Abstract:
A patient monitor has multiple sensors adapted to attach to tissue sites of a living subject. The sensors generate sensor signals that are responsive to at least two wavelengths of optical radiation after attenuation by pulsatile blood within the tissue sites.

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS  
       [0001]     This application is a continuation of U.S. application Ser. No. 11/104,720, filed Apr. 13, 2005, which is a continuation of U.S. application Ser. No. 10/668,487, filed Sep. 22, 2003, now U.S. Pat. No. 6,898,452, which is a continuation of U.S. application Ser. No. 10/026,013, filed Dec. 21, 2001, now U.S. Pat. No. 6,714,804, which is a continuation of U.S. application Ser. No. 09/323,176, filed May 27, 1999, now U.S. Pat. No. 6,334,065, which claims priority from U.S. Provisional No. 60/087,802, filed Jun. 3, 1998. 
     
    
     BACKGROUND  
       [0002]     The measurement of oxygen delivery to the body and the corresponding oxygen consumption by its organs and tissues is vitally important to medical practitioners in the diagnosis and treatment of various medical conditions. Oxygen delivery, the transport of oxygen from the environment to organs and tissues, depends on the orchestration of several interrelated physiologic systems. Oxygen uptake is determined by the amount of oxygen entering the lung and the adequacy of gas exchange within the lung. This gas exchange is determined by the diffusion of oxygen from the alveolar space to the blood of the pulmonary capillaries. Oxygen is subsequently transported to all organs and tissues by blood circulation maintained by the action of the heart. The availability of oxygen to the organs and tissues is determined both by cardiac output and by the oxygen content in the blood. Oxygen content, in turn, is affected by the concentration of available hemoglobin and hemoglobin oxygen saturation. Oxygen consumption is related to oxygen delivery according to Fick&#39;s axiom, which states that oxygen consumption in the peripheral tissues is equal to oxygen delivery via the airway.  
         [0003]     Oxygen delivery and oxygen consumption can be estimated from a number of measurable parameters. Because of the diagnostic impracticalities of measuring oxygen uptake and cardiac output, oxygen delivery is typically assessed from the oxygen status of arterial blood alone, such as arterial oxygen partial pressure, P a O 2 , and arterial oxygen saturation, S a O 2 . P a O 2  represents the relatively small amount of oxygen dissolved in the blood plasma. S a O 2  represents the much larger amount of oxygen chemically bound to the blood hemoglobin. Oxygen consumption is typically assessed from the oxygen status of mixed venous blood, i.e. the oxygen saturation of blood from the pulmonary artery, S v O 2 , which is used to estimate the O 2  concentration of blood returning from all tissues and organs of the body. These parameters can be measured by both invasive and non-invasive techniques, except S v O 2 , which requires an invasive measurement.  
         [0004]     Invasive techniques include blood gas analysis using the in vitro measurement of extracted arterial or venous blood, drawn with a syringe and needle or an intervascular catheter. Arterial blood is commonly obtained by puncturing the brachial, radial or femoral artery. Venous blood can be obtained from an arm vein, but such a sample reflects only local conditions. To obtain mixed venous blood, which represents the composite of all venous blood, a long catheter is typically passed through the right heart and into the main pulmonary artery from a peripheral vein. Extracted blood gas analysis utilizes blood gas machines or oximeters. A blood gas machine measures the partial pressure of oxygen, PO 2 , using a “Clark electrode” that detects the current generated by oxygen diffusing to a sealed platinum electrode across a gas permeable membrane. An oximeter measures the oxygen saturation, SO 2 , of oxygenated and deoxygenated hemoglobin using spectrophotometry techniques that detect the differential absorption of particular wavelengths of light by these blood components.  
         [0005]     Invasive monitoring also includes the in vivo monitoring of blood gas via a catheter sensor inserted into an artery or vein. Miniaturization of the Clark electrode allows placement of the electrode in a catheter for continuous measurement of PO 2 . A fiber optic equipped catheter attached to an external oximeter allows continuous measurement of oxygen saturation. Because of risks inherent in catheterization and the promotion of blood coagulation by certain sensors, these techniques are typically only used when vitally indicated.  
         [0006]     Non-invasive techniques include pulse oximetry, which allows the continuous in vivo measurement of arterial oxygen saturation and pulse rate in conjunction with the generation of a photoplethsymograph waveform. Measurements rely on sensors which are typically placed on the fingertip of an adult or the foot of an infant. Non-invasive techniques also include transcutaneous monitoring of PO 2 , accomplished with the placement of a heated Clark electrode against the skin surface. These non-invasive oxygen status measurement techniques are described in further detail below.  
       SUMMARY  
       [0007]     Prior art invasive oxygen assessment techniques are inherently limited. Specifically, in vitro measurements, that is, blood extraction and subsequent analysis in a blood gas machine or an oximeter, are non-simultaneous and non-continuous. Further, in vivo measurements through catheterization are not casual procedures and are to be particularly avoided with respect to neonates. Prior art noninvasive techniques are also limited. In particular, conventional pulse oximeters are restricted to measurement of arterial oxygen saturation at a single patient site. Also, transcutaneous monitoring is similarly restricted to the measurement of an estimate of arterial partial pressure at a single patient site, among other limitations discussed further below.  
         [0008]     The stereo pulse oximeter according to the present invention overcomes many of the limitations of prior art oxygen status measurements. The word “stereo” comes from the Greek word stereos, which means “solid” or three-dimensional. For example, stereophonic systems use two or more channels to more accurately reproduce sound. The stereo pulse oximeter is similarly multi-dimensional, providing simultaneous, continuous, multiple-site and multiple-parameter oxygen status and plethysmograph (photoplethysmograph) measurements. The stereo pulse oximeter provides a benefit in terms of cost and patient comfort and safety over invasive oxygen status estimation techniques. The multi-dimensional aspects of this invention further provide oxygen status and plethysmograph measurements not available from current noninvasive techniques. In addition, the stereo pulse oximeter allows the isolation of noise artifacts, providing more accurate oxygen status and plethysmograph measurements than available from conventional techniques. The result is improved patient outcome based on a more accurate patient assessment and better management of patient care.  
         [0009]     In one aspect of the stereo pulse oximeter, data from a single sensor is processed to advantageously provide continuous and simultaneous multiple-parameter oxygen status and plethysmograph measurements from a particular tissue site. This is in contrast to a conventional pulse oximeter that provides only arterial oxygen saturation data from a tissue site. In particular a physiological monitor comprises a sensor interface and a signal processor. The sensor interface is in communication with a peripheral tissue site and has an output responsive to light transmitted through the site. The signal processor is in communication with the sensor interface output and provides a plurality of parameters corresponding to the oxygen status of the site, the plethysmograph features of the site or both. The parameters comprise a first value and a second value related to the peripheral tissue site. In one embodiment, the first value is an arterial oxygen saturation and the second value is a venous oxygen saturation. In this embodiment, another parameter provided may be the difference between arterial oxygen saturation and venous oxygen saturation at the tissue site. The venous oxygen saturation is derived from an active pulse generated at the site. The signal processor output may further comprise a scattering indicator corresponding to the site, and the sensor interface may further comprise a pulser drive, which is responsive to the scattering indicator to control the amplitude of the active pulse. One of the parameter values may also be an indication of perfusion.  
         [0010]     In another aspect of the stereo pulse oximeter, data from multiple sensors is processed to advantageously provide continuous and simultaneous oxygen status measurements from several patient tissue sites. This is in contrast to a conventional pulse oximeter that processes data from a single sensor to provide oxygen status at a single tissue site. In particular, a physiological monitor comprises a plurality of sensor interfaces each in communications with one of a plurality of peripheral tissue sites. Each of the sensor interfaces has one of a plurality of outputs responsive to light transmitted through a corresponding one of the tissue sites. A signal processor is in communication with the sensor interface outputs and has a processor output comprising a plurality of parameters corresponding to the oxygen status of the sites, the plethysmograph features of the sites or both. The parameters may comprise a first value relating to a first of the peripheral tissue sites and a second value relating to a second of the peripheral tissue sites. In one embodiment, the first value and the second value are arterial oxygen saturations. In another embodiment, the first value and the second value are plethysmograph waveform phases. The physiological monitor may further comprise a sensor attachable to each of the tissue sites. This sensor comprises a plurality of emitters and a plurality of detectors, where at least one of the emitters and at least one of the detectors is associated with each of the tissue sites. The sensor also comprises a connector in communications with the sensor interfaces. A plurality of signal paths are attached between the emitters and the detectors at one end of the sensor and the connector at the other end of the sensor.  
         [0011]     In yet another aspect of the stereo pulse oximeter, data from multiple sensors is processed to advantageously provide a continuous and simultaneous comparison of the oxygen status between several tissue sites. A conventional oximeter, limited to measurements at a single tissue site, cannot provide these cross-site comparisons. In particular a physiological monitoring method comprises the steps of deriving a reference parameter and a test parameter from oxygen status measured from at least one of a plurality of peripheral tissue sites and comparing that reference parameter to the test parameter so as to determine a patient condition. The reference parameter may be a first oxygen saturation value and the test parameter a second oxygen saturation value. In that case, the comparing step computes a delta oxygen saturation value equal to the arithmetic difference between the first oxygen saturation value and the second oxygen saturation value. In one embodiment, the reference parameter is an arterial oxygen saturation measured at a particular one the tissue sites and the test parameter is a venous oxygen saturation measured at that particular site. In another embodiment, the reference parameter is a first arterial oxygen saturation value at a first of the tissue sites, the test parameter is a second arterial oxygen saturation value at a second of the tissue sites. In yet another embodiment, the reference parameter is a plethysmograph feature measured at a first of the sites, the test parameter is a plethysmograph feature measured at a second of the sites and the monitoring method comparison step determines the phase difference between plethysmographs at the first site and the second site. In a further embodiment, the comparing step determines a relative amount of damping between plethysmographs at the first site and the second site. The multi-dimensional features of these embodiments of the stereo pulse oximeter can be advantageously applied to the diagnosis and managed medical treatment of various medical conditions. Particularly advantageous applications of stereo pulse oximetry include oxygen titration during oxygen therapy, nitric oxide titration during therapy for persistent pulmonary hypertension in neonates (PPHN), detection of a patent ductus arteriosis (PDA), and detection of an aortic coarctation. 
     
