Transmission mode photon density wave system and method

Present embodiments are directed to a monitor system, such as a pulse oximetry system. The system may include a detection feature, an emission feature capable of emitting light into tissue and arranged relative to the detection feature such that the detection feature is capable of detecting the light from the tissue after passing generally through a portion of the tissue, a modulator capable of modulating the light to generate photon density waves at a modulation frequency generally in a range of 50 MHz to 3 GHz, a detector communicatively coupled with the detection feature, wherein the detector is capable of detecting characteristics of the photon density waves comprising amplitude changes and phase shifts, and a processor capable of making determinations relating to a value of a physiologic parameter of the tissue based at least in part on the detected characteristics.

BACKGROUND

The present disclosure relates generally to a photon density wave system, and, more particularly, to apparatus of a transmission mode photon density wave system.

Pulse oximetry may be defined as a non-invasive technique that facilitates monitoring of a patient's blood flow characteristics. For example, pulse oximetry may be used to measure blood oxygen saturation of hemoglobin in a patient's arterial blood and/or the patient's heart rate. Specifically, these blood flow characteristic measurements may be acquired using a non-invasive sensor that passes light through a portion of a patient's tissue and photo-electrically senses the absorption and scattering of the light through the tissue. Typical pulse oximetry technology currently utilizes two light emitting diodes (LEDs) and a single optical detector to measure pulse and oxygen saturation of a given tissue bed.

A typical signal resulting from the sensed light may be referred to as a plethysmographic waveform. Such measurements are largely based on absorption of emitted light by specific types of blood constituents. Once acquired, this measurement may be used with various algorithms to estimate a relative amount of blood constituent in the tissue. For example, such measurements may provide a ratio of oxygenated to deoxygenated hemoglobin in the volume being monitored. It should be noted that the amount of arterial blood in the tissue is generally time varying during a cardiac cycle, which is reflected in the plethysmographic waveform.

The accuracy of blood flow characteristic estimation via pulse oximetry depends on a number of factors. For example, variations in light absorption characteristics can affect accuracy depending on where the sensor is located and/or the physiology of the patient being monitored. Additionally, various types of noise and interference can create inaccuracies. For example, electrical noise, physiological noise, and other interference can contribute to inaccurate blood flow characteristic estimates. Some sources of noise are consistent, predictable, and/or minimal, while some sources of noise are erratic and cause major interruptions in the accuracy of blood flow characteristic measurements. Accordingly, it is desirable to enable more accurate and/or controlled measurement of physiologic parameters by providing a system and method that addresses inconsistencies in physiologic characteristics of patients and issues relating to noise.

DETAILED DESCRIPTION OF SPECIFIC EMBODIMENTS

Present embodiments relate to non-invasively measuring physiologic parameters corresponding to blood flow in a patient by emitting light into a patient's tissue with light emitters (e.g., lasers/LEDs) and photoelectrically detecting the light after it has passed through the patient's tissue. More specifically, present embodiments are directed to modulate the emitted light at high frequencies to generate resolvable photon density waves. Photon density waves may be described as progressively decaying waves of intensity. On a microscopic level, photons generated by a light source generally make random migrations in a scattering medium. However, the photons collectively form a photon density wave at a modulation frequency that moves away from the light source. Photon propagation is generally dictated by scattering and absorption in the medium through which the waves are moving. Like other waves, photon density waves undergo refraction, diffraction, interference, dispersion, attenuation, and so forth.

Phase changes and amplitude changes in the photon density waves after passing through a medium may be detected to facilitate measurement of changes in total scattering particles and absorber concentration, respectively, in the observed medium. Indeed, the phase of such waves may be sensitive to scattering and the amplitude of such waves may be sensitive to absorption. For example, detection of phase changes in the photon density waves generated by modulation at high frequency may correspond to total hemoglobin because the wavelength of the photon density waves may be shorter than an average absorption length of photons. Thus, detected variations in the phase may be predominantly due to the scattering coefficient and not absorption. In other words, the variation in phase may be predominantly due to the total number of scattering particles (e.g., total hemoglobin) in the observed medium and not merely a ratio of particles (e.g., oxygenated and deoxygenated hemoglobin) that absorb different colors of light. On the other hand, changes in the amplitude of the photon density waves may correspond to absorption of specific light color (e.g. red or infrared light) in the observed volume, and, thus, a ratio of different types of particles (e.g., oxygenated and deoxygenated hemoglobin) in the probed medium.

