(en)A system for identifying heart valve signals includes a non-invasive sensor unit for capturing electrical signals and composite vibration objects, a memory containing computer instructions, and one or more processors coupled to the memory, where the one or more processors are configured to perform the steps of separating a plurality of individual heart vibration events from the composite vibration objects by using at least one among bin-wise clustering and permutation alignment, or non-negative matrix factorization, or deep belief networks and tagging the plurality of individual heart vibration events using at least one among principal component analysis, Gabor filtering, generalized cross correlation, phase transform, smoothed coherent transformation, Roth correlation or band filtering.

1.ApplicationNumber: US-201715397088-A
1.PublishNumber: US-2017188868-A1
2.Date Publish: 20170706
3.Inventor: KALE KAUSTUBH
Giraldo Luis Gonzalo Sanchez
PAVA DIEGO
ESFAHANIAN MAHDI

4.Inventor Harmonized: KALE KAUSTUBH(US)
Giraldo Luis Gonzalo Sanchez(US)
PAVA DIEGO(US)
ESFAHANIAN MAHDI(US)

5.Country: US
6.Claims:
(en)A system for identifying heart valve signals includes a non-invasive sensor unit for capturing electrical signals and composite vibration objects, a memory containing computer instructions, and one or more processors coupled to the memory, where the one or more processors are configured to perform the steps of separating a plurality of individual heart vibration events from the composite vibration objects by using at least one among bin-wise clustering and permutation alignment, or non-negative matrix factorization, or deep belief networks and tagging the plurality of individual heart vibration events using at least one among principal component analysis, Gabor filtering, generalized cross correlation, phase transform, smoothed coherent transformation, Roth correlation or band filtering.

7.Description:
(en)CROSS REFERENCE TO RELATED APPLICATIONS
This application claims the priority benefit of Provisional Application Nos. 62274765, 62274761, 62274763, 62274766, and 62274770, each of which were filed on Jan. 4, 2016, the entire disclosure each are incorporated herein by reference.
FIELD
The embodiments herein relate generally to cardiopulmonary health monitoring and more particularly to analysis software combined with transducers to capture multi-channel vibration signals along with an electrocardiogram signal for the measurement of heart functions.
BACKGROUND
Heart disease is the leading cause of death accounting for more than one-third (33.6%) of all U.S. deaths. Overall cardiac health can be significantly improved by proper triage. Low invasive and non-invasive ultrasound techniques (e.g. echocardiogram) are standard procedures, but the requirement of expensive devices and skilled operators limit their applicability. The following are the various types of heart disease that can be diagnosed and treated using the separated signal, namely, Coronary artery disease, Heart murmurs and valve abnormalities, Heart failure, Heart rhythm abnormalities (arrhythmias), Vascular disease, congenital heart disease, Cardiac resynchronization and Risk factor modification. A physician can work with patients to perform a comprehensive evaluation and design a personalized plan of care aimed at keeping them healthy.
The cardio pulmonary system which consists of the respiratory components, snoring components, and cardiac components, creates vibrations during each cardiac cycle. The vibrations are the result of the lung sounds, heart sounds, acceleration and deceleration of blood due to abrupt mechanical opening and closing of the heart valves during the cardiac cycle.
SUMMARY
The exemplary embodiments herein provide a method and system based on a technique to identify the separated cardiopulmonary signals and to extract information contained in vibration objects. In some embodiments, machine learning, auditory scene analysis, or spare coding are used as approaches to the source separation problem. Data is obtained using a tri-axial accelerometer or multiple tri-axial accelerometers placed on different points of a torso. Further note that the techniques and methods herein are not limited to acoustic, electrical or vibrational data as might be used in some stethoscopes, but can also be applied to other forms of monitoring such as echo imaging or sonograms, magnetic resonance imaging (MRI), computed tomography (CT) scanning, positron emission tomography (PET) scanning, and monitoring using various forms of catheterization. The techniques and methods herein are primarily applicable to monitoring of heart valve events, but can be alternatively applied to other types of involuntary biological signaling emanating from the brain, intrauterine, pre-natal contractions, or elsewhere within both humans and other species.
Examples of cardiac vibration objects are the first sound, the second sound, the third sound, the fourth sound, ejection sounds, opening sounds, murmurs, heart wall motions, coronary artery sounds, and valve sounds of the Mitral valve opening and closing, Aortic valve opening and closing, Pulmonary valve opening and closing, Tricuspid valve opening and closing. Examples of the pulmonary vibration objects are the respiratory lung sounds, breathing sounds, tracheobronchial sounds, vesicular sounds, Broncho vesicular sounds, snoring sounds. A portion of the energy produced by these vibrations lies in the infra-sound range, which falls in the inaudible and low sensitivity human hearing range. For example, the vibration objects from the Mitral, Tricuspid, Aortic, and Pulmonary valve openings fall in a lower range of vibrations such as 0 to 60 Hertz, whereas vibration objects from the Mitral, Tricuspid, Aortic, and Pulmonary valve closings fall in a higher range of vibrations such as 50 to 150 Hertz. A portion of the energy produced by these vibrations falls in the audible hearing range. Accelerometer transducers placed on the chest capture these vibrations from both these ranges.
Source identification analysis in accordance with the methods described herein identify individual vibration objects described above from the source separated vibration signals. The individual vibration signals are identified to be from the mitral valve, aortic valve, tricuspid valve, the pulmonary valve, coronary artery, murmurs, third sound, fourth sound, respiratory sound, breathing, and snoring during individual heart beats. The identified signals are marked to indicate their start with respect to the start of the electrocardiogram or EKG.