    
     BRIEF DESCRIPTION OF THE DRAWINGS  
       [0012]     The present invention will be described in detail below in connection with the following drawing figures in which:  
         [0013]      FIG. 1A  is a top-level block diagram of a stereo pulse oximeter according to the present invention;  
         [0014]      FIG. 1B  shows a single-sensor alternative embodiment to  FIG. 1A ;  
         [0015]      FIG. 2  is a block diagram of the stereo pulse oximeter sensor interface;  
         [0016]      FIG. 3  is a graph illustrating the absorption of red and infrared wavelengths by both oxygenated and deoxygenated hemoglobin;  
         [0017]      FIG. 4  is a graph showing the empirical relationship between the “red over infrared” ratio and arterial oxygen saturation;  
         [0018]      FIG. 5  is a block diagram of the analog signal conditioning for the sensor interface;  
         [0019]      FIG. 6  is a functional block diagram of the stereo pulse oximeter signal processing;  
         [0020]      FIG. 7  is a functional block diagram of the front-end signal processing;  
         [0021]      FIG. 8  is a graph depicting the frequency spectrum of an arterial intensity signal;  
         [0022]      FIG. 9  is a graph depicting the frequency spectrum of a combined arterial and venous intensity signal;  
         [0023]      FIG. 10  is a functional block diagram of the saturation calculation signal processing;  
         [0024]      FIG. 11  is a graph illustrating a plethysmograph waveform;  
         [0025]      FIG. 12  is a graph illustrating the absorption contribution of various blood and tissue components;  
         [0026]      FIG. 13  is a graph illustrating an intensity “plethysmograph” pulse oximetry waveform;  
         [0027]      FIG. 14  is a functional block diagram of the plethysmograph feature extraction signal processing;  
         [0028]      FIG. 15  is a functional block diagram of the multiple parameter signal processing;  
         [0029]      FIG. 16A  is an illustration of a single-site stereo pulse oximeter display screen;  
         [0030]      FIG. 16B  is an illustration of a multi-site stereo pulse oximeter display screen;  
         [0031]      FIG. 17A  is a graph depicting a family of constant power curves for the electrical analog of constant oxygen consumption;  
         [0032]      FIG. 17B  is a graph depicting arterial and venous oxygen saturation versus fractional inspired oxygen;  
         [0033]      FIG. 17C  is a graph depicting arterial minus venous oxygen saturation versus fractional inspired oxygen;  
         [0034]      FIG. 18  is a three-dimensional graph depicting a delta oxygen saturation surface;  
         [0035]      FIG. 19  is an illustration of a neonatal heart depicting a pulmonary hypertension condition;  
         [0036]      FIG. 20  is an illustration of a fetal heart depicting the ductus arteriosis; and  
         [0037]      FIG. 21  is an illustration of a neonatal heart depicting a patent ductus arteriosis (PDA). 
     
    
     DETAILED DESCRIPTION  
       [0038]     Stereo Pulse Oximetry  
         [0039]      FIG. 1A  illustrates the multi-dimensional features of a stereo pulse oximeter  100  according to the present invention. Shown in  FIG. 1A  is an exemplary stereo pulse oximeter configuration in which a first sensor  110  is attached to a neonate&#39;s left hand, a second sensor  120  is attached to one of the neonate&#39;s feet, and a third sensor  130  is attached to the neonate&#39;s right hand. In general, these sensors are used to obtain oxygen status and photoplethysmograph measurements at peripheral sites, including a person&#39;s ears and face, such as the nose and regions of the mouth in addition to hands, feet and limbs, but not including internal sites such as internal organs and the brain.  
         [0040]     Each sensor  110 ,  120 ,  130  provides a stream of data through a corresponding sensor interface  114 ,  124 ,  134  to the digital signal processor (DSP)  150 . For example, the first sensor  110  is connected to an input  112  of the first sensor interface  114 , and the output  118  of the first sensor interface  114  is attached to a first data channel input  152  of the DSP  150 . Similarly, the second sensor  120  provides data to a second data channel input  154  and the third sensor  130  provides data to a third data channel input  158 .  
         [0041]      FIG. 1B  illustrates an alternative embodiment of the separate sensors  110 ,  120 ,  130  ( FIG. 1A ). A stereo sensor  140  has multiple branches  112 ,  122 ,  132  each terminating in a sensor portion  114 ,  124 ,  134 . Each sensor portion  114 ,  124 ,  134  has two light emitters and a light detector, as described below, and is attachable to a separate patient site. Thus, the stereo sensor  140  advantageously provides a single sensor device having multiple light emitters and multiple light detectors for attachment to multiple patient tissue sites. A combination of the stereo sensor  140  and a single patient cable  142  advantageously allows a single connection  144  at the stereo pulse oximeter  100  and a single connection  146  at the stereo sensor  140 .  
         [0042]     The DSP  150  can independently process each data channel input  152 ,  154 ,  158  and provide outputs  162  typical of pulse oximetry outputs, such as arterial oxygen saturation, Sp a O 2 , the associated plethysmograph waveform and the derived pulse rate. In contrast with a conventional pulse oximeter, however, these outputs  162  include simultaneous measurements at each of several patient tissue sites. That is, for the configuration of  FIG. 1A , the stereo pulse oximeter  100  simultaneously displays Sp a O 2  and an associated plethysmograph waveform for three tissue sites in addition to the patient&#39;s pulse rate obtained from any one of sites. Further, the DSP  150  can provide unique outputs unavailable from conventional pulse oximeters. These outputs  164  include venous oxygen saturation, Sp v O 2 , a comparison of arterial and venous oxygen saturation, Δ sat =Sp av O 2 =Sp a O 2 −Sp v O 2 , and pleth, which denotes plethysmograph shape parameters, for each site. In addition, the DSP  150  can provide cross-site outputs that are only available using stereo pulse oximetry. These unique cross-site outputs  168  include Δsat xy =Sp ax O 2 −Sp ay O 2 , which denotes the arterial oxygen saturation at site x minus the arterial oxygen saturation at site y. Also included in these outputs  168  is Δpleth xy , which denotes a comparison of plethysmograph shape parameters measured at site x and site y, as described in detail below. The stereo pulse oximeter also includes a display  180  capable of showing the practitioner the oxygen status and plethysmograph parameters described above. The display  180  has a multiple channel graphical and numerical display capability as described in more detail below.  
         [0043]     Pulse Oximetry Sensor  
         [0044]      FIG. 2  depicts one stereo pulse oximeter data channel having a sensor  110  and a sensor interface  114  providing a single data channel input  152  to the DSP  150 . The sensor  110  is used to measure the intensity of red and infrared light after transmission through a portion of the body where blood flows close to the surface, such as a fingertip  202 . The sensor  110  has two light emitters, each of which may be, for example, a light-emitting diode (LED). A red emitter  212 , which transmits light centered at a red wavelength and an infrared (IR) emitter  214 , which transmits light centered at an infrared wavelength are placed adjacent to, and illuminate, a tissue site. A detector  218 , which may be a photodiode, is used to detect the intensity of the emitted light after it passes through, and is partially absorbed by, the tissue site. The emitters  212 ,  214  and detector  218  are secured to the tissue site, with the emitters  212 ,  214  typically spaced on opposite sides of the tissue site from the detector  218 .  
         [0045]     To distinguish between tissue absorption at the two wavelengths, the red emitter  212  and infrared emitter  214  are modulated so that only one is emitting light at a given time. In one embodiment, the red emitter  212  is activated for a first quarter cycle and is off for the remaining three-quarters cycle; the infrared emitter  214  is activated for a third quarter cycle and is off for the remaining three-quarters cycle. That is, the emitters  212 ,  214  are cycled on and off alternately, in sequence, with each only active for a quarter cycle and with a quarter cycle separating the active times. The detector  218  produces an electrical signal corresponding to the red and infrared light energy attenuated from transmission through the patient tissue site  202 . Because only a single detector  218  is used, it receives both the red and infrared signals to form a time-division-multiplexed (TDM) signal. This TDM signal is coupled to the input  112  of the sensor interface  114 . One of ordinary skill in the art will appreciate alternative activation sequences for the red emitter  212  and infrared emitter  214  within the scope of this invention, each of which provides a time multiplexed signal from the detector  218  allowing separation of red and infrared signals and determination and removal of ambient light levels in downstream signal processing.  
         [0046]     To compute Sp a O 2 , pulse oximetry relies on the differential light absorption of oxygenated hemoglobin, HbO 2 , and deoxygenated hemoglobin, Hb, to compute their respective concentrations in the arterial blood. This differential absorption is measured at the red and infrared wavelengths of the sensor  110 . The relationship between arterial oxygen saturation and hemoglobin concentration can be expressed as:  
                 Sp   a     ⁢     O   2       =     100   ⁢       C     HbO   ⁢           ⁢   2           C   Hb     +     C     HbO   ⁢           ⁢   2                     (   1   )             
 
 That is, arterial oxygen saturation is the percentage concentration of oxygenated hemoglobin compared to the total concentration of oxygenated hemoglobin and deoxygenated hemoglobin in the arterial blood. Sp a O 2  is actually a measure of the partial oxygen saturation of the hemoglobin because other hemoglobin derivatives, such as COHb and MetHb, are not taken into consideration. 
 