In addition to the features set forth above, it should also be noted that present embodiments may relate to emitting multiple high frequency photon density waves in coordination with one another to focus on certain tissue areas (e.g., regions rich with pulsatile signals), to facilitate identification of noise artifacts, to address patient specific tissue characteristics (e.g., skin color and low blood oxygen saturation levels), and/or to reduce noise in general. For example, multiple photon density waves may be emitted in patterns such that the waves build on one another to focus intensity at certain points throughout a tissue bed. In a specific example, a tissue bed may be swept with combinations of emission frequencies to identify areas rich with pulsatile signals. Similarly, waves may be emitted such that the waves cancel one another out in a substantially noise-free environment. Thus, detection of the waves that have not been canceled out may be indicative of the presence of noise. Additionally, relative measurements may be utilized to identify and/or correct noise. For example, certain wave features may be detected at multiple detector locations and compared to one another to identify characteristics such as venous pulsation noise.

FIG. 1illustrates a perspective view of a pulse oximetry system10in accordance with some embodiments. The system10includes a pulse oximeter or monitor12that communicatively couples to a sensor14. The monitor12may include a display15, a memory, a processor, and various monitoring and control features. The sensor14may include a sensor cable16, a connector plug18, and a sensor assembly or body20configured to attach to a patient (e.g., a patient's finger, ear, lip, or toe) in a manner that facilitates transmission of light through the patient's tissue from one side to another. The system10may be utilized to observe the blood constituents of a patient's arterial blood to facilitate estimation of the state of oxygen exchange in the patient's body by emitting waves into tissue and detecting the waves after dispersion and/or reflection by the tissue. The amount of light that passes through the tissue and other characteristics of light waves may vary in accordance with the changing amount of certain blood constituents in the tissue and the related light absorption and/or scattering. For example, as with conventional pulse oximeter systems, the system10may emit light from two or more LEDs or lasers into pulsatile tissue and then detect the transmitted light with a light detector (e.g., a photodiode or photo-detector) after the light has passed through the pulsatile tissue. Such measurements may be utilized to estimate a percentage of blood oxygen saturation in the probed volume of blood. Additionally, in accordance with present embodiments, the system10may modulate the emitted light to generate photon density waves at a high frequency such that phase shifts may be detected that correlate predominantly to scattering particles in the probed volume of blood.

As generally indicated above, the system10may generate and detect light waves to facilitate non-invasive measurement of a patient's physiological characteristics. In embodiments, the system10may generate resolvable photon density waves and make relative measurements of certain detected wave characteristics after the waves have been transmitted from one side of a medium (e.g., the tissue of a patient's finger) to the other. The wave characteristics that may be measured in accordance with present embodiments may include characteristics that relate predominantly to absorption of the emitted light in the probed medium (e.g., amplitude change) and characteristics that relate predominantly to scattering in the probed medium (e.g., phase shift). It should be noted that, as will be discussed further below, the correlation of certain wave characteristic (e.g., amplitude and phase) measurements to certain medium characteristics (e.g., quantity of scattering particles and blood oxygen saturation) may be based on high frequency modulation of the system's light sources, which generate the resolvable photon density waves.

