`
`A PULSE OXIMETER AND A METHOD OF ITS OPERATION
`This application is a divisional application of U.S. patent application Ser. No. 09/939,391 filed
`Aug. 24, 2001, now abandoned.
`BACKGROUND OF THE INVENTION
`1. Field of the Invention
`This invention is generally in the field of pulse oximetry, and relates to a sensor for use in a pulse
`oximeter, and a method for the pulse oximeter operation.
`2. Background of the Invention
`Oximetry is based on spectrophotometric measurements of changes in the color of blood,
`enabling the. non-invasive determination of oxygen saturation in the patient's blood. Generally,
`oximetry is based on the fact that the optical property of blood in the visible (between 500 and
`700nm700 nm) and near-infrared (between 700 and lOOOnm1000 nm) spectra depends strongly
`on the amount of oxygen in blood.
`Referring to FigFIG. 1, there is illustrated a hemoglobin spectra measured by oximetry based
`techniques. Graphs Gl G1 and G2 correspond, respectively, to reduced hemoglobin, or
`deoxyhemoglobin (Hb), and oxygenated hemoglobin, or oxyhemoglobin (HbO2), spectra. As
`shown, deoxyhemoglobin (Hb) has a higher optical extinction (i.e., absorbs more light) in the red
`region of spectrum around 660nm660 nm, as compared to that of oxyhemoglobin (HbO2). On the
`other hand, in the near-infrared region of the spectrum around 940nm940 nm, the optical
`absorption by deoxyhemoglobin (Hb) is lower than the optical absorption of oxyhemoglobin
`(HbO2).
`Prior art non-invasive optical sensors for measuring arterial oxyhemoglobin saturation (SaO2) by
`a pulse oximeter (termed SpO2) are typically comprised of a pair of small and inexpensive light
`emitting diodes (LEDs), and a single highly sensitive silicon photodetector. A red (R) LED
`centered on a peak emission wavelength around 660nm660 nm and an infrared (IR) LED
`centered on a peak emission wavelength around 940nm940 nm are used as light sources.
`Pulse oximetry relies on the detection of a photoplethysmographic signal caused by variations in
`the quantity of arterial blood associated with periodic contraction and relaxation of a patient'
`spatient's heart. The magnitude of this signal depends on the amount of blood ejected from the
`heart into the peripheral vascular bed with each systolic cycle, the optical absorption of the
`blood, absorption by skin and tissue components, and the specific wavelengths that are used to
`illuminate the tissue. SaO2 is determined by computing the relative magnitudes of the R and IR
`photoplethysmograms. Electronic circuits inside the pulse oximeter separate the R and IR
`photoplethysmograms into their respective pulsatile (AC) and non-pulsatile (DC) signal
`components. An algorithm inside the pulse oximeter performs a mathematical normalization by
`which the time- varying AC signal at each wavelength is divided by the corresponding time-
`invariant DC component which results mainly from the light absorbed and scattered by the
`bloodless tissue, residual arterial blood when the heart is in diastole, venous blood and skin
`pigmentation.
`Since it is assumed that the AC portion results only from the arterial blood component, this
`scaling process provides a normalized R/IR ratio (i.e., the ratio of AC/DC values corresponding
`to R- and IR-spectrum wavelengths, respectively), which is highly dependent on SaO2, but is
`largely independent of the volume of arterial blood entering the tissue during systole, skin
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`pigmentation, skin thickness and vascular structure. Hence, the instrument does not need to be
`re-calibrated for measurements on different patients. Typical calibration of a pulse oximeter is
`illustrated in FigFIG. 2 by presenting the empirical relationship between SaO2 and the
`normalized R/IR ratio, which is programmed by the pulse oximeters' manufacturers.
`Pulse oximeters are of two kinds operating, respectively, in transmission and reflection modes. In
`transmission-mode pulse oximetry, an optical sensor for measuring SaO2 is usually attached
`across a fingertip, foot or earlobe, such that the tissue is sandwiched between the light source and
`the photodetector. In reflection-mode or backscatter type pulse oximetry, as shown in Fig. 3, the
`In reflection-mode or backscatter type pulse oximetry, as shown in FIG. 3, the LEDs and
`photodetector are both mounted side-by-side next to each other on the same planar substrate.
