Patent Publication Number: US-2011060197-A1

Title: Near infrared spectrophotometry with enhanced signal to noise performance

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     This patent application claims the benefit of U.S. Provisional Ser. No. 61/224,684, filed Jul. 10, 2009. This patent application is a continuation-in-part of U.S. Ser. No. 12/826,218, filed Jun. 29, 2010 (Atty. Dkt. 6949/81720), which claims the benefit of U.S. Provisional Ser. No. 61/222,099, filed Jun. 30, 2009, and which also claims the benefit of U.S. Provisional Ser. No. 61/255,851, filed Oct. 28, 2009. Each of the above-referenced patent applications is incorporated by reference herein. 
    
    
     FIELD 
     This patent specification relates to the non-invasive monitoring of a physiological condition of a patient using information from non-invasive near-infrared (NIR) optical scans. More particularly, this patent specification relates to systems, methods, and related computer program products for improving signal to noise performance in the non-invasive near-infrared spectrophotometric (NIRS) monitoring of chromophore levels in biological tissue. 
     BACKGROUND AND SUMMARY 
     The use of near-infrared (NIR) light as a basis for the measurement of biological properties or conditions in living tissue is particularly appealing because of its relative safety as compared, for example, to the use of ionizing radiation. Various techniques have been proposed for non-invasive NIR spectroscopy or NIR spectrophotometry (NIRS) of biological tissue. Generally speaking, these techniques are directed to detecting the concentrations of one or more chromophores in the biological tissue, such as blood hemoglobin in oxygenated (HbO) and deoxygenated (Hb) states. 
     As used herein, NIR tissue oxygenation level monitoring refers to the introduction of NIR radiation (e.g., in the 500-2000 nm range) into a tissue volume and the processing of received NIR radiation migrating outward from the tissue volume to generate at least one metric indicative of oxygenation level(s) in the tissue. One example of an oxygenation level metric is oxygen saturation [SO 2 ], which refers to the fraction or percentage of total hemoglobin [HbT] that is oxygenated hemoglobin [HbO]. NIRS-based oxygen saturation readings can be classified as “relative” in nature (i.e., presented only in terms of their change over time) or can be “absolute” in nature (i.e., computed from absolute concentrations of [HbO] and [HbT] in units of grams per deciliter (g/dl) or equivalent). 
     NIR cerebral oxygenation level monitoring, which refers to the transcranial introduction of NIR radiation into the intracranial compartment and the processing of received NIR radiation migrating outward therefrom to generate at least one metric indicative of oxygenation level(s) in the brain, represents one particularly important type NIR tissue oxygenation level monitoring. One exemplary need for reliable determination of oxygen saturation levels in the human brain arises in the context of the millions of surgical procedures performed under general anesthesia every year. One statistic recited in U.S. Pat. No. 5,902,235 is that at least 2,000 patients die each year in the United States alone due to anesthetic accidents, while numerous other such incidents result in at least some amount of brain damage. Certain surgical procedures, particularly of a neurological, cardiac or vascular nature, may require induced low blood flow or pressure conditions, which inevitably involves the potential of insufficient oxygen delivery to the brain. Many surgical procedures also involve the possibility that a blood clot or other clottable material can break free, or otherwise get introduced into the bloodstream, and travel to the brain to cause a localized or widespread ischemic event therein. At the same time, the brain is highly intolerant to oxygen deprivation, and brain cells will die (become infarcted) within a few minutes if not sufficiently oxygenated. Accordingly, the availability of immediate, accurate and reliable information concerning brain oxygenation levels is of critical importance to anesthesiologists and surgeons, as well as other involved medical practitioners. 
     Pulse oximetry, in which infrared sources and detectors are placed across a thin part of the patient&#39;s anatomy such as a fingertip or earlobe, has arisen as a standard of care for all operating room procedures. However, pulse oximetry provides only a general measure of blood oxygenation as represented by the blood passing by the fingertip or earlobe sensor, and does not provide a measure of oxygen levels in vital organs such as the brain. In this sense, the surgeons in the operating room essentially “fly blind” with respect to brain oxygenation levels, which can be a major source of risk for patients (e.g., stroke) as well as a major source of cost and liability issues for hospitals and medical insurers. 
     Valid NIR cerebral oxygenation level readings provide crucial monitoring data for the surgeon and other attending medical personnel, providing more direct data on brain oxygenation levels than pulse oximeters while being just as safe and non-invasive as pulse oximeters. Generally speaking, such systems involve the attachment of an NIR probe patch, or multiple such NIR probe patches, to the forehead and/or other available skin surface of the head. Each NIR probe patch usually comprises one or more NIR optical source ports for introducing NIR radiation into the cerebral tissue and one or more NIR optical receiver ports for detecting NIR radiation that has migrated through at least a portion of the cerebral tissue. One or more oxygenation level metrics are then provided on a viewable display in a digital readout and/or graphical format. 
     One issue that arises in NIR cerebral oximetry is the need for substantial signal penetration depth in order to obtain useful readings for the brain tissue itself, which lies beneath several intervening layers including the skin, scalp, skull, dura, and cerebrospinal fluid (CSF) layers. According to one thumbnail estimate provided in U.S. Pat. No. 5,853,370, which is incorporated by reference herein, the average penetration depth for a NIRS source-detector pair is about one-half of the lateral separation between the source and the detector. Thus, to acquire meaningful readings for brain tissue at a depth of about 3 cm from the skin surface, the source-detector distance needs to be about 6 cm. However, due to the high degree of signal degradation involved, such relatively large source-detector distances have not been provided in known commercially available NIR cerebral oximeters. It would be desirable to provide an NIR cerebral oximeter with improved signal-to-noise performance in order to accommodate such relatively large source-detector distances. Furthermore, improved signal to noise performance would also increase the accuracy and/or reliability of the readings provided for more closely-spaced source-detector pairs. Other issues arise as would be apparent to one skilled in the art upon reading the present disclosure. 
     It is to be appreciated that although one or more preferred embodiments is detailed hereinbelow in the particular context of NIR cerebral oxygenation level monitoring (NIR cerebral oximetry), the present teachings are readily applicable to the non-invasive spectrophotometric monitoring of any of a variety of different body parts including, but not limited to, the kidney, lung, liver, arm, leg, neck, etc., and furthermore are applicable for the monitoring of any of a variety of different chromophore types therein. 
     Provided according to one or more preferred embodiments are methods, systems, and related computer program products for non-invasive spectrophotometric monitoring of an optical property of a tissue volume during a patient monitoring session. A plurality of optical sources and a plurality of optical detectors are secured to a surface of the tissue volume. The plurality of optical sources are operated to introduce, simultaneously and on a continuous basis throughout the patient monitoring session, a plurality of optical signals into the tissue volume. Preferably, each of the optical signals has a modulation frequency different than that of each other optical signal, and any two of the optical signals that are introduced from a same one of the optical sources are at different optical wavelengths. The plurality of optical detectors are each operated to detect, simultaneously and on a continuous basis throughout the monitoring session, a portion of each of the optical signals that has propagated thereto, and each of the detected optical signal portions is processed to derive an amplitude signal and a phase signal associated therewith. The derived amplitude signals and phase signals associated with the detected optical signal portions are then processed to determine the optical property of the tissue volume. 
     Also provided is an apparatus for non-invasive spectrophotometric monitoring of an optical property of a tissue volume of a patient during a patient monitoring session. The apparatus comprises a probe patch wearable on a surface of the tissue volume, the probe patch comprising a plurality of optical sources and a plurality of optical detectors. The probe patch is configured to maintain each of the optical sources and each of the optical detectors in secured contact with the surface of the tissue volume throughout the patient monitoring session. The apparatus further comprises a source controller coupled to each of the plurality of optical sources, the source controller being configured to cause the plurality of optical sources to introduce, simultaneously and on a continuous basis throughout the patient monitoring session, a plurality of optical signals into the tissue volume, each optical signal having a modulation frequency different than that of each other optical signal, wherein any two of the optical signals that are introduced from a same one of the optical sources are at different optical wavelengths. The apparatus further comprises a detector controller coupled to each of the plurality of optical detectors, the detector controller being configured to cause each of the plurality of optical detectors to detect, simultaneously and on a continuous basis throughout the monitoring session, a portion of each of the optical signals that has propagated thereto. The apparatus further comprises at least one processor configured to process each of the detected optical signal portions to derive an amplitude signal and a phase signal associated therewith, the at least one processor being further configured to process the amplitude signals and phase signals associated with the detected optical signal portions to determine the optical property of the tissue volume. 
     Also provided is an apparatus for non-invasive spectrophotometric monitoring of an optical property of a tissue volume of a patient during a patient monitoring session, comprising a probe patch wearable on a surface of the tissue volume of the patient. A first optical source and a first optical detector are disposed on the probe patch. The probe patch is configured to maintain each of the first optical source and the first optical detector in secured contact with the surface of the tissue volume throughout the patient monitoring session. The first optical detector includes a first aperture formed in a tissue-facing surface of the wearable patch. The first aperture includes a central area, a first edge positioned nearer to the first optical source than the central area, and a second edge positioned farther from the first optical source than the central area. Preferably, the first and second edges of the first aperture are each curved concavely toward the first optical source. 