As indicated above, the system10may be utilized to make measurements that relate predominantly to scattering in the observed volume. More specifically, the system10may be utilized to make measurements relating to a total amount of scattering particles in the observed volume based on phase shifts detected in the emitted light waves. For example, the system10may emit light that is modulated at a high frequency (e.g., 50 MHz to 3.0 GHz) to generate resolvable photon density waves, and then measure the phase shift of these high frequency waves to facilitate estimation of a total number of scattering particles in the observed medium. Similarly, as set forth above, the system10may be utilized to make measurements that relate predominantly to absorption in an observed volume. For example, the system10may detect changes in AC and DC amplitudes of the resolvable photon density waves to facilitate detection of a ratio of certain constituents in the blood (e.g., a ratio of oxygenated hemoglobin to total hemoglobin). It should be noted that the amplitude changes and phase shifts measured at a detection point may be considered relative to one or more points. For example, the amplitude and phase shifts measured at a detector may be considered relative to the associated values generated at the emitter.

FIG. 2is a basic block diagram of an embodiment of the pulse oximeter system10that is capable of transmission mode photon density wave emission and detection. The configuration and operation of the system10in transmission mode may enable deep penetration of the photon density waves into a region of interest in a patient's tissue. As inFIG. 1, the system10illustrated inFIG. 2includes the monitor12and the sensor14. InFIG. 2, the monitor12and the sensor14include features capable of cooperating to transmit photon density waves into one side of a patient's tissue and out of a generally opposite side for detection. Specifically, inFIG. 2, the monitor12is illustrated as including various functional components that facilitate transmission of photon density waves through an emitter-side fiber optic cable22to the sensor14. The sensor14is physically configured such that, when properly attached to a patient's tissue24, the photon density waves from the emitter-side fiber optic cable22pass into one side of the patient's tissue24(e.g., one side of a patient's earlobe or finger), out of the generally opposite side of the patient's tissue26, and into a detector-side fiber optic cable30coupled with the sensor14. For example, the sensor14may include a clamping mechanism that positions the emitter-side fiber optic cable22generally opposite the detector-side fiber optic cable30when straddling the patient's tissue24. In accordance with present embodiments, the sensor body20may be specifically arranged such that a light emission feature (e.g., the emitter-side fiber optic cable22, a fiber optic component connector, a laser, or an LED) is generally arranged opposite a light detection feature (e.g., the detector-side fiber optic cable30, a fiber optic component connector, a lens, a transparent layer, or a detector) such that photon density waves can be passed through the tissue26from the light emission feature on a first side of the tissue to the detection feature on a generally opposite or opposite side of the tissue26.

The functional components disposed within or included as features of the monitor12may include sensor driving circuitry32, an emitter34, a detector36, phase detection circuitry38, and digital signal processing (DSP) circuitry40. While some embodiments may include differing component arrangements (e.g., certain features may be included in the sensor14instead of the monitor12), including these functional components within the monitor12may leave the sensor14to be generally composed of fiber optics, which may make the sensor14cheap and easily disposable. Indeed, the sensor14may include a sensor body42that houses fiber optic components44and/or portions of the emitter-side fiber optic cable22and the detector-side fiber optic cable30. The fiber optic components44may include features (e.g., fiber optic curves and coupling mechanisms) that may facilitate communicative coupling with monitor12and/or arranging the emission and detection points of the emitter-side fiber optic cable22and the detector-side fiber optic cable30generally opposite one another and such that light is emitted substantially directly into the patient's tissue and detected opposite the emission point. As illustrated inFIG. 2, the emitter-side fiber optic cable22is coupled to the emitter34and the detector-side fiber optic cable30is coupled to the detector36. Further, each fiber optic cable22,30may be coupled to the sensor body42and/or the fiber optic components44within the body42to facilitate transmission of the photon density waves through the patient's tissue. In the illustrated embodiment, the fiber optic components44function to turn the emitted light into the patient's tissue26and the received light back into the detector-side fiber optic cable30. In other embodiments, mirrors, prisms, the fiber optic cable itself or the like may be utilized to guide the light in a desired direction.