         [0047]      FIG. 3  shows a graph  300  of the optical absorption properties of HbO 2  and Hb. The graph  300  has an x-axis  310  corresponding to wavelength and a y-axis  320  corresponding to hemoglobin absorption. An Hb curve  330  shows the light absorption properties of deoxygenated hemoglobin. An HbO 2  curve  340  shows the light absorption properties of oxygenated hemoglobin. Pulse oximetry measurements are advantageously made at a red wavelength  350  corresponding to 660 nm and an infrared wavelength  360  corresponding to 905 nm. This graph  300  shows that, at these wavelengths  350 ,  360 , deoxygenated hemoglobin absorbs more red light than oxygenated hemoglobin, and, conversely, oxygenated hemoglobin absorbs more infrared light than deoxygenated hemoglobin.  
         [0048]     In addition to the differential absorption of hemoglobin derivatives, pulse oximetry relies on the pulsatile nature of arterial blood to differentiate hemoglobin absorption from absorption of other constituents in the surrounding tissues. Light absorption between systole and diastole varies due to the blood volume change from the inflow and outflow of arterial blood at a peripheral tissue site. This tissue site might also comprise skin, muscle, bone, venous blood, fat, pigment, etc., each of which absorbs light. It is assumed that the background absorption due to these surrounding tissues is invariant and can be ignored. Thus, blood oxygen saturation measurements are based upon a ratio of the time-varying or AC portion of the detected red and infrared signals with respect to the time-invariant or DC portion. This AC/DC ratio normalizes the signals and accounts for variations in light pathlengths through the measured tissue. Further, a ratio of the normalized absorption at the red wavelength over the normalized absorption at the infrared wavelength is computed:  
               RD   IR     =       (       Red   AC       Red   DC       )       (       IR   AC       IR   DC       )               (   2   )             
 
 where Red AC  and IR AC  are the root-mean-square (RMS) of the corresponding time-varying signals. This “red-over-infrared, ratio-of-ratios” cancels the pulsatile signal. The desired Sp a O 2  measurement is then computed from this ratio. 
 
         [0049]      FIG. 4  shows a graph  400  depicting the relationship between RD/IR and Sp a O 2 . This relationship can be approximated from Beer-Lambert&#39;s Law, as outlined below. However, it is most accurately determined by statistical regression of experimental measurements obtained from human volunteers and calibrated measurements of oxygen saturation. The result can be depicted as a curve  410 , with measured values of RD/IR shown on a y-axis  420  and corresponding saturation values shown on an x-axis  430 . In a pulse oximeter device, this empirical relationship can be stored in a read-only memory (ROM) look-up table so that Sp a O 2  can be directly read-out from input RD/IR measurements.  
         [0050]     According to the Beer-Lambert law of absorption, the intensity of light transmitted through an absorbing medium is given by:  
             I   =       I   0     ⁢     exp   ⁡     (     -       ∑     i   =   1     N     ⁢       ɛ     i   ⁢           ⁢   λ       ⁢     c   i     ⁢     x   i           )                 (   3   )             
 
 where I 0  is the intensity of the incident light, ε i,λ . is the absorption coefficient of the i th  constituent at a particular wavelength λ, c i  is the concentration coefficient of the i th  constituent and x i  is the optical path length of the i th  constituent. As stated above, assuming the absorption contribution by all constituents but the arterial blood is constant, taking the natural logarithm of both sides of equation (3) and removing time invariant terms yields: 
 
ln( l )=−[ε HbO2,λ C HbO2 +ε Hb,λ C hb   ]x ( t )  (4) 
 
 Measurements taken at both red and infrared wavelengths yield: 
 
RD( t )=−[ε HbO2,RD C HbO2 +ε Hb,RD C hb   ]x   RD   ( t )  ( 5 ) 
 
IR( t )=−[ε HbO2,IR C HbO2 +ε Hb,IR C hb   ]x   IR ( t )  (6) 
 
 Taking the ratio RD(t)/IR(t) and assuming x RD (t)≈x IR (t) yields: 
 
RD/IR=[ε HbO2,RD C HbO2 +ε Hb,RD C Hb ]/[ε HbO2,IR C HbO2 +ε Hb,IR C hb ]  (7) 
 
 Assuming further that: 
 
C HbO2 +C Hb =1  (8) 
 
 then equation (1) can be solved in terms of RD/IR yielding a curve similar to the graph  400  of  FIG. 4 . 
 