     Among other advantages, non-invasive near-infrared spectrophotometric monitoring according to one or more of the preferred embodiments provides for improved signal to noise performance. Among other advantages, the improved signal to noise performance provides an ability to increase penetration depths in the non-invasive NIRS monitoring of crucial deep-layer tissue structures including, but not limited to, the human brain. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  illustrates a prior art bilateral cerebral spectrophotometric monitoring system; 
         FIGS. 2A-2C  illustrate a prior art slope method used in spectrophotometric monitoring; 
         FIGS. 3A-3C  illustrate equations used in particular prior art phase modulated spectrophotometric (PMS) and continuous wave (CW) spectrophotometric monitoring scenarios; 
         FIG. 4  illustrates a prior art arrangement of non-ideal optical sources and detectors on a probe patch; 
         FIG. 5  illustrates an intensity-based slope computation based on a prior art symmetric source-detector layout; 
         FIG. 6  illustrates an intensity-based slope computation based on a prior art symmetric source-detector layout; 
         FIG. 7A  illustrates a near-infrared spectrophotometric (NIR) cerebral oximeter according to a preferred embodiment; 
         FIG. 7B-1  illustrates an NIR probe patch according to a preferred embodiment; 
         FIG. 7B-2  illustrates exemplary dimensions associated with the NIR probe patch of  FIG. 7B-1 ; 
         FIG. 7C  illustrates an NIR probe patch according to a preferred embodiment; 
         FIGS. 7D-1  and  7 D- 2  illustrate an NIR probe patch according to a preferred embodiment; 
         FIG. 7E  illustrates the NIR probe patch of  FIG. 7D-1  as applied to a surface of a biological volume according to a preferred embodiment; 
         FIG. 7F  illustrates dual instances of the NIR probe patch of  FIG. 7D-1  as applied to a forehead of a patient for bilateral cerebral oximetry according to a preferred embodiment; 
         FIG. 8A  illustrates near-infrared spectrophotometric (NIRS) monitoring of a biological volume of a patient according to a preferred embodiment; 
         FIG. 8B  illustrates an alternative version of the probe patch illustrated in  FIG. 8A  according to a preferred embodiment; 
         FIGS. 9A-9C  illustrate equations for adapting a slope method for NIRS monitoring of a biological volume according to a preferred embodiment; 
         FIG. 10  illustrates NIRS monitoring of a biological volume of a patient according to a preferred embodiment; 
         FIG. 11  illustrates NIRS monitoring of a tissue volume in which there is a relatively low duty cycle for any particular source/wavelength pair; 
         FIG. 12  illustrates NIRS monitoring of a tissue volume of a patient according to a preferred embodiment; 
         FIG. 13  illustrates a block diagram of signal processing circuitry for use with a detector of a NIRS monitoring system according to a preferred embodiment; 
         FIG. 14  illustrates NIRS monitoring of a tissue volume of a patient according to a preferred embodiment; 
         FIGS. 15A-15B  illustrate an NIR probe patch including curved-edge detection apertures according to a preferred embodiment; 
         FIG. 16  illustrates an NIR probe patch including curved-edge detection apertures according to a preferred embodiment; and 
         FIG. 17  illustrates NIRS monitoring of a tissue volume of a patient according to a preferred embodiment. 
     
    
    
     DETAILED DESCRIPTION 
       FIG. 1  illustrates a prior art proposal for a bilateral monitoring system in which two NIR probe patches  16  and  116  are placed on the forehead of the patient. The prior art proposal of  FIG. 1  is further described in U.S. Pat. No. 6,615,065, which is incorporated by reference herein. Separate readings for the left and right sides of the brain are acquired and displayed separately on an output display  20 . As illustrated in the proposal of  FIG. 1 , NIR probe patches are often placed on the forehead of the patient. The forehead represents a generally desirable region for attaching NIR probe patches, for at least the reason that the forehead is generally free of hair follicles. Even for a smoothly shaved head, the presence of hair follicles can introduce substantial amounts of noise and other interference into the NIR signals. 
     However, the use in  FIG. 1  of two separate NIR probe patches on the forehead is antagonistic to an even more important goal of NIR cerebral oximetry, which is to obtain “deep” readings that are relevant to the brain tissue, rather than to the intervening skin, scalp, skull, dura, and cerebrospinal fluid (CSF) tissue. According to one thumbnail estimate provided in U.S. Pat. No. 5,853,370, which is incorporated by reference herein, the average penetration depth for a NIRS source-detector pair is about one-half of the lateral separation between the source and the detector. Because the source and the detector for any particular source-detector pair are required to be present on the same NIR probe patch (due to the need for precise, predetermined source-detector separation distances), the maximum source-detector separation distance for the prior art proposal of  FIG. 1  is limited by the spatial extent of each individual NIR probe patch  16  and  116 . Moreover, the use in  FIG. 1  of two separate NIR probe patches on the forehead also brings about the need for left-right duplication of multiple source-detector pairs in order to obviate source intensity differences, detector efficiency differences, and skin coupling efficiency differences among the sources and detectors  116 . 
       FIGS. 2A-2C  illustrate a prior art slope method used in spectrophotometric monitoring.  FIGS. 3A-3C  illustrate equations used in particular prior art phase modulated spectrophotometric (PMS) and continuous wave (CW) spectrophotometric monitoring scenarios. Summarized in  FIGS. 2A-2C  and  3 A- 3 C is the well-accepted “slope method” for computing tissue oxygenation levels (see, e.g., Fantini, Franceschini, and Gratton, “Semi-Infinite-Geometry Boundary Problem For Light Migration In Highly Scattering Media: A Frequency-Domain Study In The Diffusion Approximation,” J. Opt. Soc. Am. B, Vol. 11, pp. 2128-38 (1994) and Fantini, Hueber, and Franceschini, et. al., “Non-Invasive Optical Monitoring of the Newborn Piglet Brain Using Continuous-Wave and Frequency-Domain Spectroscopy,” Phys. Med. Biol., Vol. 44, pp. 1543-1563 (1999), each of which is incorporated by reference herein), while  FIGS. 4-6  set forth one known method (see, e.g., U.S. Pat. No. 6,078,833, which is incorporated by reference herein) for using multiple source-detector pairs positioned over a common region to obviate source intensity differences, detector efficiency differences, and skin coupling efficiency differences among the sources and detectors. 
     Notationally, the prime symbol (′) is used to denote ideal intensities (I′) and ideal phases (φ′) that would result from ideal sources and ideal detectors (including ideal skin coupling), as well as ideal slopes (K′) of any plotted functions based on those ideal intensities and phases. In contrast, non-primed versions of those quantities refer to the physically measured versions of those values in the real world, and are termed herein measured intensities (I) and measured phases (φ), as well as measured slopes (K) of the plotted functions based on the measured intensities and measured phases. For PMS (phase modulated spectrophotometry) systems, also termed frequency domain spectrophotometry systems, the basis of the slope method is that for any particular NIR radiation wavelength, a plot of log (r 2 l′) versus r (where r is the source-detector distance) ( FIG. 2B ) has a relatively constant slope K a ′ over an appreciably useful range of distances, a plot of φ′ versus r ( FIG. 2C ) also has a relatively constant slope K p ′ over an appreciably useful range of distances, and the values of K a ′ and K p ′ can be used to compute the absorption coefficient μ a  ( FIG. 3A , Eq. {3A-1}) and the effective or reduced scattering coefficient μ s ′ ( FIG. 3A , Eq. {3A-2}) for that NIR radiation wavelength, where ω is the angular frequency corresponding to the source intensity modulation and v is the speed of light in the tissue. For CW (continuous wave) spectrophotometry systems in which there is no high-frequency modulation or phase measurements, the value of K a ′ can be used to compute the absorption coefficient μ a  ( FIG. 3B , Eq. {3B-1}) for that NIR radiation wavelength using a fixed estimate of the effective scattering coefficient μ s ′. Based on the absorption coefficient μ a  for multiple NIR wavelengths (on opposite sides of the isosbestic wavelength for oxygenated and deoxygenated hemoglobin) the oxygenated hemoglobin saturation value SO2 is then readily determined, with {Eq. 3C-1} setting forth the formula for the particular NIR wavelengths of 680 nm and 830 nm. Generally speaking, the SO2 reading for the PMS-based measurements can be characterized as an absolute percentage value, whereas the SO2 reading for CW measurements should be taken only as a relative value over time. 