In operation, the driving circuitry32may generate waves (e.g., sine waves) and provide timing control signals such that the emitter34is activated in specified intervals and/or such that certain varying amplitudes of light are emitted by the emitter34to produce desired qualities of the photon density waves. The driving circuitry32may also include features that are capable of controlling access to the phase detection circuitry38via clock signals provided to the phase detection circuitry38from the driving circuitry32. As an example of a typical process in accordance with present embodiments, the driving circuitry32may cause the emitter34to emit photon density waves at a high frequency into the emitter-side fiber optic cable22such that the photon density waves are transmitted into one side of the patient's tissue24and out of the other side of the patient's tissue26. The detector-side fiber optic cable30may receive at least a portion of the photon density waves and transmit the photon density waves to the detector36, which communicates with the phase detection circuitry38to identify phase information and so forth. This information may then be transmitted to the DSP circuitry40for analysis. It should be noted that in some embodiments, multiple emitters may be utilized in conjunction with multiple fiber optic cables or the multiple emitters may share a single fiber optic cable.

FIG. 3is a block diagram of another embodiment of the pulse oximeter system10that may be configured to implement embodiments of the present disclosure. As indicated above, the system10may include the monitor12and the sensor14. In accordance with present embodiments, the sensor14may be configured such that light from an emitter50can pass into a patient's tissue52on one side and out a generally opposite side when properly attached. Further, the sensor14may be configured such that after being transmitted from one side of the tissue52to the generally opposite side, the light may be received by a photo-detector54. The photo-detector54may then convert the received light into a photocurrent signal, which may then be provided to the monitor12. It should be noted that in some embodiments, multiple sensors14may be employed. Further, in some embodiments, one or more sensors may each include multiple emitters and/or detectors. If multiple emitters are employed, it will generally be desirable for each of the emitters to include red and infrared (IR) light sources, such as laser diodes (LD)56.

In some embodiments, in addition to the emitter50and the detector54, the sensor assembly or body20may also contain various other features in accordance with present embodiments. For example, the sensor14may include a phase detector58capable of detecting phase shifts in photon density waves observed by the detector54. While the phase detection feature58is positioned within the sensor assembly20in the illustrated embodiment, in some embodiments, the phase detection feature58may be located within the oximeter12. Additionally, the sensor14may include an encoder60(e.g., a resistor or chip) which may be capable of providing signals indicative of the wavelength(s) of light received from the emitter50to allow the oximeter12to select appropriate calibration coefficients for calculating oxygen saturation. The data or signal from the encoder60may be decoded by a detector/decoder feature62in the oximeter12.

In some embodiments, the oximeter12may include a microprocessor64coupled to an internal bus66. Also connected to the bus66may be a memory68(e.g., RAM and/or ROM) and a display70. Received signals from the detector54may be passed through a first amplifier72, a switch74, an analog multiplier76, a low pass filter78, and/or an analog-to-digital converter80. The digital data may then be stored in a queued serial module (QSM)82for later downloading to the memory68as the QSM82fills up. In an embodiment, there may be multiple parallel paths of separate amplifier, filter, and A/D converters for multiple light wavelengths or spectra received, and/or for phase data generated by the phase detector58. In one embodiment, a signal from the phase detector58may be processed in any suitable manner, and may be sent through a different data path than the signal from the detector54, which may be configured to detect amplitude of the photon density waves. The received optical signal may be converted into an electrical signal at the detector54. The electrical signal may then be amplified by the amplifier72and sent to a frequency mixer or analog multiplier (e.g., analog multiplier76) to generate a signal that is proportional to a phase difference between a reference oscillator (not shown) and the received signal. Similarly, the AC and DC amplitudes of the received signal may be determined with peak detection circuits and low pass filters (e.g., filter78).