         [0051]     Sensor Interface  
         [0052]      FIG. 2  also depicts the sensor interface  114  for one data channel. An interface input  112  from the sensor  110  is coupled to an analog signal conditioner  220 . The analog signal conditioner  220  has an output  223  coupled to an analog-to-digital converter (ADC)  230 . The ADC output  118  is coupled to the DSP  150 . The analog signal conditioner also has a gain control input  225  from the DSP  150 . The functions of the analog signal conditioner  220  are explained in detail below. The ADC  230  functions to digitize the input signal  112  prior to further processing by the DSP  150 , as described below. The sensor interface  114  also has an emitter current control input  241  coupled to a digital-to-analog converter (DAC)  240 . The DSP provides control information to the DAC  240  via the control input  241  for a pair of emitter current drivers  250 . One driver output  252  couples to the red emitter  212  of the sensor  110 , and another driver output  254  couples to the IR emitter  214  of the sensor  110 .  
         [0053]      FIG. 5  illustrates one embodiment of the analog signal conditioner  220 . The analog signal conditioner  220  receives a composite intensity signal  112  from the sensor detector  218  ( FIG. 2 ) and then filters and conditions this signal prior to digitization. The embodiment shown has a preamplifier  510 , a high pass filter  520 , a programmable gain amplifier  530  and a low pass filter  540 . The low pass filter output  223  is coupled to the ADC  230  ( FIG. 2 ). The preamplifier  510  converts the current signal  112  from the detector  218  ( FIG. 2 ) to a corresponding amplified voltage signal. The gain in the preamplifier  510  is selected in order to prevent ambient light in the signal  112  from saturating the preamplifier  510  under normal operating conditions. The preamplifier output  512  is coupled to the high pass filter  520 , which removes the DC component of the detector signal  112 . The corner frequency of the high pass filter  520  is set well below the multiplexing frequency of the red and infrared emitters  212 ,  214  ( FIG. 2 ). The high pass filter output  522  couples to the programmable gain amplifier  530 , which also accepts a programming input  225  from the DSP  150  ( FIG. 2 ). This gain is set at initialization or at sensor placement to compensate for variations from patient to patient. The programmable gain amplifier output  532  couples to a low-pass filter  540  to provide anti-aliasing prior to digitization.  
         [0054]     As described above, pulse oximetry measurements rely on the existence of a pulsatile signal. The natural heart beat provides a pulsatile signal that allows measurement of arterial oxygen saturation. In the systemic circulation, all arterial pulsations are damped before flow enters the capillaries, and none are transmitted into the veins. Thus, there is no arterial pulse component in the venous blood and absorption caused by venous blood is assumed canceled by the ratio-of-ratio operation described above. Venous blood, being at a relatively low pressure, will “slosh back and forth” during routine patient motions, such as shivering, waving and tapping. This venous blood sloshing creates a time-varying signal that is considered “noise” and can easily overwhelm conventional ratio-based pulse oximeters. Advanced pulse oximetry techniques allow measurement of Sp v O 2  under these circumstances. For example, such advanced techniques are disclosed in U.S. Pat. No. 5,632,272, which is assigned to the assignee of the current application. This measurement is only available during motion or other physiological events causing a time-varying venous signal.  
         [0055]     The venous blood may also have a pulsatile component at the respiration rate, which can be naturally induced or ventilator induced. In adults, the natural respiration rate is 10-15 beats per minute (bpm). In neonates, this natural respiration rate is 30-60 bpm. The ventilator induced pulse rate depends on the ventilator frequency. If this respiration induced venous pulse is of sufficient magnitude, advanced pulse oximetry techniques, described below, allow measurement of Sp v O 2 .  
         [0056]     A controlled physiological event, however, can be created that allows for a continuous measurement of venous oxygen saturation, independent of motion or respiration. U.S. Pat. No. 5,638,816, which is assigned to the assignee of the current application discloses a technique for inducing an intentional active perturbation of the blood volume of a patient, and is referred to as an “active pulse.” Because peripheral venous oxygen saturation, Sp v O 2 , is a desirable parameter for stereo pulse oximetry applications, it is advantageous to provide for a continuous and controlled pulsatile venous signal.  
         [0057]      FIG. 2  depicts an active pulse mechanism used in conjunction with a pulse oximetry sensor. An active pulser  260  physically squeezes or otherwise perturbs a portion of patient tissue  270  in order to periodically induce a “pulse” in the blood at the tissue site  202 . A pulser drive  280  generates a periodic electrical signal to a transducer  262  attached to the patient. The transducer  262  creates a mechanical force against the patient tissue  270 . For example, the pulser  260  could be a solenoid type device with a plunger that presses against the fleshy tissue to which it is attached. The DSP  150  provides pulse drive control information to a digital to analog converter (DAC)  290  via the control input  291 . The DAC output  292  is coupled to the pulser drive  280 . This allows the processor to advantageously control the magnitude of the induced pulse, which moderates scattering as described below. The pulser  260  could be a pressure device as described above. Other pressure mechanisms, for example a pressure cuff, could be similarly utilized. Other methods, such as temperature fluctuations or other physiological changes, which physiologically alter a fleshy medium of the body on a periodic basis to modulate blood volume at a nearby tissue site could also be used. Regardless of the active pulse mechanism, this modulated blood volume is radiated by a pulse oximeter sensor and the resulting signal is processed by the signal processing apparatus described below to yield Sp v O 2 .  
         [0058]     Signal Processor  
         [0059]      FIG. 6  illustrates the processing functions of the digital signal processor (DSP)  150  ( FIG. 1A ). Each data channel input  152 ,  154 ,  158  ( FIG. 1A ) is operated on by one or more of the front-end processor  610 , saturation calculator  620 , plethysmograph feature extractor  630  and multiple parameter processor  640  functions of the DSP  150 . First, a digitized signal output from the ADC  230  ( FIG. 2 ) is input  602  to the front-end processor  610 , which demultiplexes, filters, normalizes and frequency transforms the signal, as described further below. A front-end output  612  provides a red signal spectrum and an IR signal spectrum for each data channel as inputs to the saturation calculator  620 . Another front-end output  614  provides a demultiplexed, normalized IR plethysmograph for each data channel as an input to the feature extractor  630 . The saturation calculator output  622  provides arterial and venous saturation data for each data channel as input to the multiple parameter processor  640 . One feature extractor output  632  provides data on various plethysmograph shape parameters for each data channel as input into the multiple parameter processor  640 . Another feature extractor output  634 , also coupled to multiple parameter processor  640 , provides an indication of plethysmograph quality and acts as a threshold for determining whether to ignore portions of the input signal  602 . The multiple parameter processor has a numerical output  642  that provides same-channel Δsat parameters and cross-channel parameters, such as Δsat xy  or Δpleth xy  to a display  180  ( FIG. 1A ). The numeric output  642  may also provide saturation and plethysmograph parameters directly from the saturation calculator  620  or the feature extractor  630  without further processing other than data buffering. The multiple parameter processor also has a graphical output  644  that provides plethysmograph waveforms for each data channel in addition to graphics, depending on a particular application, the indicate the trend of the numerical parameters described above.  
         [0060]     Front-End Processor  
         [0061]      FIG. 7  is a functional block diagram of the front-end processor  610  for the stereo pulse oximeter. The digitized sensor output  118  ( FIG. 2 ) is an input signal  602  to a demultiplexer  710 , which separates the input signal  602  into a red signal  712  and an infrared signal  714 . The separated red and infrared signals  712 ,  714  are each input to a filter  720  to remove unwanted artifacts introduced by the demultiplexing operation. In one embodiment, the filter  720  is a finite-impulse-response, low-pass filter that also “decimates” or reduces the sample rate of the red and infrared signals  712 ,  714 . The filtered signals  722  are then each normalized by a series combination of a log function  730  and bandpass filter  740 . The normalized signals, RD(t), IR(t)  742  are coupled to a Fourier transform  750 , which provides red frequency spectrum and infrared frequency spectrum outputs, RD(ω), IR(ω)  612 . A demultiplexed infrared signal output  614  is also provided.  
         [0062]     Saturation Calculator  
         [0063]      FIG. 8  shows a graph  800  illustrating idealized spectrums of RD(t) and IR(t)  752  ( FIG. 7 ). The graph has an x-axis  810  that corresponds to the frequency of spectral components in these signals and a y-axis  820  that corresponds to the magnitude of the spectral components. The spectral components are the frequency content of RD(t) and IR(t), which are plethysmograph signals corresponding to the patient&#39;s pulsatile blood flow, as described below. Thus, the frequencies shown along the x-axis  810 , i.e. f 0 , f 1 , f 2 , are the fundamental and harmonics of the patient&#39;s pulse rate. The spectrum of RD(t), denoted RD(ω)  612  ( FIG. 7 ), is shown as a series of peaks, comprising a first peak  832  at a fundamental frequency, f 0 , a second peak  842  at a first harmonic, f 1  and a third peak  852  at a second harmonic, f 2 . Similarly, the spectrum of IR(t), denoted IR(ω)  612  ( FIG. 7 ), is shown as another series of peaks, comprising a first peak  834  at a fundamental frequency, f 0 , a second peak  844  at a first harmonic, f 1  and a third peak  854  at a second harmonic, f 2 . Also shown in  FIG. 8  is the ratio of the spectral peaks of RD(t) and IR(t), denoted RD(ω)/IR(ω). This ratio is shown as a first ratio line  838  at the fundamental frequency f 0 , a second ratio line  848  at the first harmonic f 1 , and a third ratio line  858  at the second harmonic f 2 .  
         [0064]     The magnitude of these ratio lines RD(ω)/IR(ω) corresponds to the ratio RD/IR defined by equation (2), and, hence, can be used to determine Sp a O 2 . This can be seen from Parseval&#39;s relation for a periodic signal, x(t), having a period T, where X k  is the spectral component at the kth harmonic of x(t):  
                 1   T     ⁢       ∫   0   T     ⁢         (          x   ⁡     (   t   )            )     2     ⁢     ⅆ   t           =       ∑             ⁢   k               ⁢           ⁢       (          X     k   ⁢                    )     2               (   9   )             
 
 Equation (9) relates the energy in one period of the signal x(t) to the sum of the squared magnitudes of the spectral components. The term |X k | 2  can be interpreted as that part of the energy per period contributed by the kth harmonic. In an ideal measurement, the red and infrared signals are the same to within a constant scale factor, which corresponds to the arterial oxygen saturation. Likewise, the red and infrared spectra are also the same to within a constant scale factor. Thus, in an ideal measurement, all of the ratio lines  838 ,  848 ,  858  have substantially the same amplitude. Any differences in the amplitude of the ratio lines is likely due to motion, scattering or other noise contaminations, as discussed further below. Accordingly, any of the RD(ω)/IR(ω) ratio lines is equivalent to the ratio, RD/IR, of equation (2) and can be used to derive Sp a O 2 . 
 
         [0065]     One skilled in the art will recognize that the representations in  FIG. 8  are idealized. In particular, in actual measured data, especially if contaminated by noise, the frequencies of the peaks of RD(ω) do not correspond exactly to the frequencies of the peaks of IR(ω). For example, the fundamental frequency, f 0 , found for RD(ω) will often be different from the fundamental frequency, f 0 ′, found for IR(ω) and similarly for the harmonics of f 0 .  
         [0066]      FIG. 9  shows a graph  900  illustrating idealized spectrums RD(ω) and IR(ω) and associated ratio lines measured with an active pulse sensor. The graph  900  has an x-axis  910  that corresponds to the frequency of spectral components in these signals and a y-axis  920  that corresponds to the magnitude of the spectral components. The spectrum, RD(ω), is shown as two series of peaks. One series of peaks  930  occurs at a fundamental frequency, f h0 , and associated harmonics, f h1  and f h2 , of the patient&#39;s pulse (heart) rate. Another series of peaks  940  occurs at a fundamental frequency, f p0 , and associated harmonics, f p1  and f p2 , Of the active pulse rate. Similarly, the spectrum, IR(ω), is shown as two series of peaks. One series of peaks  950  occurs at a fundamental frequency, f h0 , and associated harmonics, f h1  and f h2 , of the patient&#39;s pulse rate. Another series of peaks  960  occurs at a fundamental frequency, f p0 , and associated harmonics, f p1  and f p2 , of the active pulse rate. Accordingly, there are two series of RD/IR ratio lines. One series of ratio lines  970  are at the patient&#39;s pulse rate and associated harmonics, and another series of ratio lines  980  are at the active pulser rate and associated harmonics.  
         [0067]     Because only the arterial blood is pulsatile at the patient&#39;s pulse rate, the ratio lines  970  are only a function of the arterial oxygen saturation. Accordingly, Sp a O 2  can be derived from the magnitude of these ratio lines  970 , as described above. Further, a modulation level for the active pulse is selected which insignificantly perturbates the arterial blood while providing a measurable venous signal. This is possible because the arterial blood pressure is significantly larger than the venous pressure. The modulation level is regulated as described above with respect to  FIG. 2 , i.e. the DSP  150 , via a pulser drive control  291 , sets the magnitude of the pulser drive  280  to the pulse inducing mechanism  262 . Assuming that the active pulse modulation of the arterial blood is insignificant, only the venous blood is pulsatile at the active pulser rate. Hence, the ratio lines  980  are only a function of the venous oxygen saturation. Accordingly, Sp a O 2  can be derived from the magnitude of the pulse rate related ratio lines  980  in the same manner as Sp a O 2  is derived from the magnitude of the pulse rate related ratio lines.  
         [0068]     Scattering  
         [0069]     Propagation of optical radiation through tissue is affected by absorption and scattering processes. The operation of pulse oximeters was described qualitatively above using an analysis based on the Beer-Lambert law of absorption, equation (3). This approach, however, fails to account for the secondary effects of light scattering at pulse oximeter wavelengths. The primary light scatterer in blood is erythrocytes, i.e. red blood cells. A qualitative understanding of the effects of scattering on pulse oximetry is aided by a description of red blood cell properties within flowing blood.  
         [0070]     Human blood is a suspension of cells in an aqueous solution. The cellular contents are essentially all red blood cells, with white cells making up less the 1/600 th  of the total cellular volume and platelets less than 1/800 th  of the total cellular volume. Normally the hematocrit, which is the percentage of the total volume of blood occupied by cells, is about 50% in large vessels and 25% in small arterioles or venules.  
         [0071]     Red blood cells are extremely deformable, taking on various shapes in response to the hydrodynamic stresses created by flowing blood. For example, assuming a laminar blood flow within a vessel, a parabolic velocity profile exists that is greatest in the vessel center and smallest along the vessel walls. Nominally, red blood cells are shaped as biconcave disks with a diameter of 7.6 um and thickness of 2.8 um. Exposed to this velocity profile, the red blood cells become parachute-shaped and aligned in the direction of the blood flow. Thus, during systole, transmitted light is scattered by aligned, parachute-shaped cells. During diastole, the light is scattered by biconcave disks having a more or less random alignment.  
         [0072]     The time-varying shape and alignment of the red blood cells can have a significant effect on measured values of oxygen saturation if scattering is ignored. Analogous to the analysis using the Beer-Lambert absorption law, scattering can be qualitatively understood as a function of the scattering coefficients of various tissues. Specifically, the bulk scattering coefficient can be written as: 
 