     As would be readily understood by a person skilled in the art in view of the present disclosure, the term “intensity” (as well as the equation variable “I” in the accompanying drawings) as used herein in the context of a PMS system refers to the amplitude of the AC component of the intensity waveform. Thus, without loss of generality, the terms “amplitude” and “intensity” may be used interchangeably herein to refer to the amplitude of the AC component of the intensity waveform (see, e.g., the Fantini 1999 article, supra, at Section 2.3 thereof). 
       FIG. 4  illustrates a prior art arrangement of non-ideal optical sources and detectors on a probe patch. As illustrated in  FIG. 4 , a non-ideal source S can be modeled as an ideal source as modified by a complex coefficient η s exp(−iθ S ), which is termed herein the source intensity/coupling coefficient. For simplicity of nomenclature, although the magnitude η s  is more generally associated with variations in both source intensity and skin coupling, the magnitude η s  is simply referenced herein as “source coupling efficiency.” The phase term θ s  is referred to herein as the “source phase error.” Likewise, as illustrated in  FIG. 4 , a non-ideal detector D can be modeled as an ideal detector as modified by a complex coefficient η D exp(−iθ D ), which is termed herein the detector sensitivity/coupling coefficient. For simplicity of nomenclature, although the magnitude η D  is more generally associated with variations in both detector sensitivity and skin coupling, the magnitude η D  is simply referenced herein as “detector coupling efficiency.” The phase term θ D  is referred to herein as the “detector phase error.” 
       FIG. 5  illustrates an intensity-based slope computation based on a symmetric source-detector layout according to the prior art.  FIG. 6  illustrates an intensity-based slope computation based on a symmetric source-detector layout according to the prior art. For simplicity and clarity of explanation, the more general case of PMS modulation is detailed further herein, with it being understood that CW methods would be analogous except with omitted phase factors and omitted phase-related slope computations. In the event that a real-world source  51  (and real-world source-skin coupling) was used and two real-world detectors D 1  and D 2  (with real-world detector-skin coupling) were positioned at r 1  and r 2 , respectively, in the configuration of  FIG. 2A , it could readily be shown that the values of μ a  and μ s ′ would include unknown coupling efficiency and phase error factors in addition to the known measured intensities I 12  and I 22 . Because the coupling efficiency and phase error factors are unknown, the values of μ a  and μ s ′ would either be non-determinable, or else broad assumptions regarding coupling efficiency and phase error factors would need to be made. However, as summarized in  FIGS. 4-6  and described further in U.S. Pat. No. 6,078,833, supra, the presence of different coupling efficiencies can be obviated by (i) adding a second source S 2 , (ii) positioning the two sources S 1 , S 2  and two detectors D 1 , D 2  in a symmetric relationship such that r 21 =r 12  and r 11 =r 22 , (iii) computing a first measured slope factor K a,D1  representing the slope factor of  FIG. 2B  for the underlying tissue as “seen” by detector D 1 , (iv) computing a second measured slope factor K a,D2  representing the slope factor of  FIG. 2B  for that same underlying tissue as “seen” by detector D 2 , and (v) computing an overall measured slope K a  as the arithmetic average of K a,D1  and K a,D2 . As illustrated in  FIG. 5 , the coupling efficiencies cancel out such that the measured K a  becomes equal to the average of the ideal slopes K′ a,D1  and K′ a,D2 , which is tantamount to an overall ideal slope K′ a . As illustrated in  FIG. 6 , the presence of different phase error factors is similarly obviated when r 21 =r 12  and r 11 =r 22 , the phase error factors canceling and the overall measured phase slope K p  becoming equal to the average of the ideal slopes K′ p,D1  and K′ p,D2 , which is tantamount to an overall ideal slope K′ p . The resultant values of μ a , μ s ′, and SO2 are thus independent of the coupling efficiencies and phase error factors, which is indeed a desirable result. 
     However, as mentioned above, in order for the system of  FIG. 1  to achieve this desirable result (i.e., the obviation of source intensity differences, detector efficiency differences, and skin coupling efficiency differences) it is required that each of the left and right NIR probe patches contain a dual arrangement (see  FIG. 4 ) of source-detector pairs for each source-detector separation distance of interest. For a single source-detector separation distance, a 2×2 arrangement (two sources, two detectors, see  FIG. 4 ) is required for each NIR probe patch, thereby requiring a total of eight elements (four sources and four detectors) for the bilateral system. For two source-detector separation distances (for example, a “near” separation distance and a “far” separation distance), a 2×4 arrangement (two sources and four detectors, or four sources and two detectors) is required for each NIR probe patch, thereby requiring a total of twelve elements (four sources and eight detectors, or eight sources and four detectors) for the bilateral system. For three source-detector separation distances (for example, a “near” separation distance, a “mid-range” separation distance, and a “far” separation distance), a 2×6 arrangement (two sources and six detectors, or six sources and two detectors) is required for each NIR probe patch, thereby requiring a total of sixteen elements (four sources and twelve detectors, or twelve sources and four detectors) for the bilateral system. In general, for “N” distinct source-detector separation distances, a (2N+2) arrangement is required for each NIR probe patch, thereby requiring a total of 2(2N+2)=4(N+1) elements for the bilateral system. 
     Provided according to one preferred embodiment is an NIR cerebral oximeter comprising a unitary across-the-forehead (ATF) patch configured and dimensioned to cover both the left and right sides of the forehead simultaneously, the ATF patch comprising a lateral distribution of NIR sources and detectors including either (i) a plurality of centrally located sources and at least one detector near each of the left and right ends, or (ii) a plurality of centrally located detectors and at least one source near each of the left and right ends, wherein each of the centrally located sources or detectors is used in determining each of (i) an overall chromophore level applicable for the combined left and right sides of the brain, (ii) (ii) a left-side chromophore level separately applicable for the left side of the brain, and (iii) a right-side chromophore level separately applicable for the right side of the brain. While one or more preferred embodiments is described in terms of an across-the-forehead patch for monitoring the left and right brain hemispheres simultaneously, it is to be appreciated that the present teachings further encompass a wide variety of different probe patches capable of simultaneous monitoring of two subregions of tissue that are at least partially non-overlapping, and that the ATF forehead represents but one particularly useful example. Thus, for example, there could be provided in accordance with another preferred embodiment a user-wearable probe patch for monitoring a single kidney, where the first subregion corresponds primarily to an upper part of the kidney and the second subregion corresponds primarily to a lower part of the kidney. As another example, there could be provided in accordance with another preferred embodiment a user-wearable probe patch for monitoring both kidneys, where the first subregion corresponds primarily to a left kidney and the second subregion corresponds primarily to a right kidney. 
     Also provided according to a preferred embodiment is an algorithm for bilateral chromophore level monitoring based on measurements acquired using the ATF patch sources and detectors, wherein the bilateral chromophore levels are computed in a manner that obviates any coupling efficiency differences or phase error differences among the different sources and detectors, subject only to certain relaxed time-invariance assumptions for the centrally located sources or detectors (specifically, that they exhibit a constant coupling efficiency ratio and a constant phase error difference between them during the monitoring session). Advantageously, because each of the centrally located sources or detectors is involved in the individual monitoring of each of the left and right sides, bilateral monitoring is provided using a reduced number of elements as compared to the use of two separate forehead patches. Advantageously, the spatial geometry of the source/detector elements on the ATF patch provides for increased source-detector separation so that deeper penetration depths into the brain can be achieved in comparison to the use of two separate forehead patches. 
     As used herein, the term or subscript “whole” is used to refer to a measurement or output reading that is applicable for the combined left and right side tissue of the brain. As will be understood by a person skilled in the art, the terms “whole brain,” “left side of the brain,” and “right side of the brain” as used herein, and unless otherwise stated, refer to those portions that are forward in the skull cavity toward the forehead and reachable by a relevant portion of the NIR radiation that has been introduced into the forehead. The unitary across-the-forehead (ATF) patch can alternatively be termed a whole-forehead patch, cross-forehead patch, or total-forehead patch. Preferably, PMS (phase modulated spectrophotometry) methods are used in conjunction with the ATF sources and detectors such that the absorption coefficient and effective scattering coefficient are each computed for each of a plurality of NIR wavelengths, and absolute SO2 values are provided. However, the preferred embodiments described herein can readily be applied in CW (continuous wave) systems. For simplicity and clarity of explanation, the more general case of PMS modulation is detailed further herein. 
     It has been found that accurate, clinically useful, absolute, reduced source/detector bilateral SO2 monitoring based on an ATF patch according to one or more of the preferred embodiments can be achieved based on certain clinically reasonable usage and parameter assumptions. A first assumption is that there is a generally quiescent time period at the beginning of a monitoring session in which the whole brain, including both the left and right sides together, can be considered to have a generally uniform SO2 value. This assumption is particularly realistic and useful for exemplary scenarios such as surgery, in which it can be assumed that no blood clots have broken free and traveled to the brain prior to the surgery (for example), and it which case it will be particularly useful to localize which side of the brain a clot is affecting if such an event occurs during the surgery. 