As illustrated inFIG. 3, the emitter50may include the two LDs56. The LDs56may receive modulated drive signals from the monitor12that activate the LDs56and cause them to emit light at certain intervals. Thus, the monitor12may activate and deactivate the LDs56at high frequencies that may facilitate measurements relating to scattering in the probed medium based on phase changes in emitted photon density waves. This modulation function may be performed by a modulator84. The modulator84may include a hardware feature, a software feature, or some combination thereof. For example, a portion of the modulator84may be stored on the memory68and may be controlled by the processor64. In the illustrated embodiment, the modulator84includes a light driver86and a time processing unit (TPU)88that cooperate to modulate the light emissions of the LDs56. The TPU88, which may include a sine wave generator, may provide timing control signals to the light drive circuitry86, which controls when the emitter50is activated, and if multiple light sources are used, the multiplexed timing for the different light sources. The TPU88may also control the gating-in of signals from the detector54through the first amplifier72and the switching circuit74. These signals are sampled at the proper time, depending at least in part upon which of multiple light sources is activated, if multiple light sources are used.

In the illustrated embodiment, the modulator84is disposed in the monitor12. However, in some embodiments the modulation function may be performed by a modulator disposed within the sensor14. Indeed, it should be noted that in some embodiments, the features related to modulating and detecting the phase of the emitted light waves may be arranged within the system10to avoid potential interference. For example, high frequency modulation and detection features may be co-located within the sensor14to reduce the distance traveled by the signals, and, thus, reduce potential interference. Indeed, in a specific example, the sensor14may include a commercially available chip set for phase measurement and commonly available drive circuits (e.g., DVD R/W driver circuits) for high frequency modulation. Examples of such devices may include the AD8302 available from Analog Devices™ and the LMH6525 available from National Semiconductor™. In other embodiments the LDs56may be positioned within the monitor12and light may be transmitted from the LDs56in the monitor12to the sensor14via fiber optics to reduce potential interference.

Regardless of the modulator's location, in contrast to traditional pulse oximetry, which conducts measurements at sufficiently low frequencies (e.g., 1.5 KHz) to be considered DC, the modulator84may be configured to modulate the LDs56at sufficiently high frequencies (e.g., approximately 50 MHz to 3.0 GHz) to cause resolvable photon density waves to propagate through the tissue52. In some embodiments, the modulator84may be configured to sweep a range from 50 MHz to 2.4 GHz. In some embodiments, the modulator84may be configured to modulate between 100 MHz and 1 GHz or to sweep a range from 100 MHz to 1 GHz. Thus, present embodiments operate at much higher frequencies than the traditional pulse oximetry sampling frequency of 1 sample every 67 microseconds.

In some embodiments, for continuous modulation of the LDs56, resolvable amplitude and phase relationships of the photon density waves may be established at various positions from the emitter along the tissue bed52. By modulating the light emitters at sufficiently high frequencies, the wavelengths of photon density waves may be shorter than the average distance required for light to be absorbed. Thus, the phase changes in the photon density waves can be attributed predominantly to scattering and not absorption. Further, in view of this, it can be determined that detected phase changes correspond to a number of scattering particles or volume change in the probed medium. The frequency of the photon density waves is essentially locked to the initial light source input and the phase change is essentially locked to arterial pulsation and the introduction of scattering particles. Indeed, the variation in AC scattering to DC scattering measured by phase offset may yield information about the total arteriole volume probed.

For a modulation frequency where the product of the frequency and the mean time between absorption events is much larger than 1, the change in phase between two points located a distance r from each other on a tissue bed may be given by the relation.

Δϕ=r⁢ωμs′6⁢c,
where c is the speed of light, ω is the angular frequency of modulation, and μs′ is the reduced scattering coefficient. The reduced scattering coefficient for a tissue bed is comprised of both blood and surrounding tissue components. It can be written as,
μs′total=Vbloodμs′blood+Vtissueμs′tissue.
The time varying component of this equation at a single wavelength will generally be only the portion due to arterial blood. The time varying component of this equation at a second wavelength will allow for the deconvolution of the scattering coefficient. The scattering coefficient for blood is related to the hematocrit (HCT) through the relation,
μs′blood=σs(1−g)(HCT/Vi(1−HCT)(1.4−HCT),
where g is the anisotropy factor, σ is the scattering cross section of an erythrocyte, Vi is the volume of an erythrocyte and HCT is the hematocrit.