μ s   =V   b μ b   +V   t μ t   (10) 
 
 where V b  is the blood volume, .μ b  is the scattering coefficient of blood, V t  is the surrounding tissue volume and .μ t  is the scattering coefficient of the surrounding tissue. The volume, V t , and scattering coefficient, μ t , of the surrounding tissue are time invariant. The blood volume, V b , however, is pulsatile. The ratio of ratios computation, RD/IR, results in normalization of the time invariant or DC tissue absorption and cancellation of the time varying or AC pulsatile blood volume absorption to yield a number related to oxygen saturation. This computational approach is valid because the absorption coefficients of blood, ε HbO2,λ , ε Hb,λ  given in equation (4) were assumed to change only slowly over time. The scattering coefficient of blood .μ b , however, is time variant. As described above, this variation is due to the time-varying alignment and shape of the red blood cells. This time variation in the detected intensity of light transmitted through a tissue site is not normalized or canceled by the RD/IR calculation. Further, because the magnitude of the scattering coefficient variations is a function of blood flow, these variations become more pronounced with larger pulses in the blood supply. As a result, scattering produces frequency-dependent magnitude variations in the ratio lines RD(ω)/IR(ω). 
 
         [0073]      FIG. 9  illustrates the effect of scattering on the spectra of the detected red and infrared intensity waveforms. When these waveforms are transformed into the frequency domain, the time varying component of scattering manifests itself as spreads  978 ,  988  in the RD/IR ratio lines at each harmonic of the plethysmograph or active pulse rate. The magnitude of the ratio lines  970  at the fundamental and harmonics of the patient&#39;s pulse rate varies between a minimum  972  and a maximum  974 , resulting in a magnitude spread  978 . Similarly, the magnitude of the ratio lines  980  at the fundamental and harmonics of the active pulse rate varies between a minimum  982  and a maximum  984 , resulting in a magnitude spread  988 . Normally, absent motion artifact or noise contamination, the spread  978 ,  988  in the ratio lines is quiet small, but the magnitude of these spreads  978 ,  988 , increases with larger blood flows or pulse magnitudes. Scattering attributable to an active pulse can be regulated by adjusting the magnitude of the active pulse modulation based upon the amount of spread  978 ,  988  of the ratio line magnitudes. Thus, the active pulse magnitude can be increased to obtain a larger detected AC signal, but limited to below the point at which scattering becomes significant.  
         [0074]      FIG. 10  depicts an embodiment of the signal processing for determining oxygen saturation from the ratio lines of RD(ω)/RD(ω). The red spectrum RD(ω)  612  and infrared spectrum IR(ω)  612 , computed as described above with respect to  FIG. 7 , are input to a peak detector  1010 . The peak detector  1010  separately calculates localized maximums for RD(ω) and IR(ω). The peak detector output  1012  is a series of frequencies corresponding to the patient pulse rate fundamental and harmonics. If an active pulse is used, the peak detector output  1012  is also a series of frequencies corresponding to the active pulse rate. Although the active pulse rate is known, the detected peaks may have been shifted due to noise, motion artifact or other signal contamination. The peak detector output  1012  is coupled to a series combination of peak matcher  1020  and ratio line calculator  1030 . The ratio lines RD/IR are calculated by matching the frequency peaks of RD(ω) with the nearest frequency peaks of IR(ω). The ratio lines associated with the pulse rate harmonics  1032  are then separated into a different set from the ratio lines associated with the active pulse harmonics  1034 , assuming an active pulse is utilized. An average ratio line for each set  1032 ,  1034  is calculated by averaging  1060  all ratio lines in a set. The magnitude of the average ratio line r  1062  for the pulse rate set  1032  is then fed to a look-up table (LUT)  1090 , which provides an output  622  of the measured value of Sp a O 2 . Similarly, if an active pulse is used, the magnitude of the average ratio line μ  1064  for the active pulse rate set  1034  is then fed to a LUT  1090 , which provides an output  622  of the measured value of Sp v O 2 . A scattering detector  1080  computes the spread  988  ( FIG. 9 ) in the set of ratio lines associated with the active pulse and provides this value  1082  to the DSP  150  ( FIG. 2 ) so that the DSP can set the pulser drive control  291  ( FIG. 2 ) to regulate the magnitude of the active pulse.  
         [0075]     Alternatively, Sp v O 2  may be measured from respiration-induced pulses in the venous blood, described above, without utilizing an active pulse sensor. Specifically, a series of ratio lines  980  ( FIG. 9 ) would occur at a fundamental frequency, f r0 , and associated harmonics, f r1  and f r2 , of the respiration rate, which is either known from the ventilator frequency or derived from a separate measurement of the natural respiration. As shown in  FIG. 10 , the ratio lines associated with the respiration rate harmonics  1034  are then separated into a different set from the ratio lines associated with the pulse rate harmonics  1032 . An average ratio line for the respiration rate set  1034  is calculated by averaging  1060  all ratio lines in that set. The magnitude of the average ratio line μ  1064  for the respiration rate set  1034  is then fed to a look-up table (LUT)  1090 , which provides an output  622  of the measured value of Sp v O 2 .  
         [0076]     Plethysmograph Feature Extractor  
         [0077]      FIG. 11  illustrates the standard plethysmograph waveform  1100 , which can be derived from a pulse oximeter. The waveform  1100  is a visualization of blood volume change in the illuminated peripheral tissue caused by arterial blood flow, shown along the y-axis  1110 , over time, shown along the x-axis  1120 . The shape of the plethysmograph waveform  1100  is a function of heart stroke volume, pressure gradient, arterial elasticity and peripheral resistance. The ideal waveform  1100  displays a broad peripheral flow curve, with a short, steep inflow phase  1130  followed by a 3 to 4 times longer outflow phase  1140 . The inflow phase  1130  is the result of tissue distention by the rapid blood volume inflow during ventricular systole. During the outflow phase  1140 , blood flow continues into the vascular bed during diastole. The end diastolic baseline  1150  indicates the minimum basal tissue perfusion. During the outflow phase  1140  is a dicrotic notch  1160 , the nature of which is disputed. Classically, the dicrotic notch  1160  is attributed to closure of the aortic valve at the end of ventricular systole. However, it may also be the result of reflection from the periphery of an initial, fast propagating, pressure pulse that occurs upon the opening of the aortic valve and that precedes the arterial flow wave. A double dicrotic notch can sometimes be observed, although its explanation is obscure, possibly the result of reflections reaching the sensor at different times.  
         [0078]      FIG. 12  is a graph  1200  illustrating the absorption of light at a tissue site illuminated by a pulse oximetry sensor. The graph  1200  has a y-axis  1210  representing the total amount of light absorbed the tissue site, with time shown along an x-axis  1220 . The total absorption is represented by layers including the static absorption layers due to tissue  1230 , venous blood  1240  and a baseline of arterial blood  1250 . Also shown is a variable absorption layer due to the pulse-added volume of arterial blood  1260 . The profile  1270  of the pulse-added arterial blood  1260  is seen as the plethysmograph waveform  1100  depicted in  FIG. 11 .  
         [0079]      FIG. 13  illustrates the photoplethysmograph intensity signal  1300  detected by a pulse oximeter sensor. A pulse oximeter does not directly detect absorption, and hence does not directly measure the standard plethysmograph waveform  1100  ( FIG. 11 ). However, the standard plethysmograph can be derived by observing that the detected intensity signal  1300  is merely an out of phase version of the absorption profile  1270 . That is, the peak detected intensity  1372  occurs at minimum absorption  1272  ( FIG. 12 ), and minimum detected intensity  1374  occurs at maximum absorption  1274  ( FIG. 12 ). Further, a rapid rise in absorption  1276  ( FIG. 12 ) during the inflow phase of the plethysmograph is reflected in a rapid decline  1376  in intensity, and the gradual decline  1278  ( FIG. 12 ) in absorption during the outflow phase of the plethysmograph is reflected in a gradual increase  1378  in detected intensity.  
         [0080]      FIG. 14  illustrates the digital signal processing for plethysmograph feature extraction  630  ( FIG. 6 ). The input  614  is the IR signal output from the demultiplexer  710  ( FIG. 7 ). This signal is shifted into a first-in, first-out (FIFO) buffer, which allows fixed-length portions of the input signal  614  to be processed for feature extraction. The buffered output signal  1412  is coupled to a shape detector  1420 , slope calculator  1430 , feature width calculator  1440  and a notch locator  1450 , which perform the core feature extraction functions. The shape detector  1420  determines if a particular buffered signal portion  1412  contains specific gross features, such as a peak, a valley, an upward slope, a downward slope, a dicrotic notch or a multiple dicrotic notch. A detected shape output  1422  containing one or more flags indicating the gross feature content of the current signal portion  1412  is coupled to the other feature extraction functions  1430 ,  1440 ,  1450  and to the waveform quality determination functions  1460 ,  1470 ,  1480 . A slope calculator  1430  determines the amount of positive or negative slope in the signal portion  1412  if the shape detector output  1422  indicates a slope is present. The output slope value  1432  is coupled to the waveform quality functions  1460 ,  1470 ,  1480  in addition to the feature extraction output  632 . A feature calculator  1440  quantifies a feature in one or more signal portions  1412  specified by the shape detector  1420 , such as the magnitude, the area under, or the width of a peak or notch. The feature calculator output  1442  is a code indicating the feature and its value, which is coupled to the feature extraction output  632 . A feature locator  1450  quantifies the time of occurrence of one or more features of a signal portion  1412  as specified by the shape detector  1420 . The feature locator output  1452 , which is coupled to the feature extraction output  632 , is a code indicating a feature and an associated code indicating time of occurrence in reference to a particular epoch. The feature locator output  1452  allows a determination of the relative location of plethysmograph features in addition to a phase comparison of plethysmographs derived from two or more tissue sites. Another feature extraction output  634 , which is coupled to the multiple parameter processor  640  ( FIG. 6 ), provides an indication of waveform quality. Input signals portions  1412  not having either a sharp downward edge  1460 , a symmetrical peak  1470  or a gradual decline  1480  are not processed further.  
         [0081]     Multiple Parameter Processor  
         [0082]      FIG. 15  illustrates the multiple parameter processing portion  640  ( FIG. 6 ) of the signal processing. A differencing function  1510  has as inputs a first saturation value, Sp 1 O 2 , and a second saturation value, Sp 2 O 2 ,  622 . The saturation input values  622  can be arterial and venous saturation values from a single data channel, arterial saturation values from two different data channels or venous saturation values from two different data channels. The differences of the saturation value inputs  622  are provided as an output  1514 , which is coupled to a saturation data memory  1520 . The saturation values  622  are also directly coupled to the saturation data memory  1520 . The memory  1520  stores a record of saturation values, SpO 2 , for each channel, delta saturation values, Δsat, for each channel and cross-channel delta saturation values, Δsat xy , as required for a particular application. A flow calculator  1530  utilizes a plethysmograph input  614  or a bio-impedance probe input  1534  to provide a flow value  1538 , which is also coupled to the saturation data memory  1520 . For example, the flow value  1538  may be a perfusion index, PI, defined as follows:  
             PI   =         IR   max     -     IR   min         IR   DC               (   11   )               
 where IR max  is the maximum value, IR min  is the minimum value, and IR DC  is the average value of the IR plethysmograph signal  614  ( FIG. 7 ). 
 