     A second assumption is that the coupling efficiencies and phase errors of the centrally located sources (or centrally located detectors) exhibit certain time-invariance requirements that are “relaxed” in the sense that it is not strictly required that each of them remains absolutely fixed during the monitoring session. More particularly, it only needs to be assumed that the ratio of the coupling efficiencies of the centrally located sources (or centrally located detectors) remains constant during the monitoring session, and that the difference between phase errors for the centrally located sources (or centrally located detectors) remains constant during the monitoring session. These time-invariance criteria are more relaxed than a “strict” time-invariance criteria in which all coupling efficiencies and phase errors of all sources and detectors must remain fixed during the monitoring session. Notably, because the centrally located sources (or centrally located detectors) are physically nearby to each other and nestled well within the interior confines of the ATF patch, it is believed particularly realistic that the ratio of their coupling efficiencies, if not the actual values of their coupling efficiencies, will tend to remain constant throughout the monitoring session. More generally stated, one or more of the preferred embodiments described further herein is advantageously applied when it can be assumed that the particular biological volume under study has a characteristic at the beginning of the monitoring period (which can be termed a calibration period) in which both of the localized subregions (or “N” subregions if there are more than two subregions being monitored) can be considered to have a generally uniform value for the optical property to be monitored. 
       FIG. 7A  illustrates an NIR cerebral oximeter  702  according to a preferred embodiment, comprising an across-the-forehead (ATF) probe patch  704  coupled via optical, electro-optical, or electrical cables  706  to a console unit  708 . Console unit  708  comprises one or more optical sources  710  and optical detectors  712 , each of which may be fully optical, electro-optical, or fully electrical in nature depending on the nature of the sources and detectors on the probe patch  704 . For one preferred embodiment, the optical sources  710  comprise one or more laser sources, the optical detectors  712  comprise one or more photomultiplier tubes (PMTs), and the probe patch  704  consists of passive optical sources and detectors and has a general overall construction similar to one or more of the NIR probe patches disclosed in the commonly assigned and U.S. Ser. No. 12/483,610 filed Jun. 12, 2009 with the dimensions, source locations, and detector locations being as set forth herein. Console unit  708  further comprises a processor  714  coupled to control and receive information from the optical sources  710  and optical detectors  712 , the processor  714  being configured, dimensioned, and programmed to achieve the functionalities described herein. Console unit  708  further comprises an output display  716  coupled to the processor  714  that simultaneously displays left, right, and whole-brain SO2 readings (and, optionally, intermediate values such as slopes, absorption coefficients, and scattering coefficients) in any of a variety of numerical and/or graphical formats. Among a variety of other control inputs, the console unit  708  further comprises a “start” button  718  that allows for user initiation of the SO2 monitoring session. The “start” button  718  can alternatively be termed a calibration button, as it instantiates a calibration process in which particular algorithm compensations (and/or other parameters) are determined based on a presumption that the optical property to be monitored is spatially homogenous throughout the different subregions of monitored tissue at that “start” time or calibration time. 
       FIGS. 7B-1  and  7 B- 2  illustrate a simplified version of the probe patch  704  and dimensions relevant thereto according to one preferred embodiment, the probe patch  704  having only two sources S 1 -S 2  and two detectors D 1 -D 2  positioned as shown. Different ATF probe patches having different source-detector separation distances can be provided for differently size foreheads. In other preferred embodiments there are additional sets of detectors for providing readings that are applicable for additional source-detector separation distances. 
       FIG. 7C  illustrates a simplified version of an alternative probe patch  754  that can be used in conjunction with the NIR cerebral oximeter  702  according to a preferred embodiment. Advantageously, as will be illustrated further infra, it is not required that the prior art symmetries of  FIG. 4  be present in order to achieve the desired monitoring functionalities according to the preferred embodiments, and thus the probe patch  754  is shown without those symmetries present. 
       FIGS. 7D-1  and  7 D- 2  illustrate a simplified version of an alternative probe patch  755  that can be used in conjunction with the NIR cerebral oximeter  702  according to a preferred embodiment. Whereas the non-symmetric probe patch  754  still maintains a somewhat linear configuration that defines left and right subregions (albeit non-symmetrically), analogous to that of the probe patch  704 , the non-symmetric probe patch  755  represents a more quadrilateral-shaped configuration that is applicable to a more compact region of tissue. For the probe patch  755 , it is required only that the sources and detectors be laid out so as to define plural subregions that are at least partially non-overlapping with each other. As illustrated in  FIG. 7D-2 , each partially non-overlapping subregion is defined by either a single detector with two sources of differing distances therefrom (to allow the above-described slope method to be applicable) or, alternatively, a single source with two detectors of differing distances therefrom. 
       FIG. 7E  illustrates the probe patch  755  of  FIGS. 7D-1  and  7 D- 2  as mounted on a surface  791  of a biological volume  790 , for monitoring an optical property of the subsurface tissue  792 . The biological volume  790  can generally be any part of the body, and is not limited to the head of the patient. 
       FIG. 7F  illustrates NIR cerebral oximetry based on the probe patch  755  of  FIGS. 7D-1  and  7 D- 2 , wherein there are two probe patches  755  coupled to respective sides of the forehead of the patient. For the scenario of  FIG. 7F , each probe patch  755  can provide optical property readings for two subregions (e.g., an “upper” subregion and “lower” subregion, see  FIG. 7D-2 ) for its respective hemisphere, and/or each probe patch  755  can provide a single reading for its respective hemisphere based on an averaging or other combination of the two subregions. 
     In keeping with the bidirectional nature of light, for each of the preferred embodiments herein there exists a converse configuration in the form of swapped source-detector positions that is also a preferred embodiment within the scope of the present teachings and that operates in essentially the same way. For example, with reference to  FIG. 7B-1 , an alternative converse configuration exists in which the detectors D 1  and D 2  are in the center of the probe patch, and the sources S 1  and S 2  are at the lateral peripheries of the probe patch. The relevant mathematical formulae and functional operation of these conversely configured preferred embodiments would be readily apparent to a person skilled in the art in view of present disclosure, and need not be discussed further herein. 
     For any particular ATF patch, the operational methods and computations for the different source-detector quadruplets thereon are generally independent of each other. For example, referring briefly to the probe patch of  FIG. 14 , infra, measurements corresponding to the S 1 -S 2 /D 1 -D 2  quadruplet shown in  FIG. 14  can be processed to compute a first absolute SO2 value, and a separate set of measurements corresponding to the S 1 -S 2 /D 3 -D 4  quadruplet can be processed to compute a second absolute SO2 value, with there being no dependencies between the two sets of computations. The multiple SO2 readings (and/or the multiple underlying values of the slopes, absorption coefficients, effective scattering coefficients, etc., at each wavelength) for the multiple source-detector quadruplets can be processed in any of a variety of different advantageous ways without departing from the scope of the present teachings. For example, in a two-quadruplet scenario (see  FIG. 14 ) in which there is a “near” quadruplet (S 1 -S 2 /D 3 -D 4 ) and a “far” quadruplet (S 1 -S 2 /D 1 -D 2 ), the “near” readings associated with lesser penetration depths can be processed in conjunction with the “far” readings associated with deeper penetration depths to extract outputs more specific to the deep brain tissue. In one preferred embodiment, the different “near” and “far” readings are processed as described in the commonly assigned U.S. Ser. No. 12/815,696, filed Jun. 15, 2010, which is incorporated by reference herein. Because the computations for different source-detector quadruplets are substantially the same and generally independent of each other, the preferred methods for bilateral and whole-head SO2 monitoring will be detailed further herein for the simplified, single quadruplet system (S 1 -S 2 /D 1 -D 2 ) of  FIG. 7C . 
       FIG. 8A  illustrates near-infrared spectrophotometric (NIR) monitoring of a biological volume of a patient according to a preferred embodiment. At step  802 , the NIR sources and detectors, as contained for example on the probe patch  754 , are secured to a surface of the biological volume. Referring ahead briefly to  FIG. 8B , in keeping with the bidirectional nature of light, there exists a converse probe patch  754 ′ for which the present description is equivalently applicable, in the form of swapped source-detector positions relative to the probe patch  754 . Upon mounting and securing of the probe patch, a calibration interval can begin, such as by the user pressing the “start” button  718 , which is followed by a monitoring interval. The calibration interval should usually last a few seconds, but can be substantially lesser or greater without departing from the scope of the present teachings. The monitoring interval can be anywhere from a few minutes to several hours, depending on the nature of the clinical procedure (e.g., during surgery, during post-operative recovery, during other patient testing, etc.) in association with which the patient monitoring may be taking place. During each of a calibration interval and the subsequent monitoring interval (step  804 ), a first portion of light (denoted “A” in  FIG. 8A ) is propagated from a first optical source S 1  through the medium to the first optical detector D 1 , a second portion of light (“B”) is propagated from the second optical source S 2  through the medium to the first optical detector D 1 , a third portion of light (“C”) is propagated from the first optical source S 1  through the medium to the second optical detector D 2 , and a fourth portion of light (“D”) is propagated from the second optical through the medium to the second optical detector. 