DESCRIPTION OF THE DRAWINGS
FIG. 1A illustrates a system for the extraction, identification, marking and display of the heart valve signals in accordance with one embodiment;
FIGS. 1B and 1C illustrate cardio pulmonary signal capture at the chest in accordance with various embodiments;
FIG. 2 is a flowchart of a method practiced by the system in accordance with one embodiment;
FIG. 3 illustrates multichannel signals captured from the sensor array on the chest shown in accordance with one embodiment;
FIG. 4 illustrates a cardiac cycle in relation with Electrocardiogram, acoustic and accelerometer sensors of the system in accordance with one embodiment;
FIG. 5 illustrates a schematic for the source identification in accordance with one embodiment;
FIGS. 6A, 6B, 6C, 6D, 6E, 6F and 6G illustrate the identification of various sources from separated sources in accordance with one embodiment;
FIGS. 7A, 7B, and 7C illustrate the source identification of various heart valve signals across the different heart beats in accordance with one embodiment;
FIGS. 8A and 8B illustrate the source identification of various heart valve signals with reference to an electrocardiogram in accordance with one embodiment;
FIG. 9 illustrates the pseudo code for the classification algorithm for source identification in accordance with one embodiment.



DETAILED DESCRIPTION
The exemplary embodiments may be further understood with reference to the following description and the appended drawings, wherein like elements are referred to with the same reference numerals. The exemplary embodiments describe a system and method of source identification of the individual cardiopulmonary signals from the composite vibrations captured on the chest wall. Specifically, psychoacoustics are considered in identifying the cardiac vibration signals captured through the transducers. The system, the psychoacoustics, and a related method will be discussed in further detail below.
The embodiments can include different source identification techniques specifically used for tagging individual cardiopulmonary signals for application in a non-linear time variant system, such as Principal component analysis, Gabor filtering, Generalized Cross Correlation (GCC), Phase transform (PHAT), ROTH, SCOT and Band Filtering. The techniques herein can use one or more among 1) Spectral information, 2) Relations among channels, and 3) Relations among events in the form of relative times of occurrence.
The exemplary embodiments provide a novel approach for small, portable, robust, fast and configurable source separation based software with transducer hardware 103 , 203 . The use of the vibration signal pattern and novel psychoacoustics help bypass conventional issues faced by linear time invariant systems.
The following are the various types of heart disease that can be diagnosed and treated using the identifies signals, namely, Coronary artery disease, Heart murmurs and valve abnormalities, Heart failure, Heart rhythm abnormalities (arrhythmias), Vascular disease, congenital heart disease, and Risk factor modification. A physician can work with patients to perform a comprehensive evaluation and design a personalized plan of care aimed at keeping them healthy.
The exemplary embodiments of the system and method proposed here are shown in FIGS. 1A, 1B, and 1C . System 100 shown in FIG. 1 is an embedded platform which can be any smart processing platform with digital signal processing capabilities, application processor, data storage, display, input modality like touch-screen or keypad, microphones, speaker, Bluetooth, and connection to the internet via WAN, Wi-Fi, Ethernet or USB. This embodies custom embedded hardware, smartphone, iPad-like and iPod-like devices. Area 101 in FIGS. 1A and 1B is the auditory scene at the chest locations. Array 102 in FIGS. 1A and 1B is a transducer array used to capture the heart signal. In some embodiments, the transducer array includes a pad that includes a vibration sensor such as a vibration sensor 102 b and an electrode 102 a for an ECG sensor. In some embodiments, the transducer array can include a single pad, two pads as shown in FIG. 1B or more than two pads as shown in FIG. 1C . In the particular embodiment of FIG. 1C , a transducer array 110 includes three pads ( 102 ) where each pad includes the vibration sensor 102 b and the ECG electronic 102 a . Other embodiments can include three or more pads where each pad would have at least a vibration sensor and optionally an electrode for the ECG sensor. Hardware 103 in FIGS. 1A-C is the wearable microprocessor hardware with digital signal processing capabilities, application processor, analog-to-digital frontend, data storage, input modality like buttons, and wireless connection via Bluetooth, Bluetooth low energy, near field communication transceiver, Wi-Fi, Ethernet or USB. Processor 112 shown in FIG. 1C comprises of the signal processing module on the wearable device that captures synchronized sensor data from the transducer array 102 . The processor 112 is configured to save the synchronized sensor data to memory and communicate it with the system 100 for data transfer. 