         [0083]     The saturation data memory  1520  provides a buffered output  1522  that is coupled to a numerical display driver  1540 . The numerical display driver  1540  provides an output  642  to a standard display, such as LED or LCD numerical display modules or a CRT monitor. The memory output  1522  is also coupled to a saturation data analyzer  1530 , one function of which calculates a long-term trend of the values in memory  1520 . For example, the saturation data analyzer may average a saturation value over time, or provide samples of the saturation values taken at regular time intervals. The output  1532  can either be numerical, which is coupled to the numerical display driver  1540 , or graphical, which is coupled to the graphical display driver  1570 . The graphical display driver  1570  provides an output  644  to a standard graphical display device, such as LED or LCD graphical display modules or a CRT monitor.  
         [0084]     A pleth data memory  1550  has as inputs the IR plethysmograph signals  614  ( FIG. 7 ) from each data channel and the associated extracted features  632  ( FIG. 14 ). The memory  1550  also has an input indicating waveform quality  634  ( FIG. 14 ). The pleth memory  1550  provides a buffered output  1558  that is coupled to the graphical display driver  1570 , allowing display of the plethysmograph waveforms for each data channel. The memory output  1558  is also coupled to a pleth data analyzer  1560 , one function of which calculates a long-term trend of the plethysmograph and shape parameters in pleth memory  1520 . For example, the pleth data analyzer  1560  may provide an average of particular shape parameters over time. As another example, the pleth data analyzer  1560  may provide a graphic showing an accumulation of many overlaid plethysmographs. The output  1562  can either be numerical, which is coupled to the numerical display driver  1540 , or graphical, which is coupled to the graphical display driver  1570 .  
         [0085]     Another function of the saturation data analyzer  1530  and the pleth data analyzer  1560  is to compare oxygen status and plethysmograph parameters derived from multiple sites in order to isolate noise artifacts and to derive a more accurate estimate of these parameters. For example, it is unlikely that motion artifact will affect each peripheral site in the same manner. If the quality input  634  indicates a noisy plethysmograph for one channel during a particular time period, the pleth data analyzer  1560  can exchange this information  1565  with the saturation data analyzer  1530 . The saturation data analyzer  1530  can then ignore the saturation data for that channel for that time period in lieu of saturation data from another channel. In a similar fashion, noisy data from multiple channels can be averaged, correlated or otherwise processed to provide an estimate of Sp a O 2 , Sp v O 2  or pulse rate, or to provide a plethysmograph that is more accurate than can be derived from a single data channel.  
         [0086]      FIG. 16A  illustrates detail of a single-site display screen  180  for the stereo pulse oximeter. The display has a numerical display portion  1610  controlled by the numerical display driver  1540  ( FIG. 15 ) and a graphical display portion  1660  controlled by the graphical display driver  1570  ( FIG. 15 ). The numerical display portion  1610  displays a value for Sp a O 2    1620 , Sp v O 2    1630  and pulse rate  1640  for a particular tissue site. The graphical display portion  1660  displays a plethysmograph  1662  for the corresponding tissue site, which can be displayed as a single waveform or an accumulated multiple of overlayed waveforms that may reveal a waveform trend. A push button or menu selection allows the user to switch to a display of data from any single one of the multiple tissue sites to which a sensor is attached.  
         [0087]      FIG. 16B  illustrates detail of a multi-site display screen  180  for the stereo pulse oximeter. The numerical display portion  1610  displays a value for Sp a O 2    1622  and Sp v O 2    1632  for a first tissue site. Also displayed is a value for Sp a O 2    1624  and Sp v O 2    1634  for a second tissue site. In addition, a value for pulse rate  1642  derived from either the first or second tissue site, or both, is displayed. The graphical display portion  1660  displays a first plethysmograph  1664  and a second plethysmograph  1666  corresponding to the first and second tissue sites, respectively. A push button, menu selection allows the user to manually switch between the single site display ( FIG. 16A ) and the multi-site display ( FIG. 16B ). Also, a triggering event, such as an alarm based on multiple-site oxygen status parameters, causes the display to automatically switch from the single-site display to the multi-site display, enabling the user better view the conditions that caused the triggering event.  
         [0088]     One of ordinary skill will appreciate many display screens variations from those shown in  FIGS. 16A and 16B  that are within the scope of this invention. For example, the stereo pulse oximeter could be configured to provide several push button or menu selectable display screens. One display screen might display more than two channels of oxygen status data. Another display screen could display cross-channel parameters such as Δsat xy  or a comparison of plethysmograph shape parameters from two channels. One of ordinary skill will also appreciate many variations and modifications of layout and design for the graphical and numerical displays within the scope of this invention.  
         [0089]     Stereo Pulse Oximetry Applications  
         [0090]     Oxygen Titration  
         [0091]     Oxygen is one of the most commonly used drugs in an intensive care unit and is an integral part of all respiratory support. The goal of oxygen therapy is to achieve adequate delivery of oxygen to the tissues without creating oxygen toxicity. Too little oxygen results in organ damage and, in particular, brain damage. Too much oxygen can result in, for example, pulmonary edema and, in neonates, retinopathy of prematurity (ROP). Infants receiving oxygen therapy, in particular, must have inspired oxygen concentration and blood oxygen levels monitored closely.  
         [0092]     Oxygen titration in neonates is currently accomplished with either transcutaneous monitoring or monitoring with a conventional pulse oximeter. As mentioned above, transcutaneous monitoring involves the placement of a heated Clark electrode against the skin surface. The electrode is secured to the skin surface with an airtight seal to eliminate contamination by room air gases. The skin surface beneath the electrode is then heated, which opens pre-capillary sphincters allowing localized arteriolar blood flow beneath the sensor. The so-called T c O 2  value that is measured correlates well with P a O 2 . However, there are several drawbacks to this approach. Because the skin surface must be heated, a fifteen minute elapsed time after application is necessary before stable readings are acquired. Further, the required temperature is 43-45° C. (110° F.), with an associated risk of burns. In addition, titration is often accomplished by simply maintaining T c O 2  within acceptable limits for this parameter, e.g. an equivalent P a O 2  of 50-80 mm Hg for neonates. However, P a O 2  alone does not provide an indication of balance between inspired oxygen and the rate of tissue oxygen consumption. If the patient is particularly anemic or hypovolemic, has an abnormal hemoglobin, or a small cardiac output, then oxygen delivery may be inadequate even in the presence of a normal P a O 2 . Titration with a conventional pulse oximeter is similarly accomplished by maintaining Sp a O 2  within acceptable limits, which also fails to consider tissue oxygen consumption.  
         [0093]     Oxygen titration can be more adequately monitored with a continuous indication of oxygen consumption, which is equal to oxygen delivery according to Fick&#39;s algorithm, as noted above. Further, continuous monitoring of oxygen consumption at a peripheral tissue site, although not necessarily indicative of overall oxygen consumption, may be indicative of an oxygen supply dependency. A measure of peripheral oxygen consumption can be expressed in terms of Δsat=Sp a O 2 −Sp v O 2  and perfusion, which, as noted above, are parameters advantageously provided by the stereo pulse oximeter according to the present invention. Oxygen consumption at a peripheral site is obtained by multiplying the difference between peripheral arterial and venous oxygen content by perfusion at the site. 
 