     At step  806 , detections of the first light portion “A”, second light portion “B”, third light portion “C”, and fourth light portion “D” that were acquired during the calibration time interval are processed to compute at least one algorithm compensation that causes (i) a first result related to the optical property based on the first and second light portions “A” and “B”, which correspond to the subregion A-B (i.e., the “left” side), to be substantially equal to (ii) a second result related to the optical property based on the third and fourth light portions “C” and “D”, which correspond to the subregion C-D (i.e., the “right” side). The first and second results to which algorithm compensation is applied can be, for example, a left-side SO2 reading and a right-side SO2 reading, respectively, computed according to the “slope” method. Alternatively, the first and second results to which algorithm compensation is applied can be intermediate values, such as the intensity-based slope factor K a , for the left and right sides as would be computed on the way to computing an eventual SO2 end result. Shown by way of example in  FIG. 8A  is a plot  850  of the SO2 results for the left side (SO2 A-B ) and the right side (SO2 C-D ) as would be computed by the slope method in a direct or uncompensated form based on readings taken during the calibration interval. Then, shown in  FIG. 8A  in the plot  851  are the results SO2 A-B  and SO2 C-D  as they appear in compensated form, wherein the algorithms for computing these results have been compensated in a way that forces these values to be equal. 
     Examples of algorithm compensations applied to cause the identical results for the two respective subregions are disclosed further infra with respect to  FIGS. 9A-9C  and  FIG. 10 . For the example of  FIG. 8A , the algorithm compensations are simply represented by the use of primed (′) versions of the result computation algorithms. According to one preferred embodiment, the applied algorithm compensation(s) are selected to relate to at least one non-ideality associated with one or more of the intensity of the optical sources, the sensitivity of the optical detectors, the coupling efficiency of light from the optical sources into the medium, and the coupling efficiency of light from the medium to the optical detectors. For example, one or more correction factors can be applied to change the values of the source intensity/coupling coefficients, detector efficiency/coupling coefficients, and/or phase error coefficients (for PMS implementations) used on the slope method equations such that the results of the slope method equations yield the same result for the two different subregions. Stated differently, the calibration process for a multi-subregion monitoring system according to a preferred embodiment harnesses a presumption that the optical property itself is spatially homogenous throughout the multiple subregions during the calibration interval, and that any differences between readings taken during that calibration interval are attributable to determinable non-idealities in the measurement system. The readings taken during the calibration interval are then used to determine the extent of those non-idealities and to compensate for them during the remainder of the monitoring session. Subsequently, if the multiple localized readings begin to depart from each other during the monitoring interval, those differences are indeed attributed to actual biological fluctuations in the patient (e.g., an ischemic condition in the left or right side of the brain), under a presumption that the non-idealities (or at least particular ratios related to those non-idealities, as described further infra) have remained constant during the post-calibration monitoring interval. 
     At step  808 , subsequent to the calibration process of step  806 , detections of the light portions “A” through “D” proceed throughout the monitoring interval, and the optical property is computed using the detected light in conjunction with the one or more compensation factors computed at step  806 . At step  810 , the resultant optical property is displayed on an output display, as illustrated by the plots  852  showing the SO2 level for the left (A-B) and right (C-D) sides of the brain, respectively. Notably, as described above in relation to step  806 , it is not required that the ultimate result (in this case, SO2) be computed for each of the different subregions in determining the algorithm compensations during the calibration phase. Rather, it can be an intermediate result that is computed for each subregion (such as a slope factor), or some other property for each subregion for which homogeneity among subregions would be implicated under an assumption that the ultimate property to be measured is known to be homogeneous throughout the subregions. 
     For one preferred embodiment, during each of the calibration interval and monitoring interval, each of the light portions “A” through “D” comprises a combination of light portions corresponding to two (or more) different wavelengths (e.g., 680 nm and 830 nm), wherein only a single source is emitting at any particular instant in time, and that emitting source is emitting only a single wavelength at any particular instant in time. The different sources and wavelengths are individually cycled through on a repeated basis through successive periods that are termed herein acquisition intervals. By way of example, for an exemplary acquisition interval of one second, the following sequence may be carried: S 1  emitting at 680 nm for 0.25 seconds to provide light portions A(680) and C(680), followed by S 1  emitting at 830 nm for 0.25 seconds to provide light portions A(830) and C(830), followed by S 2  emitting at 680 nm for 0.25 seconds to provide light portions B(680) and D(680), followed by S 2  emitting at 830 nm for 0.25 seconds to provide light portions B(830) and D(830). The process then repeats every second throughout the calibration and monitoring intervals. Any particular light portion at any particular wavelength thereby only has an active duty cycle of 25% (0.25 seconds out of every second). 
     For another preferred embodiment similar to one or more preferred embodiments detailed further hereinbelow in relation to  FIGS. 12-17 , each of the light portions “A” through “D” comprises a combination of light portions corresponding to two (or more) different wavelengths (e.g., 680 nm and 830 nm), wherein both sources are emitting simultaneously and continuously at both wavelengths, and wherein a frequency division multiplexing scheme is used so that the detectors can individually detect each distinct light portion at each distinct wavelength. By way of example, source S 1  may be continuously emitting at 680 nm at a modulation frequency of 155.001 MHz, source S 1  may be continuously emitting at 830 nm at a modulation frequency of 155.002 MHz, source S 2  may be continuously emitting at 680 nm at a modulation frequency of 155.003 MHz, and source S 2  may be continuously emitting at 830 nm at a modulation frequency of 155.004 MHz. Each of the detectors D 1  and D 2  receives all signals simultaneously and separates (demultiplexes) them from each other based on their distinct modulation frequencies. Any particular light portion at any particular wavelength thereby has an active duty cycle of 100% which, as described further hereinbelow, can provide for enhanced signal to noise performance as compared to scenarios in which the there is a lesser duty cycle. 
       FIGS. 9A-9C  and  FIG. 10  illustrate a particular application of the general method of  FIG. 8A , in the context of a PMS-based spectrophotometry system using the probe patch  754  based on two representative wavelengths of 680 nm and 830 nm.  FIGS. 9A-9C  illustrate equations for adapting the slope method of absorption coefficient and effective scattering coefficient computation to a bilateral NIR cerebral oxygenation monitor using a reduced-element across-the-forehead (ATF) patch according to a preferred embodiment.  FIG. 9A  illustrates equations that represent the measured slopes K a  and K p  as “seen” by the left side detector D 1  for the distance interval r 11  to r 21 , which are denoted K a,LEFT (t) and K p,LEFT (t), respectively. The left-side measured slope K a,LEFT (t) is computed from the measured light intensity values I 21 (t) and I 11 (t) as shown, while the measured left-side phase slope K p,LEFT (t) is computed from the measured phase values φ 21 (t) and φ 11 (t) as shown.  FIG. 9B  illustrates equivalent equations applicable for the right side detector D 2 . 
     As illustrated in  FIG. 9C , which collects and compares the slope equations from  FIGS. 9A-9B , the measured left-side slope K a,LEFT (t) differs from the ideal left-side slope K′ a,LEFT (t) only by the log of the ratio of the coupling efficiencies of the centrally located sources S 1  and S 2 , termed herein a source intensity and coupling coefficient ratio factor (SICCRF), divided by the known quantity r 21 -r 11  {Eq. 9C-5}. The measured right-side slope K a,RIGHT (t) differs from the ideal right-side slope K′ a,RIGHT (t) only by the SICCRF (oppositely signed), divided by the known quantity r 12 -r 22  {Eq. 9C-6}. Moreover, the measured left-side phase slope K p,LEFT (t) differs from the ideal left-side phase slope K′ p,LEFT (t) only by the difference of the phase errors of the centrally located sources S 1  and S 2 , termed herein a source phase error factor (SPEF), divided by the known quantity r 21 -r 11  {Eq. 9C-7}. The measured right-side phase slope K p,RIGHT (t) differs from the ideal right-side phase slope K′ p,RIGHT (t) only by the SPEF (oppositely signed), divided by the known quantity r 12 -r 22  {Eq. 9C-8}. According to a preferred embodiment, these relationships are uniquely combined with the bilaterality assumptions set forth above (including homogeneity at time 0) to permit the separate computation of K′ a,LEFT (t), K′ a,RIGHT (t), K′ p,LEFT (t), and K′ p,RIGHT (t) throughout the monitoring session, which are then used to compute separate, absolute left-side (SO2 LEFT (t)) and right-ride (SO2 RIGHT (t)) oxygen saturation values throughout the monitoring session. Briefly stated, when the user presses the “Start” button at the beginning (t=0) of the monitoring session, the algorithm compensation referenced at step  806  of  FIG. 8A  proceeds by a determination of the values for SICCRF and SPEF (calibrated) for each NIR radiation wavelength based on (i) measured intensity and phase values at t=0, and (ii) the assumption that K′ a,LEFT (0)=K′ a,RIGHT (0) and K′ p,LEFT (0)=K′ p,RIGHT (0). Then, for all times t&gt;0 after the calibration is complete, the values of K′ a,LEFT (t), K′ a,RIGHT (t), K′ p,LEFT (t), and K′ p,RIGHT (t) are computed based on (i) the measured intensity and phase values at time “t”, and (ii) the determined (calibrated) values of SICCRF and SPEF. 