105 in FIG. 1A is the module that calculates vital sign from the input sensor stream coming from hardware 103 for the Heart rate, breathing rate, EKG signal, skin temperature, and associated vitals. The hardware 103 can optionally encrypt the raw sensor data for transmission to the cloud computing module 106 . It can also communicate with a dashboard on a computing device 106 for data exchange, login, alerts, notifications, display of processed data. Computing device 106 in FIG. 1A can be the cloud module that processes the individual streams for eventual source separation, identification and marking of the heart valve signals. The system 100 allows a user to visually see the individual streams and information from the heart valves. In some embodiments, the system 100 could include a connected display or any other modality of display or presentation device. In some embodiments, the system 100 allows a user to visually see the individual streams and information from the heart valves and in some embodiments the system could present streams or information on a connected display or any other modality of display. The transducer array 102 can include multiple sensor transducers that capture the composite signal that includes the electrocardiogram signals, heart sounds, lung sounds and snoring sounds for example. The module 103 can be in the form of wearable hardware that synchronously collects the signals across the transducers and is responsible for the analog to digital conversion, storage and transmission to a portable unit 104 . Note that the embodiments herein are not limited to processing the individual streams for source separation, identification and marking of the heart valve signals at the cloud computing module 106 only. Given sufficient processing power, the aforementioned processing can occur at the microprocessor hardware module 103 , at the module 105 , or at the cloud-computing module 106 , or such processing can be distributed among such modules 103 , 105 , or 106 .
The exemplary embodiments of the system and method proposed here for the source identification of the cardiopulmonary signals 200 are shown in FIG. 2 . Block 201 indicates the separation of sources from the composite signals. Block 202 represents the phase estimation between the separated sources at each of the sensor position. Block 203 represents calculating the time stamps of individual sources at each heartbeat with respect to the synchronized EKG signal and the other sensor or sensors. Block 204 represents the source identification module responsible for tagging each of the separated source in individual heart beats to be one of the heart valve event, namely Mitral valve closing and opening, Tricuspid valve closing and opening, Aortic valve opening and closing, and the Pulmonic valve opening and closing. Block 205 represents the time marking module to estimate the time of occurrence of the above mentioned valve events with respect to the start of the EKG signal.
The exemplary embodiments of the system and method proposed here for the source identification of the cardiopulmonary signals from the composite signal 300 are shown in FIG. 3 . Area(s) 101 in FIG. 1B indicate the locations at which the composite heart signal can be captured. A vibration signal 302 as charted on the first line in FIG. 3 represents a signal captured at the aortic auscultation location. A vibration signal 303 shows the vibration signal captured at the pulmonic auscultation location. A vibration signal 304 shows the vibration signal captured at the tricuspid auscultation location. A vibration signal 305 represents a vibration signal captured at the mitral auscultation location. The last or bottom line in FIG. 3 represents an electrocardiogram signal 306 captured. In some embodiments, note that the number of sensors used (such as in the sensor array 102 of FIG. 1 ), are less than the number of vibration sources. For example, 3 sensors can be used to ultimately extract signals for 4 (or more) vibration sources; or 2 sensors can be used to ultimately extract signals for 3 or 4 (or more) vibration sources; or 1 sensor can be used to ultimately extract signals for 2, or 3, or 4 (or more) vibration sources.
The exemplary embodiments of the system and method proposed here draw inspirations from biology with respect to the cardiac cycle in-relation with electrocardiogram and accelerometer transducer captured cardiac signal. A timeline chart 400 in FIG. 4 shows a cardiac cycle. Lines or signals 401 a , 401 b , and 401 c represent or indicate the pressure changes during a cardiac cycle for aortic pressure ( 401 a ), atrial pressure ( 401 b ) and ventricular pressure ( 401 c ) measured in measured in millimeters of mercury (mmHg). Line or signal 402 represents or indicates the volume changes during a cardiac cycle in milliliters (ml). Line or signal 403 represents or indicates the electrical changes during a cardiac cycle captured by an electrocardiogram. Line or signal 404 represents or indicates the acoustic changes during a cardiac cycle captured by an acoustic sensor such as a phonocardiogram or PCG. S1 represents the first heart sound or the “lub” sound and the S2 represents the second heart sound or “dub” sound. Line or signal 405 represents or indicates the vibration changes during a cardiac cycle captured by an accelerometer transducer at the location of the device. Pattern 406 in FIG. 4 indicates the different valve opening and closing seen in line or signal 405 as captured by the accelerometer sensor or sensors. More specifically, a closer inspection of the pattern 406 reveals the closing of the mitral valve (M1) and tricuspid valve (T1) during the S1 or first heart sound and the closing of the aortic valve (A2) and pulmonary valve (P2). The bottom half of FIG. 4 , goes on to further show a representation of the anatomy of the human heart relevant for the generation of the sounds and a corresponding graph representing the sounds belonging to coronary artery, murmurs, first sound, second sound, third sound, fourth sound, ejection sounds, opening sounds, respiratory sound, breathing, and snoring during individual heart beats, with respect to the electrocardiogram signal.
The exemplary embodiments of the system and method proposed here provide a source identification analysis algorithm that allows for the tagging of the vibrations from the cardiopulmonary system. The sensors can be modeled as:
x 1 ( t )= s 1 ( t )+ n 1 ( t )
x 2 ( t )=α s 1 ( t+D )+ n 2 ( t ),