Accordingly, when the modulator84operates at a high enough frequency, measured phase changes in the photon density waves may be utilized to calculate a number of scattering particles in the observed volume. For example, the monitor12may be configured to receive phase shift and/or amplitude data from the sensor14and calculate a value related to a quantity of scattering particles in the probed tissue for display on the monitor12. Specifically, the monitor12may include instructions or an algorithm stored on the memory68and configured to perform such calculations.

As an example of the correlation of phase change measurements of photon density waves modulated at high frequency to a number of scattering particles in the probed medium,FIGS. 4A and 4Binclude a pair of graphs that represent simulations of phase changes due to scattering at two different frequencies. Specifically,FIG. 4Aincludes a first graph102andFIG. 4Bincludes a second graph104that each represent simulations of phase change (measured in degrees) due to scattering variation of an arterial pulse (Hemoglobin 15 g/dL) for photon density waves at 890 nm that are modulated with a frequency of 103.4 MHz and 1.034 GHz respectively. It should be noted that the increase in frequency from 103.4 MHz in the first graph102to 1.034 GHz in the second graph104results in a phase change of approximately 3-4 degrees. This change correlates to the wavelengths of the photon density waves. In other words, because the wavelength is reduced even further from the 103.4 MHz modulation rate (first graph102) to the 1.034 GHz modulation rate (second graph104) and there is less opportunity for absorption, the phase change of the higher modulation rate corresponds more specifically to scattering. In some embodiments, a range of frequencies between those shown inFIGS. 4A and 4Bmay be swept through to profile the characteristics of the tissue at different photon density wave frequencies.

Scattering may be quantified based on phase change. Specifically, as set forth above, a modulation frequency where the product of the frequency and the mean time between absorption events is much larger than 1, the change in phase between two points may be given by the relation,

Δϕ=r⁢ωμs′6⁢c.
Changes in phase due to arterial pulsation may be directly related to the change in scattering coefficient of the medium which is due to the change in the concentration of the number of erythrocytes. It should be noted that a second method for correlating the scattering changes from the phase could involve a calibration curve determined from tissue phantoms or clinical data.

FIG. 5illustrates an example of a source modulation signal as driven by cross-coupled LEDs in accordance with some embodiments. Specifically,FIG. 5illustrates a control signal200that may be generated by the modulator84to activate and/or deactivate an emitter including red and IR light sources, such as the LDs56. In other embodiments, separate modulators may be utilized for each light source and/or additional light sources. Indeed, when multiple emitters are utilized, each emitter may be modulated by a separate modulator.

In the illustrated embodiment, the control signal200is representative of dark intervals202, intervals of power204being supplied to a red LD, and intervals of power206being supplied to an IR LD over time. Further, the control signal200has a period designated by reference number208. This period208may be adjusted such that each of the LDs56may be modulated with a desired frequency (e.g., approximately 100-1000 MHz) to generate photon density waves. Such adjustments to the modulation frequency may facilitate detection of phase shifts in the photon density waves, and, thus, variations in scattering based on such phase shifts. As may be appreciated by those of ordinary skill in the art, the control signal200may be adjusted or modified for different scenarios. For example, the control signal200may be adjusted to be generally sinusoidal, adjusted to include various intensity levels, and so forth. The sinusoidal nature of the wave may be controlled by a wave generator and the intensity levels may be adjusted by providing more power and/or by reducing dark intervals and increasing the length of time that light is emitted.

As indicated above, the phase of the photon density waves may be sensitive to changes in the scattering coefficient, while the amplitude of the photon density waves may be sensitive to the concentration of absorbers in the medium. Specifically, with regard to amplitude measurements, the AC amplitude and DC amplitude may yield information about absorption in the volume. Thus, detection of amplitude changes in the photon density waves may be utilized to calculate absorber concentration values in the observed medium, such as blood oxygen saturation values. Such calculations may be made using the standard ratio of ratios (i.e., ratrat) technique for the constant and modulated values of the photon density wave amplitudes at two wavelengths. Once the ratio of ratios values is obtained, it may be mapped to the saturation from clinical calibration curves.