`This arrangement allows for measuring SaO2 from multiple convenient locations on the body
`(e.g. the head, torso, or upper limbs), where conventional transmission-mode measurements are
`not feasible. For this reason, non-invasive reflectance pulse oximetry has recently become an
`important new clinical technique with potential benefits in fetal and neonatal monitoring. Using
`reflectance oximetry to monitor SaO2 in the fetus during labor, where the only accessible location
`is the fetal scalp or cheeks, or on the chest in infants with low peripheral perfusion, provides
`several more convenient locations for sensor attachment.
`Reflection pulse oximetry, while being based on similar spectrophotometric principles as the
`transmission one, is more challenging to perform and has unique problems that can not always be
`solved by solutions suitable for solving the problems associated with the transmission-mode
`pulse oximetry. Generally, comparing transmission and reflection pulse oximetry, the problems
`associated with reflection pulse oximetry consist of the following:
`In reflection pulse oximetry, the pulsatile AC signals are generally very small and, depending on
`sensor configuration and placement, have larger DC components as compared to those of
`transmission pulse oximetry. As illustrated in FigFIG. 4, in addition to the optical absorption and
`reflection due to blood, the DC signal of the R and IR photoplethysmograms in reflection pulse
`oximetry can be adversely affected by strong reflections from a bone. This problem becomes
`more apparent when applying measurements at such body locations as the forehead and the
`scalp, or when the sensor is mounted on the chest over the ribcage. Similarly, variations in
`contact pressure between the sensor and the skin can cause larger errors in reflection pulse
`oximetry (as compared to transmission pulse oximetry) since some of the blood near the
`superficial layers of the skin may be normally displaced away from the sensor housing towards
`deeper subcutaneous structures. Consequently, the highly reflective bloodless tissue
`compartment near the surface of the skin can cause large errors even at body locations where the
`bone is located too far away to influence the incident light generated by the sensor.
`Another problem with currently available reflectance sensors is the potential for specular
`reflection caused by the superficial layers of the skin, when an air gap exists between the sensor
`and the skin, or by direct shunting of light between the LEDs and the photodetector through a
`thin layer of fluid which may be due to excessive sweating or from amniotic fluid present during
`delivery. ' It is important to keep in mind the two fundamental assumptions underlying the
`conventional dual- wavelength pulse oximetry, which are as follows:
`It is important to keep in mind the two fundamental assumptions underlying the conventional
`dual-wavelength pulse oximetry, which are as follows:
`(1) the path of light rays with different illuminating wavelengths in tissue are substantially equal
`and, therefore, cancel each other; and (2) each light source illuminates the same pulsatile change
`in arterial blood volume.
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`Furthermore, the correlation between optical measurements and tissue absorptions in pulse
`oximetry are based on the fundamental assumption that light propagation is determined primarily
`by absorbanceabsorbable due to LambertLambent-Beer's law neglecting multiple scattering
`effects in biological tissues. In practice, however, the optical paths of different wavelengths in
`biological tissues is known to vary more in reflectance oximetry compared to transmission
`oximetry, since it strongly depends on the light scattering properties of the illuminated tissue and
`sensor mounting.
`Several human validation studies, backed by animal investigations, have suggested that
`uncontrollable physiological and physical parameters can cause large variations in the calibration
`curve of reflectance pulse oximeters primarily at low oxygen saturation values below 70%. It
`was observed that the accuracy of pulse oximeters in clinical use might be adversely affected by
`a number of physiological parameters when measurements are made from sensors attached to the
`forehead, chest, or the buttock area. While the exact sources of these variations are not fully
`understood, it is generally believed that there are a few physiological and anatomical factors that
`may be the major source of these errors. It is also well known for example that changes in the
`ratio of blood to bloodless tissue volumes may occur through venous congestion,
`vasoconstriction/vasodilatation, or through mechanical pressure exerted by the sensor on the
`skin.