Where x are the sensor signals and s and n are the sources and the noise respectively (assumed to be uncorrelated). A system 500 can include filtering or weighting (using transforms 501 (H 1 ) and 502 (H 2 )) of the respective input sensor signals X1 and X2. Cross correlating the filtered sensor signals (using blocks 504 , 505 , and 506 ) and finding the peak using peak detector 507 will result in the estimate of the delay D as can be seen in the block diagram of FIG. 5 . The cross correlation between x1(t) and x2(t) is related to the cross power spectral density function by the well-known Fourier transform relationship. The process takes the filters 501 and 502 into account, which can be considered a weighting function. This weighting function and the delay 503 is used to compensate for the smearing effect of the Fourier Transform. The weighting function can be of different forms depending on the conditions of the signal, predominant between the weighting functions are: GCC-ROTH, GCC-SCOT, GCC-PHAT, PASSBAND FILTERING (The weighting function is an ideal pass band filter that eliminates all frequencies above and bellow the expected range).







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The chart 600 of FIG. 6A , shows the solid lines with manual marking of the M1 and T1 (the first and second vertical lines, respectively) showing peaks in the composite Mitral auscultation site. The adjacent broken lines correspond to the maximum point of the signals obtained by source separation. A feature similar to GCC-PHAT can be obtained from the aligned cluster centers shown in chart 610 of FIG. 6B . Each signal is represented with a different color or shading in grayscale. The peak of each one of the separate colored curves (represented by the vertical lines in FIG. 6B ) corresponds to the delay estimate of arrival between two sensors for each one of the sources. Permutation alignment is important for source separation. It can also help with the identification problem. Since the cluster centers are related to source locations. If we compute the phase differences between two sensors, we can obtain a histogram of the phase differences for the full recording as depicted for a mixture of 4 sources as shown in FIG. 620 of FIG. 6C . The relative times of arrival are clearly different for the M1 and T1 sounds as shown in the charts 630 of FIG. 6D and chart 640 of FIG. 6E . This is a clear indicator of the different location of the sources. Analysis of the propagation of sources can be based on signal markings. First notice the periodic pattern, which is more evident for the relative TOA of the T1 sound. This pattern correlates with the breathing intervals of the subject as shown in charts 650 of FIG. 6F and 660 of FIG. 6G . It suggests that propagation speed is directly affected by breathing. Propagation is not time invariant while working with cardiopulmonary signals.
One of the proposed embodiments of the source identification is using Gabor analysis on the source separated signals. After the signals have been separated the next step is to identify them, this process is usually known as source tagging. The need to do source tagging arises from the permutation problem caused by the source separation algorithm. The separated sources obtained by the source separation are similar to the original sources up to a scale and a permutation factor.
In another embodiment, the source can be identified by manually tagging them against a synchronized EKG signal and taking advantage of the timings relative to the QRS wave. This way, however is usually slow and time consuming and an automatic tagging algorithm is thus preferable. Since the different heart sounds comes from different locations in the heart, it is expected that each source will have a unique phase relation between the sensors that are located at fixed points during the data gathering phase. This phase relation can be made evident by using signal representation with a dictionary suited for highlighting frequency and phase relations with a greedy algorithm such as the matching pursuit algorithm. One of the proposed algorithms herein performs a Gabor analysis in the source separated signals using a finite Gabor Dictionary of fixed frequencies and with variable phase delay. The system finds the delay that minimizes the reconstruction error for each source and creates a group of features that is incorporated into a decision-making and classification algorithm. The classification algorithm will combine the features extracted by the Gabor analysis with other features that comes from the Principal Component Algorithm (PCA) and cross correlation analysis that uses a set of manually tagged patient's tracks as training. The system is composed of three modules or stages: A suitable Gabor dictionary is created that can serve as the basis representation of the signals. An optimization algorithm aims to find the delay that will minimize the reconstruction error in a giving delay range. Collection and organization of all the features extracted from the second stage. For each one of the stages different techniques were tried in order to achieve the best results. The Gabor dictionary selected has the function:





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In the equation 8 is the delay, a is the Gaussian decay of the atom and f is the frequency of operation of the Gabor. The system initially searches for the Gabor atoms that best represent the signal by doing a sweep in frequency and selecting those Gabor atoms that produced a larger weight matrix after the matching pursuit. The system was subsequently changed to a group of fixed frequency Gabor atoms exploiting the fact that for the M1, T1, A2 and P2 sounds vibration reside mainly in the range of 50-300 Hz and the Aortic opening vibration resides mainly in the low frequency range (from 0-50 Hz). The fixed Gabor approach sacrifices a little bit of reconstruction error in favor of a faster computation and easier analysis since the frequencies were fixed. The first optimization technique used was the, Gradient Descent algorithm: A first-order optimization algorithm that looks for the minimum of a given function by taking steps proportional to the negative of the gradient. The system is very dependent on the initial guess and number of minima of the system. A second approach was using the minimize function, which uses the Polack-Ribiere algorithm of conjugate gradients to compute search directions, and a line search using quadratic and cubic polynomial approximations. The system is also dependent on the initial value of the search but is more efficient in choosing the parameters to find the minimum. Finally, a brute force approach was implemented that sweeps the signal with different delays and then selects the one that resulted in a minimum function. In order to improve speed on this algorithm, a broad search is done first and then a refined search is done around the minimum point found by the broad search.
The exemplary embodiments of the system and method proposed here provide a source identification algorithm for the vibrations from the cardiopulmonary system. Referring to the chart 700 of FIG. 7A , in order to find the time stamps for events such as Mitral closing & opening, Tricuspid closing & opening, Aortic opening & closing, Pulmonic opening and closing, we look at all the individual source separated signals 701 to 706 of the composite signal 707 and first try to find the location of max peak in the source separated (SS) signal for each source and then find delay between two channels. The chart 710 of FIG. 7B shows the frequency spectrum of the source separated signals. Cross correlated vibrations in aortic and pulmonic channels for each interval for each source are calculated to find a consistent delay between two channels as represented by chart 720 of FIG. 7C . Given the start of QRS complex (indicative of ventricular depolarization triggered by contraction of the ventricles) and end point of each vibration, the vibration(s) within this interval is cross correlated with all vibrations in each source. This is done for both aortic and pulmonic channels. At the end, Principle Component Analysis (PCA) was applied to find the timing information and delay between two channels. PCA uses (separated source) SS signal from each source in each channel to find the template which represents a majority of the vibrations within that source. The template is then cross correlated with the whole source and maximum of PCA signal in each interval is found and compared with the start of QRS. In another implementation, in the second attempt, SS signals from two channels but the same source are fed into the PCA to find the template. Then the aortic template is used for both channels' cross correlation to identify the different vibrations into valve events, breathing sounds, and vibrations of the heart walls. To accurately estimate M1, T1, A2, P2, AO, PO from the frequency signal captured by digital accelerometer on the wearable, several sub-frequencies are considered. Some are noted here for example: 0-30 Hz, 0-60 Hz, 30-150 hz, 30-250 Hz.
The exemplary embodiments of the system and method proposed here provide a source marking algorithm for the vibrations from the cardiopulmonary system. Next step is to use the PCA to determine which source is associated with which event (Mitral closing & opening, Tricuspid closing & opening, Aortic opening & closing, Pulmonic opening and closing). The architecture for automatic source tagging and timing of valvular events is further described. One way to identify which events are relevant to a source is by manually tagging the sources against the synchronized EKG signal and taking advantage of the timings relative to the QRS wave (identification of the S1 and S2 sounds using the EKG signal as the reference has been widely researched in studies. Another approach is an automatic tagging algorithm. The tagging is composed of a classifier preceded by a feature extraction algorithm. For the timing, the system can exploit the computations of one of the feature extraction algorithms to obtain an energy contour from which the time location of a given event can be inferred. Because the embodiments herein build upon having the ability to capture the signal at different locations simultaneously, to the system exploits the relations among channels to extract additional information about the sources. Likewise, since some source separation algorithms where channels relations are associated with location, the system can leverage on the intrinsic relations among the channels to extract relevant information that helps distinguish among the events. The system herein hypothesizes that phase information between channels is relevant for distinguishing among cardiac events since valves are located at different positions within the heart. Perhaps, one of the most distinctive features of the cardiac events is their relative order of occurrence, which repeats periodically with each heartbeat. Time information extracted from the set of sources can be utilized to localize the occurrence of each source signal within the heart cycle. Therefore, the features proposed here are conceived to provide three aspects: 1) Spectral information, 2) Relations among channels, and 3) Relations among events in the form of relative times of occurrence. In some embodiments, the system includes a feature extraction algorithm based on multichannel Gabor basis decomposition using matching pursuit, and a second algorithm that uses eigen-decomposition of a covariance matrix extracted from the source signals to obtain a cross correlation function between channels. For the Gabor basis, consider the reconstruction of a 2-channel signal x (t). In this approach to the system can extract a phase relation between the sensors which are located at standard aortic and pulmonic auscultation positions. As mentioned above, the system is interested on features that reflect spectral content as well as the channel interrelations. In the two channel case, each basis function is a pair of Gabor functions with equal frequency f and envelope width h and a phase difference θ. The tunable parameter θ can be swept across a range of values for which the matching pursuit reconstruction error can be obtained. The behavior of the reconstruction error for a source signal over different values of θ provides information about the phase different between the two channels for the particular source signal. In particular, the value of θ that attains the minimum error within the defined range of values is taken as the optimal channel delay, and the time average of the activation coefficients obtained by matching pursuit at the optimal θ provide and spectral characterization of the source. For the feature Extraction by Self-Similarity Template