With regard to phase shift measurements, when the wavelengths of the photon density waves get below that of the mean absorption distance, the phase becomes almost exclusively a function of the scattering coefficient. While dependent upon the tissue bed being probed, this is generally believed to occur at a modulation frequency in the range of approximately 500 MHz. Thus, the phase shift measurement may yield information about the number of erythrocytes or red blood cells in the local probed volume. The HCT discussed above is proportional to the number of erythrocytes. Accordingly, by sweeping frequencies, a multi-parameter output may be obtained that relates to standard pulse oximetry measurements as well as the puddle hematorcit.

The amplitude and phase at a given frequency may be proportional to the scattering and absorption coefficient at a given wavelength until the product of the frequency and the mean time between absorption events is much larger than 1. When the product of the frequency and the mean time between absorption events is much larger than 1, the amplitude is a function of the absorption and phase is only a function of the scattering. Thus, a frequency sweep may be used to reduce the error in the determination of a single value of reduced scattering coefficient for the blood and a single value of absorption coefficient. Indeed, in some embodiments, the amplitude and phase information may be utilized together to yield a value of total hemoglobin per unit volume.

In some embodiments, by modulating the light sources at a sufficient frequency, and, thus, facilitating a detectable phase shift that corresponds to scattering particles, present embodiments may provide an extra degree of certainty for blood flow parameter measurements. Indeed, the detected amplitude for the photon density waves may be utilized to calculate traditional pulse oximetry information and the phase may be utilized to confirm that such values are correct (e.g., within a certain range of error). For example, the amplitude information may be utilized to calculate a blood oxygen saturation (SpO2) value and empirical data may indicate that a particular SpO2value should correspond to a particular phase variation at a given frequency. In other words, there may be a certain phase change that should accompany a given increase in absorber observed as a change in amplitude. Various algorithms (e.g., learning based algorithms such as support vector machines, cluster analysis, neural networks, and PCA) based on the measured phase shift and amplitude change may be compared to determine if the amplitude shift and phase shift correlate to a known SpO2. If both the measured amplitude shift and phase shift correlate to a known SpO2, the measured SpO2value may be deemed appropriate and displayed or utilized as a correct SpO2value. Alternatively, if the measured amplitude shift and phase shift do not agree, the calculated SpO2value may be identified as being corrupt or including too much noise and, thus, be discarded.

In some embodiments, as illustrated byFIGS. 6-8, multiple emitter and/or detector arrangements may be utilized in conjunction with one another to provide a transmission mode photon density wave system. Specifically,FIG. 6illustrates a first emitter302and a second emitter304, wherein each of the emitters302,304includes a red and an IR light source (e.g., LED). Waves306represent photon density waves propagating through tissue307from the emitters302,304to a first detector310and a second detector312. During transmission mode operation, the emitters302,304are positioned on a first side of the tissue307and the detectors are positioned on a second side of the tissue307generally opposite the first side. As will be understood by one of ordinary skill in the art, because the multiple emitters are generating separate waves in the same tissue bed, the waves can be made to interfere with one another by adjusting the modulation frequencies of each emitter302,304. Further, it should be noted that the transmission mode setup enables deep penetration to facilitate access to many different regions of interest. Accordingly, multiple emitters may be utilized to steer intensities through the tissue and adjust intensity patterns in the different areas of tissue. For example, the phase of the photon density waves could be adjusted in such a way as to completely cancel out any signal at the first detector310. Thus, if the first detector310detects a signal, it may be an indication of noise.

FIG. 7illustrates an embodiment including multiple emitters400and a single detector402positioned adjacent a patient's tissue404. During transmission mode operation, the emitters400are positioned on a first side of the tissue404and the detector402is positioned on a second side of the tissue404generally opposite the first side. This embodiment may be utilized to generate an adaptive constructive/destructive interference pattern in the tissue bed, including deep within the tissue bed, by adjusting the relative phases of the emitters (at a given wavelength) that would allow for the measurement of local tissue components. These would be visible in the phase and amplitude changes determined by the single detector.