     Stated somewhat more broadly, operation of a bilateral NIR cerebral oximeter using a reduced-element ATF patch according to one preferred embodiment is based on a modified version of the slope method in which left-side slopes and right-side slopes are individually computed, wherein (i) at the quiescent beginning of the monitoring session, it is presumed that any differences in the left-side slopes versus the right-side slopes are attributable to coupling efficiency and/or phase error differences among the sources and detectors because the SO2 distribution is assumed uniform across both left and right hemispheres, and (ii) during the subsequent course of the monitoring session, it is presumed that any change in the left-side slopes or right-side slopes is attributable to timewise physical changes in the SO2 values in that hemispheres because the coupling efficiency and/or phase error differences are presumed to be fixed in time. 
     Notably, for the converse preferred embodiment in which the detectors D 1 -D 2  are centrally located and the sources S 1 -S 2  are at the left and right ends, it can be shown that the equations turn out similarly to  FIG. 9C  except that the source intensity and coupling coefficient ratio factor (SICCRF) becomes a detector sensitivity and coupling coefficient ratio factor (DSCCRF) equal to the log of the ratio of the coupling efficiencies of the centrally located detectors D 1  and D 2 , and the source phase error factor (SPEF) becomes a detector phase error factor (DPEF) equal to the difference of the phase errors of the centrally located detectors D 1  and D 2 . Thus, in the a more general expression of the preferred embodiments, the SICCRF could be replaced in the present description by a factor termed the centrally located element coupling coefficient ratio factor (CLECCRF) and the SPEF could be replaced in the present description by a factor termed the centrally located element phase error factor (CLEPEF). 
       FIG. 10  illustrates bilateral NIR cerebral oxygenation level monitoring according to a preferred embodiment. As the process begins at step  1002 , the ATF patch has been mounted and the system has begun to acquire intensity and phase measurements during a calibration interval (the time is arbitrarily set to “0” for the time at which calibration, i.e., algorithm compensation, takes place). A set of quiescent readings for the measured intensities and measured phases is established and maintained at this time, based for example on a running 10-second averaging interval (or other suitable averaging interval) to ensure a set of smooth and reliable intensity and phase values at t=0 when the calibration process will begin. Then, with the patient in a quiescent state such that the bilateral assumptions supra are valid (e.g. the surgery operation has not yet begun and the ATF patch is safely secured to the forehead), the user presses the start button (step  1004 ) at time t=0 to start the calibration process, which is carried out separately for each wavelength. At steps  1008 - 1014 , the measured slopes K a,LEFT (0), K a,RIGHT (0), K p,LEFT (0), and K p,RIGHT (0) are computed from the quiescent measured intensities and phases I 11 (0), φ 11 (0), I 12 (0), φ 12 (0), I 21 (0), φ 21 (0), I 22 (0), and φ 22 (0). At steps  1016 - 1018 , the SICCRF and SPEF are computed based on (i) the measured slopes K a,LEFT (0), K a,RIGHT (0), K p,LEFT (0), and K p,RIGHT (0), and (ii) the assumptions that K′ a,LEFT (0)=K′ a,RIGHT (0) and K′ p,LEFT (0)=K′ p,RIGHT (0). The calibration process for that wavelength is then complete (step  1020 ), and the process is repeated for each wavelength such that separate values of SICCRF and SPEF are established for each wavelength. 
     Subsequent to the calibration process, for all times t&gt;0 (it can be assumed for purposes of this description that the calibration process took a negligible amount of time immediately after t=0), the known (calibrated) values of SICCRF and SPEF are used in conjunction with the ongoing measured slope values to compute the ideal slope values for the left side, right side, and whole-brain for each wavelength, which are then used as the basis for the left side, right side, and whole-brain SO2 values. Thus, at step  1024 , the measured slope values K a,LEFT (t), K a,RIGHT (t), K p,LEFT (t), and K p,RIGHT (t) are computed from the measured intensities and phases at time “t”. At step  1026 , the ideal slope values K′ a,LEFT (t), K′ a,RIGHT (t), K′ p,LEFT (t), and K′ p,RIGHT (t) are computed based on K a,LEFT (t), K a,RIGHT (t), K p,LEFT (t), and K p,RIGHT (t) and the values of SICCRF and SPEF. At step  1028 , the absorption coefficients and effective scattering coefficients are computed from K′ a,LEFT (t), K′ a,RIGHT (t), K′ p,LEFT (t), and K′ p,RIGHT (t). For whole-brain monitoring, the value of K′ a,WHOLE (t) is computed as the average of K′ a,LEFT (t) and K′ a,RIGHT (t), the value of K′ p,WHOLE (t) is computed as the average of K′ p,LEFT (t) and K′ p,RIGHT (t), and the corresponding absorption coefficients and effective scattering coefficients are computed therefrom at step  1029 . Finally, at steps  1030 - 1033  the values of SO2 LEFT (t), SO2 RIGHT (t), and SO2 WHOLE (t) are computed from the absorption coefficients at the multiple wavelengths, and at step  1034  they are displayed on the output display  716 . 
       FIG. 11  illustrates NIR probe patches  1104  and  1105  mounted on the forehead of a patient and a corresponding source timing diagram corresponding to a scenario in which there is a relatively low duty cycle for any particular optical signal at any particular wavelength. Probe patch  1104  includes two source ports S 1  and S 2  and four detector ports D. The probe patch  1105  also includes two source ports S 3  and S 4  and four detector ports. For the example of  FIG. 11 , it is presumed that a PMS (phase modulated spectrophotometry) scheme is used in which there are two NIR wavelengths (680 nm and 830 nm) and a modulation frequency of 155 MHz. During each acquisition cycle T A , which is typically on the order of 1 second, there needs to be provided individually measured amplitudes and phases for each of the individual wavelengths 680 nm and 830 nm for each individual source port/detector port pair on each of the NIR probe patches  1104  and  1105 . According to the example of  FIG. 11 , this is achieved by firing each source port/wavelength pair during a distinct time interval that does not overlap with any other source port/wavelength pair. Thus, each of the following source port/wavelength pairs emits during a distinct time interval: S 1 —680 nm; S 1 —830 nm; S 2 —680 nm; S 2 —830 nm; S 3 —680 nm; S 3 —830 nm; S 4 —680 nm; and S 4 —830 nm. Each detector actively detects (“listens”) whenever any of the source ports on that same NIR probe patch are firing. 
     By firing each source port/wavelength pair during a distinct time interval, it is ensured that each detector port achieves a clear, individualized “channel” with each source port/wavelength pair (i.e., with each individual wavelength emitted at each individual source port), without interference or stray radiation from other sources or other wavelengths. As used herein, “duty cycle” refers to the percentage of time that any particular “channel” (i.e., any particular source port/detector port/wavelength triplet) is actively providing measured amplitudes and phases during the tissue monitoring session. It can be readily seen that all detector ports will have the same duty cycle for any particular source port/wavelength pair, because the detector ports can operate (“listen”) independently of each other. Accordingly, unless indicated otherwise, duty cycles are presented herein only in terms of the particular source port/wavelength pair (e.g., the duty cycle for S 1 —680 nm, the duty cycle for S 1 —830 nm, etc.), with it being understood that such duty cycle applies across all of the different detectors on the NIR probe patch. 
     For the example of  FIG. 11 , assuming that all source port/wavelength pairs are given equal treatment, the maximum achievable duty cycle is 12.5% for each source port/wavelength pair. More generally, for systems having “N” different source ports and “M” different wavelengths, the maximum achievable duty cycle is 1/(NM). Most prior art cerebral oximetry systems exhibit duty cycles that are well below the maximum achievable duty cycle due to various hardware considerations, such as detection “setup time” for synchronizing to the next active source port/wavelength pair. Some known prior art cerebral oximetry systems exhibit duty cycles that are even as low as 1% for each source port/wavelength pair. 
       FIG. 12  illustrates NIR probe patches  1204  and  1205  mounted on the forehead of a patient and a corresponding source timing diagram for an NIR cerebral oximetry system according to a preferred embodiment, wherein each source port/wavelength pair emits at a different modulation frequency, and wherein an overall received signal at each detector port is processed to separately extract therefrom a plurality of individual received signals based on their different modulation frequencies, each individual received signal corresponding to a respective one of the source port/wavelength pairs. This provides the ability for multiple source port/wavelength pairs to be emitting simultaneously, because each detector is able to distinguish each individual “channel” based on its modulation frequency. For one preferred embodiment, for a system having “N” different source ports and “M” different wavelengths, all “NM” source port/wavelength pairs emit simultaneously and continuously throughout the monitoring session, each having a different modulation frequency, thereby providing 100% duty cycle (“full duty cycle”). 