VpO 2 =[O 2  content (arterial)−O 2  content (venous)]Φ  12) 
 
 where oxygen content is measured in milliliters (ml) of O 2  per deciliters (dl) of blood and Φ denotes perfusion in deciliters per minute. Oxygen content, however, can be expressed in terms of the amount of oxygen bound to the hemoglobin plus the amount of oxygen dissolved in the plasma. The amount of bound oxygen is equal to the hemoglobin concentration, C hb , in grams per deciliter of blood, times the hemoglobin carrying capacity, which is 1.34 milliliters of O 2  per gram of hemoglobin times the hemoglobin oxygen saturation, SO 2 . The amount of dissolved oxygen is simply the partial pressure of oxygen, PO 2 , times the O 2  solubility coefficient in blood, which is 0.003 milliliters of O 2  per deciliter. The sum of these two terms yields: 
 
O 2  content=1.34C Hb SO 2 +0.003PO 2   (13) 
 
 Substituting equation (13) into equation (12) yields the following equation for tissue oxygen consumption: 
 
VpO 2 =[1.34C Hb (Sp a O 2 −Sp v O 2 )+0.003(P a O 2 -P v O 2 )]φ (14) 
 
 Except when the fractional inspired oxygen, FiO 2 , is high, blood plasma plays a minimal role in oxygen delivery. Thus, peripheral oxygen consumption is approximately: 
 
VpO 2 =[1.34C Hb Δsat]Φ  (15) 
 
         [0094]     In order to illustrate a schema of oxygen titration, it is convenient to characterize the relationship between oxygen supplied at the airway to oxygen consumed at a peripheral tissue site. Specifically, characterization of the relationship between Δsat, Φ and FiO 2  is useful. Assuming constant oxygen consumption at the tissue site, equation (15) is: 
 
Δsat Φ=constant  (16) 
 
 Equation (16) has a simple analog in electronic circuits, i.e. a variable resistor across a current or voltage source adjusted to maintain constant power. In this analog circuit, the current through the resistor, I, is equivalent to perfusion, the voltage across the resistor, V, is equivalent to Δsat and the constant of equation (16) is equivalent to the constant power, P, consumed by the resistor. The equation representing this electrical analog is: 
 
 V×I=P   (17) 
 