In other embodiments utilizing multiple emitters, the interference of photon density waves may facilitate sweeping photon density waves through a probed volume by changing the relative phase between the emitters. For example, such techniques may be utilized to establish a “phased array” of photon density waves for use in pulse oximetry and hemometry techniques. Indeed, such a “phased array” technique may facilitate identification of regions rich with pulsatile signals in the probed tissue and/or calibration of a sensor through the interference of photon density waves. For example, the phases of individual waves may be controlled to determine the intensity profile within the medium.

It may be desirable to detect regions rich with pulsatile signals to facilitate obtaining a strong pulsatile signal. For example, it may be desirable to focus on a specific location in tissue that includes an artery or even a specific portion of the artery. The transmission mode arrangement may facilitate access to such specific locations by enabling deep penetration. Periodic sweeps may be performed to insure that the focus remains on the pulsation-rich regions. Further, such a technique may define an adaptive measurement system that may be utilized to identify regions of low saturation and/or regions in the probed tissue where blockage may result in anemic conditions. Additionally, it is believed that the use of multiple emitters may facilitate adaptation of the sensor to different physiological variations between patients, such as different skin and/or tissue characteristics.

FIG. 8illustrates an embodiment including multiple detectors500and a single emitter502capable of emitting and detecting photon density waves passed through tissue504. During transmission mode operation, the emitter502may be positioned on a first side of the tissue504and the detectors500may be positioned on a second side of the tissue504generally opposite the first side. The illustrated embodiment may be utilized to identify non-physiological artifact. Each of the multiple detectors500may have a different phase and amplitude relationship with respect to each other. Uncorrelated changes in phase and amplitude between the multiple detectors500would result in a non-physiological artifact such as noise artifact, sensor off, and so forth.

The inclusion of multiple detectors around a tissue bed may facilitate detection of and/or compensation for a variety of noise artifacts that typically plague existing pulse oximetry technologies. Indeed, for a given wavelength, a time-varying phase and amplitude relation between multiple detectors may be established which is correlated to arterial pulse. The phase and amplitude information may form a phase space that yields a bounded parameter space for a single wavelength that contains physiological measurements. Noise artifacts will typically lie outside of this bounded area, as will be discussed in further detail below. Further, the addition of a second wavelength may facilitate formation of a 4-dimensional physiological measurement space that facilitates noise artifact reduction due to constraints of decision planes in the hyperspace. Correlated phase and amplitude changes for a single wavelength are bounded by physiological parameters such as arteriole density, realistic hematocrit numbers, and so forth. At a single wavelength, these bounds result in bounds on the detected amplitude and phase in a 2D space. These same bounds are applicable for a second wavelength. The 4 factor correlation (phase(wavelength1), phase wavelength2), amplitude(wavelength1), amplitude(wavelength2)) is bounded by physiological factors in a linked 4D space. The bounds can be drawn as hyperplanes in that space. For example, cluster analysis, Neural Networks, and partial least squares (PLS) algorithms may be used to generate the decision planes and compensate for a variety of noise artifact.

In some embodiments, and as an example,FIG. 9includes a 2-dimensional plot600that represents a physiological state602characterized by amplitude604and phase shifts606. Once phase shift and/or amplitude data has been properly characterized based on empirical data, certain correlations may be indicative of a change in pressure (e.g., a sensor is attached too tightly), a certain area of tissue being subject to exsanguination, a sensor being off noise being present, and so forth. The plot600is representative of a single wavelength at a given frequency. Thus, multiple wavelengths at a given frequency would each have this type of physiological space for expected amplitude and phase variation. Noise artifact608will generally lie outside of this bounded parameter space or physiological regime. Accordingly, if a measurement falls outside of the physiological regime, it may be discarded as including too much noise. When a measurement is discarded, it may be replaced with the previous measurement or some combination of historical values. For example, historical values may be averaged using an averaging routine to provide a replacement for the noisy current measurement value.