Based-on PCA, we define a self-similarity using the concept of eigenfilter. Let yi the a two-channel source signal defined in (1), and let Yi its discretized version of size Nx2. For filter of length Nw, we compute the tap-delay matrices Δ(j). The joint-channel eigenfilter correspond to the leading right-singular vector u of the tap-delay centered matrix Δi obtained by removing the mean of each column of Δi. The eigenfilter u can be split into the first and second channel components. The first channel component u(1) corresponds to the first Nw entries of u, and u(2) to the remaining entries going from index Nw+1 to last index 2Nw. The proposed feature is the cross-correlation function between u(1) and u(2). Note that this function contains the main frequency components of the source signals expressed in the time domain and the peak provides information about the channel relations. For different delays between the two channels the cross correlation peaks at different locations, accordingly. To extract relative time information contains cues for the classification of the cardiac events, we adopt a simple approach that uses the energy contours. The process consists of three basic steps: Compute energy contours for all source signals, Compute timings for each source signal, Compute features of source i by averaging the sum of the remainder sources centered at the timings. The energy of each source signal Yi is calculated using the leading, quadrature pair of right singular vectors of their corresponding tap-delay matrix. For classification, a set of training exemplars (sources) have been manually tagged with the respective events. Each source signal is then represented by the feature vectors described above. Tests have been carried out on pooled covariance linear discriminant analysis, quadratic discriminant analysis, k-nearest neighbours and support vector machines both linear kernel and Gaussian kernel. The features described in the previous sections can be used in different manners for tagging, for example: Tagging using only Gabor-based features, Tagging using only self-similarity features, Tagging using only time-based features, Tagging using all Gabor-based, self-similarity, and time-based features. There are even sub cases of the situations considered above, For instance, the combination of Gabor and self-similarity features can be done on a single classifier or using ensembles. The first validation examples only address the cases in the bullet items. In addition to the classifier, preprocessing of the feature vectors is also performed. A centering vector x cent is calculated using the mean over the training set x train . This step is followed by a linear transformation based on PCA to produce a much smaller dimension feature vector. The preprocessing is summarized in Algorithm 1, in FIG. 9 . The choice of p is typically driven by the training data itself. The singular values of x train can be used to decide the dimensionality of the transformed feature vectors after.
The exemplary embodiments of the system and method proposed here provide a source marking algorithm that allows from the explanation earlier for the marking of the Mitral valve closing (MC), Mitral valve opening (MO), Aortic valve opening (AO), Aortic valve closing (AC), Tricuspid valve closing (TC), Tricuspid valve opening (TO), Pulmonary valve closing (PC) and Pulmonary valve opening (PO) signals. The extracted individual valve vibration objects are aligned into a signal for each of the four valves across multiple heart beats. The chart 800 in FIG. 8A shows the source separation of heart valve opening and closing signals. Line 801 indicates length or duration of the vibration signal for the Mitral valve closing (M1). Line 802 indicates length or duration of the vibration signal for the Tricuspid valve closing (T1). Line 803 indicates the length or duration of the vibration signal for the Aortic valve closing (A2). Line 804 indicates the length or duration the vibration signal for the Pulmonic valve closing (P2). Line 805 indicates the composite vibration signal captured by a particular transducer. Signal 806 indicates the EKG signal captured by the system. Referring to chart 810 of FIG. 8B , the line 807 indicates the length or duration of the vibration of the Aortic valve opening (AO). Line 808 indicates the length or duration of the vibration of the Pulmonic valve opening (PO). Further note that the lines or signals 809 in FIG. 8A or 811 in FIG. 8B are actually several separated superimposed signals representing the vibration signals from separate sources coming from the mitral valve, tricuspid valve, aortic valve, and pulmonary valve (using less than 4 vibration sensors to extract such separated signals in some embodiments).
The exemplary embodiments of the system and method proposed here provide a source marking algorithm for the vibrations from the cardiopulmonary system, using information about the time of occurrence of the event. Automated finding of M1, T1, A2 & P2: After going through the timing plots of different patients, it was decided to first determine A2 and P2 timing and their corresponding sources and then find M1 and T1 from the rest of sources. Automatic A2 & P2 finding: 1) The number of zero-crossings is found for each SS source. Then the noisy source that is associated with the maximum zero-crossing is discarded. 2) Automated calculation of QRS onset points for all heartbeats in each source. 3) Automated calculation of beginning and ending points of all vibrations in each source. 4) Applying PCA approach on all sources that outputs the time difference between QRS onset and peak of PCA signal (timing vector) as well as delay between aortic and pulmonic channels. 5) Based on probability density estimation of each source's timing vector, strong peak(s) that correspond(s) to timing(s) more than 300 ms are accepted as A2 and P2 candidates and those sources that don't satisfy this condition are discarded. Also new candidate timings are sorted in ascending order. 6) Variation, length of samples more than 300 ms (L) and some other measurements are calculated from accepted timing vectors of accepted sources and are stored in a structure for further analysis. 7) If there is only one strong peak in probability density estimation plot whose timing is more than 300 ms, then a variable called “ind_both” is set to 1. If there are two or more strong peaks whose timing is more than 300 ms, if max peak's timing is over 300 ms, then “ind_both” variable is set to 1 but if max peak's timing is lower than 300 ms but another strong peak's timing is more than 300 ms, then “ind_both” is set to 0. If there are two or more sources whose “both_indic” variables are equal to 1, then “both_indic”=1.8) In case of having no valid timing, A2 and P2 are set to zero. 9) In case of having one valid timing, if L>6 then it is marked as A2 and its corresponding source number is also saved. P2 is set to zero. 10) In case of having two valid timings:
if both_indic=1,
if L1>=6 samples, then A2 is set to first timing in sorted timing vector Else A2=0
if A2=0 and L2>=6, then A2 is set to second timing in sorted timing vector Else P2=0
if A2˜=0 and L2>=6, then P2 is set to second timing in sorted timing vector Else P2=0
if both_indic˜=1,
if L1>=4 samples & V1<=800 then A2 is set to first timing in sorted timing vector Else A2=0
if A2=0 and L2>=4 & V1<=800, then A2 is set to second timing in sorted timing vector Else P2=0
if A2˜=0 and L2>=6 & V2<=800 then P2 is set to second timing in sorted timing vector Else P2=0
11) In case of having more than two valid timings,
if both_indic=1