`Additionally, the empirically derived calibration curve of a pulse oximeter can be altered by the
`effects of contact pressure exerted by the probe on the skin. This is associated with the following.
`The light paths in reflectance oximetry are not well defined (as compared to transmission
`oximetry), and thus may differ between the red and infrared wavelengths. Furthermore, the
`forehead and scalp areas consist of a relatively thin subcutaneous layer with the cranium bone
`underneath, while the tissue of other anatomical structures, such as the buttock and hmbslimbs,
`consists of a much thicker layer of skin and subcutaneous tissues without a nearby bony support
`that acts as a strong light reflector.
`Several in vivo and in vitro studies have confirmed that uncontrollable physiological and
`physical parameters (e.g., different amounts of contact pressure applied by the sensor on the skin,
`variation in the ratio of bloodless tissue-to-blood content, or site-to-site variations) can often
`cause large errors in the oxygen saturation readings of a pulse oximeter, which are normally
`derived based on a single internally- programmed calibration curve. The relevant in vivo studies
`are disclosed in the following publications:
`1. Dassel, et al., "“Effect of location of the sensor on reflectance pulse oximetry",”, British
`Journal of Obstetrics and Gynecology, vol. 104, pp. 910-916, (1997);
`2. Dassel, et al., 'Reflectance“Reflectance pulse oximetry at the forehead of newborns: The
`influence of varying pressure on the probe",”, Journal of Clinical Monitoring, vol. 12, pp. 421-
`428, (1996).]).
`The relevant in vitro studies are disclosed, for example in the following publication:
`3. Edrich et al., "“Fetal pulse oximetry: influence of tissue blood content and hemoglobin
`concentration in a new in- vitro model",”, European Journal of Obstetrics and Gynecology and
`Reproductive Biology, vol. 72, suppl. 1, pp. S29-S34, (1997).
`Improved sensors for application in dual-wavelength reflectance pulse oximetry have been
`developed. As disclosed in the following publication: Mendelson, et al., "“Noninvasive pulse
`oximetry utilizing skin reflectance photoplethysmography",”, IEEE Transactions on Biomedical
`Engineering, vol. 35, no. 10, pp. 798-805 (1988), the total amount of backscattered light that can
`be detected by a reflectance sensor is directly proportional to the number of photodetectors
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`placed around the LEDs. Additional improvements in signal-to-noise ratio were achieved by
`increasing the active area of the photodetector and optimizing the separation distance between
`the light sources and photodetectors.
`Another approach is based on the use of a sensor having six photodiodes arranged symmetrically
`around the LEDs that is disclosed in the following publications:
`4. Mendelson, et al., "“Design and evaluation of a new reflectance pulse oximeter sensor",”,
`Medical Instrumentation, vol. 22, no. 4, pp. 167-173 (1988); and
`5. Mendelson, et al., "“Skin reflectance pulse oximetry: in vivo measurements from the forearm
`and calf,”, Journal of Clinical Monitoring, vol. 7, pp. 7-12, (1991).
`According to this approach, in order to maximize the fraction of backscattered light collected by
`the sensor, the currents from all six photodiodes are summed electronically by internal circuitry
`in the pulse oximeter. This configuration essentially creates a large area photodetector made of
`six discrete photodiodes connected in parallel to produce a single current that is proportional to
`the amount of light backscattered from the skin. Several studies showed that this sensor
`configuration could be used successfully to accurately measure SaO2 from the forehead, forearm
`and the calf on humans. However, this sensor requires a means for heating the skin in order to
`increase local blood flow, which has practical limitations since it could cause skin burns.
`Yet another prototype reflectance sensor is based on eight dual-wavelength LEDs and a single
`photodiode, and is disclosed in the following publication: Takatani et al., "“Experimental and
`clinical evaluation of a noninvasive reflectance pulse oximeter sensor",”, Journal of Clinical
`Monitoring, vol. 8, pp. 257-266 (1992). Here, four R and four IR LEDs are spaced at 90-degree
`intervals around the substrate and at an equal radial distance from the photodiode.