     By providing full duty cycle for each individual source port/detector port/wavelength triplet (“channel”) in the preferred embodiment of  FIG. 12 , as compared to a duty cycle of 1/(NM) for each such channel in the example of  FIG. 11 , there is provided a factor of NM more data points over any particular sampling period for each channel. This, in turn, provides for an increase in the signal to noise ratio (SNR) for each channel. It can be shown that the improvement in signal to noise performance for each channel can be estimated by the square root of the factor by which the number of data points per sampling interval has increased. Thus, where the number of data points per sampling interval has increased by a factor of NM, the improvement in signal to noise performance is roughly the square root of NM. For the preferred embodiment of  FIG. 12 , where the number of source ports “N” is 4, the number of wavelengths “M” is 2, and the number “NM” of source port/wavelength pairs is 8, the improvement in signal to noise performance over the example of  FIG. 11  is roughly 280%. Another advantage is that, since the monitoring is continuous, there are no inefficiencies caused by the need for repeated “setup times” during each acquisition interval, as is the case for the example of  FIG. 11 . 
       FIG. 13  illustrates a block diagram of signal processing circuitry associated with NIRS monitoring according to a preferred embodiment. The block diagram of  FIG. 13  is individually applicable for each distinct detector port (detector). The overall received signal at a particular detector port, which is in analog electrical form (e.g., as the output of a photomultiplier tube or semiconductor photodiode), is processed to separately extract therefrom a plurality of individual received signals (amplitudes I and phases φ) based on their different modulation frequencies. As illustrated, each individual received signal (e.g., I S1-680 , φ S1-680 ) corresponds to a respective one of the source port/wavelength pairs (e.g., S 1 —680 nm) with which that detector port establishes NM/2 corresponding respective source port/detector port/wavelength triplets (“channels”). For the more general case in which all sources and detectors are on the same probe patch, the number of source port/detector port/wavelength triplets (“channels”) established for each detector port is NM. For one preferred embodiment, the circuit of  FIG. 13  is replicated for each different detector port in the NIRS monitoring system. 
     It is to be appreciated that the particular modulation frequencies, channel spacings, etc. that are set forth  FIGS. 12-13  are presented by way of example only, and not by way of limitation, although for one preferred embodiment, it has been found advantageous from a hardware and signal processing perspective to use channel spacings (here, 1 kHz) that are relatively low compared to the “base” modulation frequency (here, 155 MHz) so that only one analog mixer is needed and so that the baseband digital processing (FFT, channel filtering) can be implemented with relatively inexpensive hardware. More generally, for one preferred embodiment it has been found useful in PMS-based systems to have the different modulation frequencies each be greater than 100 MHz and yet differ from each other by less than 100 kHz. For another preferred embodiment it has been found useful in PMS-based systems to have the different modulation frequencies each be greater than 100 MHz and yet differ from each other by less than 1 MHz. 
       FIG. 14  illustrates an across-the-forehead (ATF) NIR probe patch  1402  mounted on the forehead of a patient and a corresponding source timing diagram for an NIR cerebral oximetry system according to a preferred embodiment. The ATF patch  1402  is preferably similar to that described in Ser. No. 12/826,218, supra, which is incorporated by reference herein, the ATF patch  1402  being particularly advantageous in providing bilateral monitoring of cerebral oxygenation levels using a reduced number of source/detector elements and/or in providing localized optical property readings without requiring certain source-detector symmetries. It has been found particularly advantageous to use the full-duty cycle methods of  FIGS. 12-13 , supra, in combination with the teachings of Ser. No. 12/826,218, supra, at least due to the large source-detector separation distances S 1 -D 2  and S 2 -D 1  that can be realized, which can be up to 6 cm or even greater, and whose operation can benefit greatly from the improved signal to noise performance provided by, according to a preferred embodiment, having the source port/wavelength pairs (S 1 —680 nm, S 1 —830 nm, S 2 —680 nm, and S 2 —830 nm) all emitting simultaneously at different modulation frequencies (155.001 MHz, 155.002 MHz, 155.003 MHz, and 155.004 MHz, respectively), and processing the overall received signal at each detector port to individually extract therefrom the received signals corresponding to each respective source port/wavelength pair. 
     Provided in conjunction with each of the preferred embodiments is a console unit coupled via optical, electro-optical, or electrical cables to the NIR probe patch and comprising one or more optical sources and optical detectors, each of which may be fully optical, electro-optical, or fully electrical in nature depending on the nature of the sources and detectors on the NIR probe patch. For one preferred embodiment, the optical sources comprise one or more laser sources, the optical detectors comprise one or more photomultiplier tubes (PMTs), and the NIR probe patch consists of passive optical sources and detectors and has a general overall construction similar to one or more of the NIR probe patches disclosed in the commonly assigned U.S. Ser. No. 12/483,610 filed Jun. 12, 2009, which is incorporated by reference herein, except that the dimensions, source locations, and detector locations are as set forth herein and/or in Ser. No. 12/826,218, supra. The console unit further comprises a processor and analog/digital hardware coupled to control and receive information from the optical sources and optical detectors, the processor and analog/digital hardware being configured, dimensioned, and programmed to achieve the functionalities described herein. The console unit further comprises an output display coupled to the processor that displays the SO2 readings in real time. 
     Each of the source ports on the NIR probe patch can be optically coupled to the optical sources of the console unit so as to simultaneously emit optical signals at each of the different wavelengths (e.g., 680 nm and 830 nm) being used. Alternatively, each source port can be spatially divided into multiple sub-ports, each sub-port simultaneously emitting at a different wavelength (for example, the source port/wavelength pair S 1 —680 nm being provided at a first sub-port of source port S 1  and the source port/wavelength pair S 1 —830 nm being provided at a second sub-port of source port S 1 ). 
       FIGS. 15A-15B  illustrate an NIR probe patch  1502  having curved-edge detection apertures according to a preferred embodiment. The NIR probe patch  1502  is wearable on a surface of the tissue volume of the patient and comprises a source S 1 , a first detector D 1 , and a second detector D 2 . The probe patch  1502  is configured to maintain the source S 1 , the first detector D 1 , and the second detector D 2  in secured contact with the surface of the tissue volume throughout the patient monitoring session. The detector D 1  includes an aperture  1504  formed in the tissue-facing surface of the probe patch  1502 , wherein light must travel through the aperture  1504  in order to be detected. The detector D 2  likewise includes an aperture  1506 . With reference to the more detailed drawing of the aperture  1504  in  FIG. 15B , the aperture comprises a central area  1504 C, a nearer edge  1504   i  that is nearer to the source S 1  than the central area  1504 C, and a farther edge  1504   o  that is farther from the source S 1  than the central area  1504 C. The aperture  1504  further comprises side edges  1504   x  and  1504   y.  According to a preferred embodiment, the nearer edge  1504   i  and farther edge  1504   o  are each curved inward toward the location of the source S 1  (which can also be termed a source port location). In one preferred embodiment, the side edges  504   x  and  504   y  each correspond to a radial line extending through the source port location. In one preferred embodiment, the aperture  1504  has an arcuate slit-like character in that it is relatively narrow in a depthwise direction from the source S 1  (i.e., in the direction of a radial line extending from the source S 1 ) and relatively long in a tangential direction (i.e., in the direction of a tangent to a radial line extending from the source S 1 ). For the example of  FIG. 15B , the aperture  1504  is roughly three times as long in the tangential direction than in the depthwise direction. In other preferred embodiments, the aperture  1504  is at least five times as long in the tangential direction than in the depthwise direction, making its overall shape even more slit-like. 
     In one preferred embodiment, the nearer edge  1504   i  has a generally constant radius of curvature r i  and the farther edge  1504   o  has a generally constant radius of curvature r o , wherein each of the curvatures r i  and r o  is equal to an average distance r 1  of the aperture  1504  from the source S 1 . Likewise, the nearer and farther edges of aperture  1506  each have a curvature radius equal to an average distance r 2  of the aperture  1506  from the source S 1 . In another preferred embodiment, the curvatures r i  and r o  of aperture  1504  are equal to 0.5 times r 1 , or are equal to another fixed percentage of r 1 , which can be empirically tuned. 
       FIG. 16  illustrates the use of curved apertures similar to those of  FIGS. 15A-15B  in conjunction with an across-the-forehead (ATF) NIR probe patch  1602  that is otherwise similar to the ATF patch  1402  of  FIG. 14 , supra. For one preferred embodiment, the radius of curvature of each of the nearer and farther edges of each detection port (detection aperture) is the average of the distances to the different source ports S 1  and S 2 . Thus, for example, the radius of curvature of each of the nearer and farther edges of detection aperture D 3  is equal to (r 13 +r 23 )/2. It has been found that using curved apertures as shown in  FIGS. 15A-15B  and  FIG. 16  can reduce phase measurement error and/or provide more precise phase measurements. 