for loop: if L1>=6 & V1<=800, then A2 is set to current timing in sorted timing vector; break Else A2=0; End of for loop

if A2=0, then P2=0;
if A2˜=0 & last timing in the sorted timing vector
if L>=6 & V1<=800, then P2 is set to last timing in the sorted timing vector
if A1˜=0 & there are two timings after A1 timing,
if V2<V1, then P2 is set to second timing after A2.
Else P2 is set to first timing after A2.
if A1˜=0 & there are more than two timings after A1 timing,
if L1>=6 & V1<=800, then P2 is set.
if both_indic ˜=1,

for loop: if L1>=8 samples & V1<=1000 then A2 is set to current timing in sorted timing vector Else A2=0; End of for loop.

if A2=0, then P2=0;
if A2˜=0
for loop: if L>=8 & V1<=1000,
then P2 is set to current timing in the sorted timing vector; break
Else P=0;
End of for loop

At the end, A2 and P2 timings and their corresponding source numbers can be saved in an excel spreadsheet along with the patient name.

Automatic M1 & T1 finding:
1) The same previous steps from 1 to 7 are implemented in this phase 8) In case of having no valid timing, M1 and T1 are set to zero. 9) In case of having one valid timing, if L>6 and timing is less than 120 ms, then it is marked as M1 and its corresponding source number is also saved.
P2 is set to zero.
10) In case of having two valid timings,

if both_indic=1,


if L1>=6, then M1 is set to first timing in the sorted timing vector Else M1=0
if M1=0 & L2>=6, then M1 is set to second timing in the sorted timing vector and T1=0
if M1˜=0 & L2>=6 and second timing is less than double the first timing
then T1=second timing in the sorted timing vector Else T1=0
if both_indic ˜=1

L1>=4 & V1<=800, then M1 is set to first timing in the sorted timing vector Else M1=0

If M1=0 & L2>=4 & V2<=800, then M1=second timing in the sorted timing vector Else T1=0

If M1˜=0 & L2>=4 & V2<=800, then T1=second timing in the sorted timing vector Else T1=0

11) In case of having more than two valid timings,
if both_indic ˜=1,

for loop: if L>=8 & V<=1000, then M1 is set to current timing in the sorted timing vector Else M1=0; break; end of for loop

if M1=0, then T1=0;
for loop: if M1˜=0 & L>=6 & V<=1000, then T1 is set to current timing in the sorted timing vector; break, end of for loop

if M1˜=0 & L2>=6 and second timing is less than double the first timing

then T1=second timing in the sorted timing vector Else T1=0
if both_indic=1

low_var=Find sources with variations less than 15.

From these low_var sources, find those whose length are more than 9 samples (len_low_var).

If length of low_var>=2 & len_low_var>=2

M1 is set to the timing corresponding to the lowest variation
T1 is set to the timing corresponding to the second lowest variation
Else
for loop: If L>=5 & V<800, then M1 is set to the current timing in the sorted timing vector Else M1=0

If M1=0, then T1=0

If M1˜=0
for loop:

If there is only one timing after M1 & L>=5 & V<=800, then T1=current timing in the sorted timing vector; break

If there are two timings after M1,

if V1<300, then T1=first timing after M1 in the sorted timing vector; break

Else if V2<V1, then T1=second timing after M1 in the sorted timing vector; break

Else, T1=first timing after M1 in the sorted timing vector; break

If there are more than two timings after M1,

if V1<5, then T1=current timing after M1 in the sorted timing vector; break

Else if V2<V1, then T1=current timing after M1 in the sorted timing vector; break

Else, T1=current timing after M1 in the sorted timing vector; break

Automated Finding of Aortic Opening (AO):




Track 1 with 2 channels 8 by 2 by N
M1, T1 timings and EKG times
Remove the noisy source using max zero-crossing. Data down to 7 by 2 by N
Sub-source markings including start and end of vibrations, envelope and max peak
Save to non-zero-source vector for both channels
Remove missing start and end points and too large ones by comparing between two channels as well as considering each source
Calculate PCA for two channels and all vibrations. Outputs are delay_vec, PCA signal and delay between two channels