`A similar sensor configuration based on six photodetectors mounted in the center of the sensor
`around the LEDs is disclosed in the following publication: Konig, et al., "“Reflectance pulse
`oximetry - —principles and obstetric application in the Zurich system",”, Journal of Clinical
`Monitoring, vol. 14, pp. 403-412 (1998).
`According to the techniques disclosed in all of the above publications, only
` LEDs of two wavelengths, R and IR, are used as light sources, and the computation of
` SaO2 is based on reflection photoplethysmograms measured by a single photodetector,
`regardless of whether one or multiple photodiodes chips are used to construct the sensor. This is
`because of the fact that the individual signals from the photodetector elements are all summed
`together electronically inside the pulse oximeter. Furthermore, while a radially-symmetric
`photodetector array can help to maximize the detection of backscattered light from the skin and
`minimize differences from local tissue inhomogeneity, human and animal studies confirmed that
`this configuration can not completely eliminate errors caused by pressure differences and site-to-
`site variations.
`The use of a nominal dual-wavelength pair of 735/890nm890 nm was suggested as providing the
`best choice for optimizing accuracy, as well as sensitivity in dual- wavelength reflectance pulse
`oximetry, in USU.S. Pat. Nos. 5,782,237 and 5,421,329. This approach minimizes the effects of
`tissue heterogeneity and enables to obtain a balance in path length changes arising from
`perturbations in tissue absorbance. This is disclosed in the following publications:
`6. Mannheimer at al., "“Physio-optical considerations in the design of fetal pulse oximetry
`sensors",”, European Journal of Obstetrics and Gynecology and Reproductive Biology, vol. 72,
`suppl. 1, pp. S9-S19, (1997); and
`7. Mannheimer at al., "“Wavelength selection for low-saturation pulse oximetry",”, IEEE
`Transactions on Biomedical Engineering, vol. 44, no. 3, pp. 48-158 (1997)].
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`However, replacing the conventional R wavelength at 660nm660 nm, which coincides with the
`region of the spectrum where the difference between the extinction coefficient of Hb and HbO2 is
`maximal, with a wavelength emitting at 735nm735 nm, not only lowers considerably the overall
`sensitivity of a pulse oximeter, but does not completely eliminate errors due to sensor placement
`and varying contact pressures.
`Pulse oximeter probes of a type comprising three or more LEDs for filtering noise and
`monitoring other functions, such as carboxyhemoglobin or various indicator dyes injected into
`the blood stream, have been developed and are disclosed, for, example, in WO 00/32099 and
`USU.S. Pat. No. 5,842,981. The techniques disclosed in these publications are aimed at
`providing an improved method for direct digital signal formation from input signals produced by
`the sensor and for filtering noise.
`None of the above prior art techniques provides a solution to overcome the most essential
`limitation in reflectance pulse oximetry, which requires the automatic correction of the internal
`calibration curve from which accurate and reproducible oxygen saturation values are derived,
`despite variations in contact pressure or site-to- site tissue heterogeneity.
`In practice, most sensors used in reflection pulse oximetry rely on closely spaced LED
`wavelengths in order to minimize the differences in the optical path lengths of the different
`wavelengths. Nevertheless, within the wavelength range required for oximetry, even closely
`spaced LEDs with closely spaced wavelengths mounted on the same substrate can lead to large
`random error in the final determination of SaO2.
`SUMMARY OF THE INVENTION AND ADVANTAGES
`The object of the invention is to provide a novel sensor design and method that functions to
`correct the calibration relationship of a reflectance pulse oximeter, and reduce measurement
`inaccuracies in general. Another object of the invention is to provide a novel sensor and method
`that functions to correct the calibration relationship of a reflectance pulse oximeter, and reduce
`measurement inaccuracies in the lower range of oxygen saturation values (typically below 70%),
`which is the predominant range in neonatal and fetal applications.