       FIG. 17  illustrates NIRS monitoring of a tissue volume of a patient according to a preferred embodiment. At step  1702 , a probe patch containing a plurality of optical sources and a plurality of optical detectors is secured to a surface of the tissue volume. Illustrated in  FIG. 17  is a probe patch  1752  as secured to the forehead of a patient for cerebral oximetry. The probe patch is optically, electrically, or electrooptically coupled to a console unit  1754  that includes a source controller, a detector controller, at least one processor, and a display that are collectively configured and/or programmed to achieve the functionalities described herein. Preferably, as illustrated in  FIG. 17 , the probe patch  1752  includes curved-aperture detectors D 1  and D 2  similar to those described above with respect to  FIGS. 15A ,  FIG. 15B , and/or  FIG. 16 . 
     At step  1704 , the plurality of optical sources are operated to introduce, simultaneously and on a continuous basis throughout the patient monitoring session, a plurality of optical signals into the tissue volume, wherein each of the optical signals has a modulation frequency different than that of each other optical signal, and wherein any two of the optical signals that are introduced from a same one of the optical sources are at different optical wavelengths. Illustrated by way of example in  FIG. 17  is a side view of the probe patch  1752  as secured to a surface  1756  of a tissue volume, wherein first and second optical signals (OS 1  and OS 2 ) are introduced into the tissue volume from source S 1 , and wherein third and fourth optical signals (OS 3  and OS 4 ) are introduced into the tissue volume from source S 2 , all signals being introduced simultaneously and continuously throughout the patient monitoring session. The optical signals OS 1 , OS 2 , OS 3 , and OS 4  are all at different modulation frequencies, and any two of them emitted from a common optical source are at different optical wavelengths. 
     At step  1706 , the plurality of optical detectors are operated to detect, simultaneously and on a continuous basis throughout the monitoring session, a portion of each of the optical signals that has propagated thereto, and the detected signal portions are processed to derive an amplitude signal and a phase signal associated therewith. Thus, for example, the first optical signal OS 1  as introduced into the tissue volume by source S 1  will have a first optical signal portion OSP 11  that propagates to the detector D 1 , and will have a second optical signal portion OSP 12  that propagates to the detector D 2 . Each detector will receive a portion of each of the optical signals OS 1 , OS 2 , OS 3 , and OS 4  that has propagated thereto. For example, detector D 1  will receive the optical signal portions OSP 11 , OSP 21 , OSP 31 , and OSP 41 , while detector D 2  will receive the optical signal portions OSP 12 , OSP 22 , OSP 32 , and OSP 42 . Each detector will generate a first signal representative of an overall combination of the optical signal portions as received at that detector. For example, detector D 1  will generate an overall signal O 1  representative of the combination of the optical signal portions OSP 11 , OSP 21 , OSP 31 , and OSP 41  received thereat. 
     Based on the different modulation frequencies of the optical signal portions OSP 11 , OSP 21 , OSP 31 , and OSP 41  and using a circuit similar or analogous to that of  FIG. 13 , the console unit  1754  will demultiplex the first signal O 1  into individual components corresponding to the optical signal portions OSP 11 , OSP 21 , OSP 31 , and OSP 41 , and then process these individual components to generate an amplitude signal and a phase signal associated with each of the optical signal portions. Likewise, based on the different modulation frequencies of the optical signal portions OSP 12 , OSP 22 , OSP 32 , and OSP 42  and using a circuit similar or analogous to that of  FIG. 13 , the console unit  1754  will demultiplex the signal O 2  into individual components corresponding to the optical signal portions OSP 12 , OSP 22 , OSP 32 , and OSP 42 , and then process these individual components to generate an amplitude signal and a phase signal associated with each of the optical signal portions. 
     Finally, at step  1708 , the amplitude signals and a phase signals associated with the detected optical signal portions are processed to determine the optical property of the tissue volume. For one preferred embodiment, with reference generally to  FIG. 9C-1  through  FIG. 10 , supra, the optical property can be computed at step  1708  according to the steps of: (i) for a nearer-spaced source-detector pair selected from the pluralities of optical sources and detectors, receiving the amplitude signals and phase signals for two corresponding optical signal portions having distinct wavelengths; (ii) for a farther-spaced source-detector pair selected from the pluralities of optical sources and detectors and including either the optical source or the optical detector of the nearer-spaced source-detector pair, receiving the amplitude signals and phase signals for two corresponding optical signal portions having distinct wavelengths; and (iii) processing the amplitude signals and phase signals corresponding to the nearer-spaced and farther-spaced source-detector pairs according to a slope-based phase modulation spectroscopy (PMS) algorithm to compute an absorption property and a scattering property relevant to at least a portion of the tissue volume. 
     Whereas many alterations and modifications of the present invention will no doubt become apparent to a person of ordinary skill in the art after having read the foregoing description, it is to be understood that the particular embodiments shown and described by way of illustration are in no way intended to be considered limiting. By way of example, although 100% or “full” duty cycle operation is particularly advantageous in the context of PMS (phase modulated spectrophotometry) systems, the scope of the preferred embodiments can also include CWS (continuous wave spectrophotometry) systems. For CWS schemes, even though phases are not measured, there is usually some modulation of the NIR signals performed to avoid 1/f effects, with typical modulation frequencies being on the order of 25 kHz. For these cases, the different source port/wavelength pairs can simply be modulated at distinct frequencies of 25 kHz, 26 kHz, 27 kHz, etc., with the overall received signal at each detector being separated into individual received signals based on these different frequencies. 
     By way of further example, although 100% or “full” duty cycle provides the most increase in signal to noise performance, it is also within the scope of the preferred embodiments to provide a less-than-full duty cycle system in which more than one, but fewer than all, of the “NM” different source port/wavelength pairs are emitting simultaneously. For example, a first half of the source port/wavelength pairs can simultaneously emit only during the first half of the acquisition cycle T A , and a second half of the source port/wavelength pairs can simultaneously emit only during the second half of the acquisition cycle T A . Such less-than-full duty cycle strategies could provide for relaxed demodulation/filtering hardware requirements and/or improved channel separation, while still providing for appreciably significant increases in signal to noise performance over the example of  FIG. 11 . 
     By way of still further example, one or more of the preferred embodiments supra are readily applicable for improving the signal to noise performance of NIRS monitoring systems that employ more than one “base” modulation frequency. The preferred embodiment of  FIG. 12 , for example, involved a single “base” modulation frequency of 155 MHz. Some known proposals for NIR spectrophotometric monitoring, however, are based on the use of multiple “base” modulation frequencies, such as that disclosed in the commonly assigned U.S. Pat. No. 7,551,950, in which two modulation frequencies (120 MHz and 150 MHz) are used in one of its examples. In such case, there is provided in accordance with one preferred embodiment a 50% duty cycle system, wherein all source port/wavelength pairs simultaneously emit at distinct frequencies around 120 MHz (e.g., 120.001 MHz, 120.002 MHz, 120.003 MHz, etc.) during the first half of the acquisition cycle, and then all source port/wavelength pairs simultaneously emit at distinct frequencies around 150 MHz (e.g., 150.001 MHz, 150.002 MHz, 150.003 MHz, etc.) during the second half of the acquisition cycle. Alternatively, there is provided in accordance with another preferred embodiment a 100% duty cycle system in which all source port/wavelength pairs emit simultaneously at all modulation frequencies (e.g., 120.001 MHz, 120.002 MHz, 120.003 MHz . . . , 150.001 MHz, 150.002 MHz, 150.003 MHz, etc.). Accordingly, it can be readily seen that the preferred embodiments are applicable across a wide variety of different NIRS implementations. Among other advantages, the improved signal to noise performance provided according to one or more preferred embodiments provides an ability to increase penetration depths in the non-invasive NIRS monitoring of crucial deep-layer tissue structures such as the human brain. 
     By way of even further example, in one preferred embodiment an NIR cerebral oximetry system is provided using the full-duty-cycle aspects and/or the curved aperture-shape aspects of one or more preferred embodiments supra in conjunction with the deep-layer-specific monitoring methods of the commonly assigned U.S. Ser. No. 12/815,696, supra. By way of even further example, there can be provided in an alternative preferred embodiment a scenario in which a same source is emitting at two different wavelengths simultaneously, wherein the modulation frequency for the two wavelengths is also identical. For this case, each detector port can be provided with a wavelength separation filter (e.g., a filter that passes light at 680 nm and reflects light at 830 nm) that separates the two optical signals based on optical wavelength, and then proceeds to separately demodulate those two optical signals. Therefore, reference to the details of the embodiments are not intended to limit their scope, which is limited only by the scope of the claims set forth below.