Begin AO Detection




Select above data for AO channel
Remove zero delay values
Remove delays between two channels that are greater than 50 ms and count <4
Remove positive delays between two channels and count <5
Calculate ksdensity
Find max peak as well as other peaks that their distance together is more than 5 samples
Keep negative delays if delay is less than 150 ms
Calculate ksdensity of event_vec
Find more max peak and peaks >half of max peak
Find peaks >10 ms and with 0.05 of max probability value
Save all the candidate sources with negative delay btw two channels
Select peaks that fall in range of M1 and M1+95 ms (if M1=T1=0, in range 20 to 150 ms)
Sort all timings and save
Find AO




First Case:
If there is only on candidate source:
1) If delay variation <500 ms but source variation >1000 ms, 2) remove timings >200 ms, 3) If length of samples >4, remove samples whose timings relative to median are >60 ms 4) If empty timing vector:

I. then get the original timing vector, calculate kdensity and find max and other peaks whose probability is more than half of max's probability

II. If time diff btw AO candidate and max peak's timing <10 ms,


set max's peak timing as ref time, remove samples with diff more than 50 ms relative to ref time and update the AO candidate and its variation.
III. If time diff btw AO candidate and max peak's timing >10 ms, find a qualified peak and its corresponding timings that is closest to AO candidate, set that peak's timing as ref time, remove samples with diff more than 50 ms relative to ref time and update the AO candidate and its variation.
5) If non-empty timing vector: I. calculate kdensity and find max and other peaks whose probability is more than half of max's probability. Find the one that is closest to candidate, set is as ref time, remove samples with diff more than 50 ms relative to ref time and update the AO candidate and its variation. 6) Check if var_delay <500 && var_SRC<900 && AO_candidate <=M1+96

Second Case:
If there are two candidate sources:
The following steps from the first case are implemented:
1, 2, 3 with 50 ms threshold, and 6 After founding all valid AO timings from all candidate sources, the closet AO candidate to “M1+30 ms” is selected.

Third Case:
If there are more than two candidate sources:
The following steps from the first case are implemented:
1, 2, 3 with 90 ms threshold, 4 with 90 ms threshold, and 6 After founding all valid AO timings from all candidate sources, the closet AO candidate to “M1+30 ms” is selected.

To improve the results in the exemplary embodiment the system can perform AO detection using a new PCA approach such as: Mean removal, Min-Max removal, or Different window length for correlation. Detailed analysis on delay calculation from cross-correlation template. Source tagging using PCA features and different classifiers on both tagged and untagged tracks for both High Pass and Low Pass. Defined two timing calculation algorithm 1) with threshold 2) without threshold and got the timing distributions for different test cases. To improve AO detection using new PCA algorithm, a few different tests were done such as: Removed the mean of each source's template (X matrix) before calculating covariance matrix. For some sources, cross-correlation template shows zero delay at a discontinued point although there is a delay between peaks of templates. Removed min-max of eigenvector before performing cross-correlation showing a continuous behavior with a delay. Then ran different SSs on the same patient to see if these observations are consistent. There are different scenarios when calculating the delay such as: Maximum peak is positive or negative, The Delay of Max peak is positive or negative. The approach herein is to select positive maximum peak with the positive delay. The goal in some embodiments is to tag or label the sources with an event (HP: M1, T1, A2, P2, LP: AO). The cross-correlation signals generated by new PCA algorithm are used as features to train the classifiers and then tag sources of a test signal. A similar approach can be extended for the identification of other cardiopulmonary signals, like coronary artery sounds, snoring, murmurs, respiration sounds in one embodiment.
The exemplary embodiments develop a novel method of source identification, which in one embodiment using the Pulmonary and Aortic, and in addition possibly the Tricuspid and Mitral auscultation locations, lends the system capable of identifying individual cardiopulmonary events from the vibrations.
The exemplary embodiment of system and method described is the development on an embedded hardware system, the main elements required to capture body sounds are the sensor unit that captures the body sounds, digitization, and digital processing of the body sounds for noise reduction, filtering and amplification. Of course, more complicated embodiments using the techniques described herein can use visual sensors, endoscopy cameras, ultrasound sensors, MRI, CT, PET, EEG and other scanning methods alone or in combination such that the monitoring techniques enable improvement in terms of source separation or identification, and/or marking of events such as heart valve openings, brain spikes, contractions, or even peristaltic movements or vibrations. Although the focus of the embodiments herein are for non-invasive applications, the techniques are not limited to such non-invasive monitoring. The techniques ultimately enable diagnosticians to better identify or associate or correlate detected vibrations or signaling with specific biological events (such as heart valve openings and closings, brain spikes, fetal signals, or pre-natal contractions.)
It will be apparent to those skilled in the art that various modifications may be made in the present invention without departing from the spirit or scope of the invention. Thus, it is intended that the present invention cover the modifications and variations of this invention provided they come within the scope of the method and system described and their equivalents.