`Yet another object of the present invention is to provide automatic correction of the internal
`calibration curve from which oxygen saturation is derived inside the oximeter in situations where
`variations in contact pressure or site-to-site tissue heterogeneity may cause large measurement
`inaccuracies.
`Another object of the invention is to eliminate or reduce the effect of variations in the calibration
`of a reflectance pulse oximeter between subjects, since perturbations caused by contact pressure
`remain one of the major sources of errors in reflectance pulse oximetry. In fetal pulse oximetry,
`there are additional factors, which must be properly compensated for in order to produce an
`accurate and reliable measurement of oxygen saturation. For example, the fetal head is usually
`the presenting part, and is a rather easily accessible location for application of reflectance pulse
`oximetry. However, uterine contractions can cause large and unpredictable variations in the
`pressure exerted on the head and by the sensor on the skin, which can lead to large errors in the
`measurement of oxygen saturation by a dual-wavelength reflectance pulse oximeter. Another
`object of the invention is to provide accurate measurement of oxygen saturation in the fetus
`during delivery.
`The basis for the errors in the oxygen saturation readings of a dual-wavelength pulse oximeter is
`the fact that, in practical situations, the reflectance sensor applications affect the distribution of
`blood in the superficial layers of the skin. This is different from an ideal situation, when a
`reflectance sensor measures light backscattered from a homogenous mixture of blood and
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`bloodless tissue components. Therefore, the R and IR DC signals practically measured by
`photodetectors contain a relatively larger proportion of light absorbed by and reflected from the
`bloodless tissue compartments. In these uncontrollable practical situations, the changes caused
`are normally not compensated for automatically by calculating the normalized R/IR ratio since
`the AC portions of each photoplethysmogram, and the corresponding DC components, are
`affected differently by pressure or site-to-site variations. Furthermore, these changes depend not
`only on wavelength, but depend also on the sensor geometry, and thus cannot be eliminated
`completely by computing the normalized R/IR ratio, as is typically the case in dual-wavelength
`pulse oximeters.
`The inventor has found that the net result of this nonlinear effect is to cause large variations in
`the slope of the calibration curves. Consequently, if these variations are not compensated
`automatically, they will cause large errors in the final computation of SpO2, particularly at low
`oxygen saturation levels normally found in fetal applications.
`Another object of the present invention is to compensate for these variations and to provide
`accurate measurement of oxygen saturation. The invention consists of, in addition to two
`measurement sessions typically carried out in pulse oximetry based on measurements with two
`wavelengths centered around the peak emission values of 660nm660 nm (red spectrum) and
`940nm ± 20nm940 nm±20 nm (IR spectrum), one additional measurement session is carried out
`with an additional wavelength. At least one additional wavelength is preferably chosen to be
`substantially in the IR region of the electromagnetic spectrum, i.e., in the NIR-IR spectrum
`(having the peak emission value above 700nm700 nm). In a preferred embodiment the use of at
`least three wavelengths enables the calculation of an at least one additional ratio formed by the
`combination of the two IR wavelengths, which is mostly dependent on changes in contact
`pressure or site-to-site variations. In a preferred embodiment, slight dependence of the ratio on
`variations in arterial oxygen saturation that may occur, is easily minimized or eliminated
`completely, by the proper selection and matching of the peak emission wavelengths and spectral
`characteristics of the at least two IR-light sources.
`Preferably, the selection of the IR wavelengths is based on certain criteria. The IR wavelengths
`are selected to coincide with the region of the optical absorption curve where HbO2 absorbs
`slightly more light than Hb. The TR.IR wavelengths are in the spectral regions where the
`extinction coefficients of both Hb and HbO2 are nearly equal and remain relatively constant as a
`function of wavelength, respectively.
