherein said calculation comprises: calculating an acoustic signals envelope for each transducer at each breathing phase (inspiration and expiration); calculating average acoustic signal envelopes for each said areas of the lungs, using information about the transducers locations; calculating the average air volume per each said breathing phases; and using said calculated average acoustic signal envelopes to determine differential air volume per each said breathing phase.

8. The method of claim 7, wherein said calculating an acoustic signals envelope comprises : calculating the absolute value of each acoustic signal produced at each said breathing phase; finding the peak absolute values; calculating a median filter for the vector of said peak values using a first sliding window; calculating a running average for the resultant median peaks vector using a second sliding window; and interpolating the resultant peaks vector to generate the acoustic signals envelope in the original time resolution.

9. The method of claim 8, wherein said first sliding window comprises X percent (e.g. 10%) of the length of said peak values vector.

10. The method of claim 8, wherein said second sliding window comprises X percent (e.g. 10%) of the length of said peak values vector.

11. The method of claim 7, wherein each said breathing phases comprises a complete breathing cycle.

12. The method of claim 7, wherein each said breathing phases comprises one of an inhaling phase and an exhaling phase.

13. The method of claim 7, wherein said determining differential air volume per each said breathing phase, comprises splitting the average air flow at each said breathing phases by the ratio of said acoustic signal envelopes in said different areas of the lungs.

14. The method of claim 7, wherein said determining differential air volume per each said breathing phase, comprises: integrating the average air volume per each said breathing phases over all the recorded breathing phase; integrating the average acoustic signal envelopes of each said areas of the lungs over all the recorded breathing phase; and splitting the air flow at each said breathing phases according to the relative integrated acoustic signals envelopes.