`In a preferred embodiment, tracking changes in the ratio formed by the two IR wavelengths, in
`real-time, permits automatic correction of errors in the normalized ratio obtained from the R-
`wavelength and each of the IR-wavelengths. The term "“ratio"” signifies the ratio of two values
`of AC/DC corresponding to two different wavelengths. This is similar to adding another equation
`to solve a problem with at least three unknowns (i.e., the relative concentrations of HbO2 and Hb,
`which are used to calculate SaO2, and the unknown variable fraction of blood-to-tissue volumes
`that effects the accurate determination of SaO2), which otherwise must rely on only two
`equations in the case of only two wavelengths used in conventional dual-wavelength pulse
`oximetry. In a preferred embodiment, a third wavelength provides the added ability to compute
`SaO2 based on the ratio formed from the R-wavelength and either of the IR- wavelengths. In a
`preferred embodiment, changes in these ratios are tracked and compared in real-time to
`determine which ratio produces a more stable or less noisy signal. That ratio is used
`predominantly for calculating SaO2.
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`The present invention utilizes collection of light reflected from the measurement location at
`different detection locations arranged along a closed path around light emitting elements, which
`can be LEDs or laser sources. Preferably, these detection locations are arranged in two
`concentric rings, the so-called "“near"” and "“far"” rings, around the light emitting elements.
`This arrangement enables optimal positioning of the detectors for high quality measurements,
`and enables discrimination between photodetectors receiving "“good"” information (i.e., AC and
`DC values which would result in accurate calculations of SpO2) and "“bad"” information (i.e.,
`AC and DC values which would result in inaccurate calculations of SpO2).
`There is thus provided according to one aspect of the present invention, a sensor for use in an
`optical measurement device for non-invasive measurements of blood parameters, the sensor
`comprising:
`(1) a light source for illuminating a measurement location with incident light of at least three
`wavelengths, the first wavelength lying in a red (R) spectrum, and the at least second and third
`wavelengths lying substantially in the infrared (IR) spectrum; and
`(2) a detector assembly for detecting light returned from the illuminated location, the detector
`assembly being arranged so as to define a plurality of detection locations along at least one
`closed path around the light source.
`The term "“closed path"” used herein signifies a closed curve, like a ring, ellipse, or polygon,
`and the like.
`The detector assembly is comprised of at least one array of discrete detectors (e.g., photodiodes)
`accommodated along at least one closed path, or at least one continuous photodetector defining
`the closed path.
`The term "“substantially IR spectrum"” used herein signifies a spectrum range including near
`infrared and infrared regions.
`According to another aspect of the present invention, there is provided a pulse oximeter utilizing
`a sensor constructed as defined above, and a control unit for operating the sensor and analyzing
`data generated thereby.
`According to yet another aspect of the present invention, there is provided a method for non-
`invasive determination of a blood parameter, the method comprising the steps of:
`illuminating a measurement location with at least three different wavelengths λlλ1, λ2 and λ3,
`the first wavelength λl λ1 lying in a red (R) spectrum, and the at least second and at least third
`wavelengths λ2 and λ3 lying substantially in the infrared (IR) spectrum;
`detecting light returned from the measurement location at different detection locations and
`generating data indicative of the detected light, wherein said different detection locations are
`arranged so as to define at least one closed path around the measurement location; and analyzing
`the generated data and determining the blood parameter.
`analyzing the generated data and determining the blood parameter.
`BRIEF DESCRIPTION OF THE DRAWINGS
`Other advantages of the present invention will be readily appreciated as the same becomes better
`understood by reference to the following detailed description when considered in connection
`with the accompanying drawings wherein:
`FigFIG. 1 illustrates hemoglobin spectra as measured by oximetry based techniques; Fig
`FIG. 2 illustrates a calibration curve used in pulse oximetry as typically programmed by the
`pulse oximeters manufacturers;
`FigFIG. 3 illustrates the relative disposition of light source and detector in reflection-mode or
`backscatter type pulse oximetry;
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`FigFIG. 4 illustrates light propagation in reflection pulse oximetry; Figs
`FIGS. 5A and 5B illustrate a pulse oximeter reflectance sensor operating under ideal and
`practical conditions, respectively;
`FigFIG. 6 illustrates variations of the slopes of calibration curves in reflectance pulse oximetry
`measurements;
`FigFIG. 7 illustrates an optical sensor according to the invention; Fig
`FIG. 8 is a block diagram of the main components of a pulse oximeter utilizing the sensor of
`FigFIG. 7;
`FigFIG. 9 is a flow chart of a selection process used in the signal processing technique according
`to the invention; and
`FigsFIGS. 10A to IOC 10C are flow charts of three main steps, respectively, of the signal
`processing method according to the invention.
`DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT EMBODIMENTS
`Referring to the Figures, wherein like numerals indicate like or corresponding parts throughout
`the several views, FigsFIGS. 1 and 2 illustrate typical hemoglobin spectra and calibrations curve
`utilized in the pulse oximetry measurements.
`The present invention provides a sensor for use in a reflection-mode or backscatter type pulse
`oximeter. The relative disposition of light source and detector in the reflection-mode pulse
`oximeter are illustrated in Fig. FIG. 3.
`FigFIG. 4 shows light propagation in the reflection-mode pulse oximeter where, in addition to
`the optical absorption and reflection due to blood, the DC signal of the R and IR
`photoplethysmograms can be adversely affected by strong reflections from the bone.
`Figs. 5 AFIGS. 5A and 5B illustrate a pulse oximeter reflectance sensor operating under,
`respectively, ideal and practical conditions. Referring now to FigFIG. 5A, it is shown that, under
`ideal conditions, reflectance sensor measures light backscattered from a homogenous mixture of
`blood and bloodless tissue components. Accordingly, the normalized R/IR ratio in dual-
`wavelength reflection type pulse oximeters, which relies on proportional changes in the AC and
`DC components in the photoplethysmograms, only reflect changes in arterial oxygen saturation.
`Referring now to FigFIG. 5B, in practical situations, the sensor applications affect the
`distribution of blood in the superficial layers of the skin. Accordingly, the R and IR DC signals
`measured by photodetectors contain a relatively larger proportion of light absorbed by and
`reflected from the bloodless tissue compartments. As such, the changes in DC signals depend not
`only on wavelength but also sensor geometry and thus cannot be eliminated completely by
`computing the normalized R/IR ratio, as is typically the case in dual-wavelength pulse oximeters.
`The result is large variations in the slope of the calibration curves, as illustrated in Fig. FIG. 6.
`Referring now to FigFIG. 6, graphs ClC1, C2 and C3 show three calibration curves, presenting
`the variation of the slope for oxygen saturation values between 50% and 100%.
`Referring to FigFIG. 7, there is illustrated an optical sensor 10 designed according to the
`invention aimed at minimizing some of the measurement inaccuracies in a reflectance pulse
`oximeter. The sensor 10 comprises such main constructional parts as a light
`source 12 composed of three closely spaced light emitting elements (e.g., LEDs or laser
`sources) 12a, 12b 12 a, 12 b and 12c 12 c generating light of three different wavelengths,
`respectively; an array of discrete detectors (e.g., photodiodes), a "“far"” detector 16 and a
`"“near"” detector 18, arranged in two concentric ring-like arrangements (constituting closed
`paths) surrounding the light emitting elements; and a light shield 14. In the present example, six
`photodiodes form each ring. All these elements are accommodated in a sensor housing 17. The
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`light shield 14 is positioned between the photodiodes and the light emitting elements, and
`prevents direct optical coupling between them, thereby maximizing the fraction of backscattered
`light passing through the arterially perfused vascular tissue in the detected light.
`It should be noted that more than three wavelengths can be utilized in the sensor. The actual
`numbers of wavelengths used as a light source and the number of photodetectors in each ring are
`not limited and depend only on the electronic circuitry inside the oximeter. The array of discrete
`photodiodes can be replaced by one or more continuous photodetector rings.
`In addition to the R and IR light emitting elements 12a 12 a and 12b 12 b as used in the
`conventional pulse oximeter sensors, the sensor 10 incorporates the