         [0095]      FIG. 17A  shows a graph  1701  that depicts a family of curves each corresponding to different values of P in equation (17). The graph  1701  has an x-axis  1710  indicating current, I, and a y-axis  1720  indicating voltage, V. A first curve  1730  shows V versus I for a constant power, P, of 0.5 watts; a second curve  1740  shows V versus I for a constant P of 1 watt; and a third curve  1750  shows V versus I for a constant P of 2 watts. Using the analogy between equations (16) and equation (17), whenever Φ (current) is small, the Δsat (voltage) is large and vice-a-versa. Also, a change in consumption (power) causes a shift in the curve along with a change in its curvature. That is, if the body suddenly changes its metabolic rate at the peripheral tissue site, the curve will accordingly shift up or shift down and will change its shape. Equation (16) and the analogous constant consumption curves of  FIG. 17A  assume a supply independent condition, i.e. that peripheral oxygen consumption is satisfied by peripheral oxygen delivery. If the peripheral tissue site is starved for oxygen, then the locus of points for Δsat versus Φ is quite different from a hyperbola The amount of tissue oxygen extraction is at a maximum and is independent of Φ. Accordingly Δsat is at a maximum and independent of Φ. The above analysis provides insight into the relationship between Δsat and Φ. The relationship between Δsat and FiO 2  can also be characterized.  
         [0096]      FIG. 17B  shows a graph  1702  of saturation along a y-axis  1760  and fractional inspired oxygen along an x-axis  1770 . A curve of Sp a O 2    1780  and a curve of Sp a O 2    1790  are depicted versus FiO 2 . The difference between these curves  1780 ,  1790  yields Δsat  1785  versus FiO 2 . When FiO 2  is zero  1772 , oxygen saturation and, hence, both Sp a O 2    1780  and Sp v O 2    1790  are zero. As FiO 2  is increased, Sp a O 2    1780  also increases until virtually reaching 100 percent saturation  1762 . As FiO 2  increases further, Sp a O 2    1780  stays at virtually 100 percent saturation  1762 . As FiO 2  is increased from zero  1772 , Sp v O 2    1790  also increases. In this low FiO 2  region  1774 , the peripheral tissue site is supply dependent and Δsat  1785  also increases. At a certain point, the tissue site oxygen demand is met by supply. In this supply independent region  1776 , oxygen consumption is constant and equation (16) is valid. Also, Δsat  1785  is at a constant maximum, which is a function of the metabolism at the tissue site. As FiO 2  increases further, eventually the partial pressure of oxygen becomes significant and the second term of equation (14) must be considered. In this high FiO 2  region  1778 , Δsat  1785  decreases because some of the tissue oxygen consumption is supplied by oxygen dissolved in the plasma.  
         [0097]      FIG. 17C  shows a graph  1704  of saturation difference along a y-axis  1764  and fractional inspired oxygen along an x-axis  1770 . A curve of Δsat  1786  is depicted versus FiO 2 , corresponding to the region Δsat  1785  depicted in  FIG. 17B . The curve  1786  has a first deflection point  1766  occurring at the transition between the low FiO 2  region  1774  ( FIG. 17B ) and the supply independent region  1776  ( FIG. 17B ). The curve  1786  also has a second deflection point  1768  occurring at the transition between the supply independent region  1776  ( FIG. 17B ) and the high FiO 2  region  1778  ( FIG. 17B ). The curve  1786  illustrates how the trend for Δsat, as measured by the stereo pulse oximeter, can be used to accurately titrate oxygen. The goal of oxygen titration is to supply sufficient oxygen to supply tissue demand and avoid unnecessarily high amounts of FiO 2 . Thus, the Δsat parameter should be monitored so that FiO 2  is adjusted between the two deflection points  1766 ,  1768 . For neonates, FiO 2  should be adjusted just beyond the first deflection point  1766 . For adults, FiO 2  should be adjusted just before the second deflection point  1768 .  
         [0098]      FIG. 18  illustrates a graph having a three-dimensional surface  1800  generally depicting the relationship between Δsat, Φ and FiO 2  from the combined graphs of  FIGS. 17A and 17C . The graph has an x-axis  1810  showing FiO 2 , a y-axis  1820  showing Φ and a z-axis  1830  showing Δsat. The surface  1800  has a supply dependent region  1840 , a perfusion-limited region  1850 , a constant consumption region  1860  and a plasma dependent region  1870 . The surface describes the oxygen status of a peripheral tissue site. The supply dependent region  1840  corresponds to the low FiO 2  region  1774  ( FIG. 17B ) described above. That is, inspired oxygen into the lungs is so low that, at the tissue site, oxygen extraction by the tissues is limited by oxygen delivery, and Δsat falls rapidly as FiO 2  is reduced. The perfusion-limited region  1850  along the x-axis  1810  represents a low perfusion state where equation (16) is not valid. That is, perfusion at the tissue site is so low that oxygen extraction by the tissues is at a maximum, and, hence, Δsat is at a maximum and is independent of FiO 2 . A cross-section of the surface taken parallel to the y-axis  1820  yields a hyperbole-shaped constant consumption region  1860 , consistent with the constant metabolic rate curves illustrated above with respect to  FIG. 17A . The plasma dependent region  1870  corresponds to the high FiO 2  region  1778  ( FIG. 17B ) described above. That is, inspired oxygen into the lungs is so high that the tissue site is partially dependent on oxygen dissolved in the plasma The surface  1800  illustrates that perfusion should be monitored simultaneously with Δsat to avoid the perfusion-limited region  1850 , where Δsat is an unresponsive indicator of FiO 2 , and to avoid misinterpreting hyperbolic changes in Δsat that result from changes in perfusion.  
         [0099]     Persistent Pulmonary Hypertension in Neonates  
         [0100]      FIG. 19  illustrates the heart/lung circulation of a hypertensive neonate. Persistent Pulmonary Hypertension in Neonates (PPHN) is a neonatal condition with persistent elevation of pulmonary vascular resistance and pulmonary artery pressure. Shown is a neonatal heart  1902  and a portion of a neonatal lung  1904 . The pulmonary artery  1910  that normally feeds oxygen depleted “blue” blood from the right ventricle  1920  to the lung  1904  is constricted. The back pressure from the constricted artery  1910  results in a right-to-left shunting of this oxygen depleted blood through the ductus arteriosus  1930 , causing it to mix with oxygen rich “red” blood flowing through the descending aorta  1940 . PPHN treatment options include vasodilators, such as nitric oxide (NO). Inhaled exogenous NO causes a dose-dependent decrease in pulmonary artery pressure and pulmonary vascular resistance, as well as a parallel increase in pulmonary blood flow, without affecting systemic arterial pressure. However, the response to NO therapy is a function of the cause of the PPHN as well as the time elapsed before initiation of therapy. Potential toxic effects of NO dictate the proper titration of NO gas. Too little NO may not effectively relieve pulmonary hypertension, and too much NO may cause cellular injury or toxicity. NO therapy is currently monitored using intermittent ultrasound imaging and/or in vitro blood gas measurements. The drawbacks to these techniques are noncontinuous monitoring and disturbances to the neonate that can exacerbate or not reflect the hypertension in the non-disturbed state.  
         [0101]     The stereo pulse oximeter according to the present invention allows noninvasive, continuous monitoring of a neonate for detection and managed treatment of PPHN that does not disturb the patient. A right hand sensor  130  ( FIG. 1 ) provides arterial oxygen saturation and a plethysmograph for blood circulating from the left ventricle  1950  through the innominate artery  1960 , which supplies the right subclavian artery. Because the innominate artery  1960  is upstream from the shunt at the ductus arteriosus  1930 , the oxygen saturation value and plethysmograph waveform obtained from the right hand are relatively unaffected by the shunt and serve as a baseline or reference for comparison with readings from other tissue sites. Alternatively, a reference sensor can be placed on a facial site, such as an ear, the nose or the lips. These sites provide arterial oxygen saturation and a plethysmograph for blood circulating from the left ventricle  1950  to the innominate artery  1960 , which supplies the right common carotid artery (not shown), or to the left common carotid artery  1965 .  
         [0102]     A foot sensor  120  ( FIG. 1 ) provides oxygen status for blood supplied from the descending aorta  1940 . The shunt  1930  affects both the oxygen saturation and the blood flow in the descending aorta  1940 . As stated above, the shunt  1930  causes oxygen-depleted blood to be mixed with oxygen-rich blood in the descending aorta  1940 . Because the descending aorta  1940  supplies blood to the legs, the oxygen saturation readings at the foot will be lowered accordingly. The PPHN condition, therefore, is manifested as a higher arterial oxygen saturation at the right hand reference site and a lower saturation at the foot site.  
         [0103]     The shunt also allows a transitory left to right flow during systole, which distends the main pulmonary artery  1980  as the result of the blood flow pressure at one end from the right ventricle and at the other end from the aortic arch  1990 . A left-to-right flow through the shunt  1930  into the distended artery  1980  alters the flow in the descending aorta  1940  and, as a result, the plethysmograph features measured at the foot. The PPHN condition, therefore, also is manifested as a plethysmograph with a narrow peak and possibly a well-defined dicrotic notch at the left hand baseline site and a broadened peak and possibly no notch at the foot site.  
         [0104]     An optional left hand sensor  110  ( FIG. 1 ) provides oxygen status for blood circulating from the left ventricle through the left subclavian artery  1970  that supplies the left arm. Because the left subclavian artery  1970  is nearer the shunt  1930  than the further upstream innominate artery  1960 , it may experience some mixing of deoxygenated blood and an alteration in flow due to the shunt  1930 . The PPHN condition, therefore, may also be manifested as a reduced saturation and an altered plethysmograph waveform at the left hand site as compared with the right hand baseline site, although to a lesser degree than with a foot site. Thus, the PPHN condition can be detected and its treatment monitored from Δsat and plethysmograph morphology comparisons between a right hand baseline sensor site and one or more other sites, such as the left hand or foot.  
         [0105]     Patent Ductus Arteriosus  
         [0106]      FIG. 20  illustrates the fetal heart/lung circulation. Shown is a fetal heart  2002  and a portion of a fetal lung  2004 . The lung  2004  is non-functional and fluid-filled. Instead, oxygenated blood is supplied to the fetus from gas-exchange in the placenta with the mother&#39;s blood supply. Specifically, oxygenated blood flows from the placenta, through the umbilical vein  2006  and into the right atrium  2022 . There, it flows via the foramen  2024  into the left atrium  2052 , where it is pumped into the left ventricle  2050  and then into the aortic trunk  2092 . Also, oxygenated blood is pumped from the right atrium  2022  into the right ventricle  2020  and directly into the descending aorta  2040  via the main pulmonary artery  2080  and the ductus arteriosus  2030 . Normally, the ductus arteriosus  2030  is only open (patent) during fetal life and the first 12 to 24 hours of life in term infants. The purpose of the ductus arteriosus  2030  is to shunt blood pumped by the right ventricle  2020  past the constricted pulmonary circulation  2010  and into the aorta  2040 .  
         [0107]      FIG. 21  illustrates a neonatal heart  2002  with a patent ductus arteriosus  2030 . The ductus arteriosus frequently fails to close in premature infants, allowing left-to-right shunting, i.e. oxygenated “red” blood flows from the aorta  2040  to the now unconstricted pulmonary artery  2010  and recirculates through the lungs  2004 . A persistent patent ductus arteriosus (PDA) results in pulmonary hyperperfusion and an enlarged right ventricle  2020 , which leads to a variety of abnormal respiratory, cardiac and genitourinary symptoms. Current PDA diagnosis involves physical examination, chest x-ray, blood gas analysis, echocardiogram, or a combination of the above. For example, large PDAs may be associated with a soft, long, low-frequency murmur detectable with a stethoscope. As another example, two-dimensional, color Doppler echocardiography may show a retrograde flow from the ductus arteriosus  2030  into the main pulmonary artery  2080 . Once a problematic PDA is detected, closure can be effected medically with indomethacin or ibuprofen or surgically by ligation. Multiple doses of indomethacin are commonplace but can still result in patency, demanding ligation. A drawback to current diagnostic techniques is that clinical symptoms of a PDA can vary on an hourly basis, requiring extended and inherently intermittent testing.  
         [0108]     The stereo pulse oximeter according to the present invention allows for continuous evaluation of PDA symptoms using non-invasive techniques. A right hand sensor  130  ( FIG. 1 ) provides arterial oxygen saturation and a plethysmograph for blood circulating from the left ventricle  2050  through the innominate artery  2160 , which supplies the right subclavian artery leading to the right arm. Because the innominate artery  2160  is upstream from the shunt at the ductus arteriosus  2030 , the oxygen saturation value and plethysmograph waveform obtained from the right hand are relatively unaffected by the shunt and serve as a baseline for comparison with readings from other tissue sites.  
         [0109]     A foot sensor  120  ( FIG. 1 ) provides oxygen status for blood supplied from the descending aorta  2040 . Unlike a PPHN condition, the shunt  2030  does not affect oxygen saturation in the descending aorta  2040 , because the relatively low pressure in the pulmonary artery  2010  does not allow a mixing of deoxygenated blood into the relatively high pressure flow of oxygenated blood in the aorta  2040 . However, like a PPHN condition, the shunt  2030  does affect the aortic flow. In particular, the shunt allows a transitory left-to-right flow during systole from the high pressure aorta  2040  to the low pressure pulmonary circulation  2010 . This left-to-right flow through the shunt  1930  alters the flow in the descending aorta  1940  and, as a result, the plethysmograph features measured at the foot. The PDA condition, therefore, is manifested as a normal plethysmograph with a characteristically narrow peak and well-defined dicrotic notch at the right-hand baseline site compared with a damped plethysmograph with a broadened peak and reduced or missing notch at the foot site. Further, the foot site waveform is phase shifted from the baseline waveform. These plethysmograph differences are accompanied by comparable arterial oxygen saturation values between the right-hand site and the foot site.  
         [0110]     An optional left hand sensor  110  ( FIG. 1 ) provides oxygen status for blood circulating from the left ventricle through the left subclavian artery  2170  that supplies the left arm. Because the left subclavian artery  2170  is nearer the shunt  2030  than the further upstream innominate artery  2160 , it may experience some alteration in flow due to the shunt  2030 . The PDA condition, therefore, may also be manifested as an altered plethysmograph waveform at a left hand site as compared with the right hand baseline site, although to a lesser degree than with a foot site. Thus, the PDA condition can be detected and its treatment monitored from Δsat xy ≈0 and plethysmograph morphology and phase comparisons between a right hand baseline sensor site and one or more other sites, such as the left hand or foot. One of ordinary skill will recognize that multiple site comparisons using the stereo pulse oximeter of the current invention may also be used to detect other cardiac abnormalities that cause mixing of oxygenated and deoxygenated blood, such as a ventricular hole or a patent foramen. Further, abnormal mixing of oxygenated and deoxygenated blood may also be manifested in measurements provided by the stereo oximeters other than Δsat xy  and .Δpleth xy  as described above. For example, an inversion in Δsat at a particular tissue site, i.e., Sp v O 2  being larger than Sp a O 2  at that site, would indicate such an abnormal condition.  
         [0111]     Aortic Coarctation  
         [0112]     Coarctation of the aorta is a congenital cardiac anomaly in which obstruction or narrowing occurs in the distal aortic arch or proximal descending aorta. It occurs as either an isolated lesion or coexisting with a variety of other congenital cardiac anomalies, such as a PDA. If the constriction is preductal, lower-trunk blood flow is supplied predominantly by the right ventricle via the ductus arteriosus, and cyanosis, i.e. poorly oxygenated blood, is present distal to the coarctation. This can be detected by the stereo pulse oximeter from a comparison of Sp a O 2  between an upper body and a lower body site. If the constriction is postductal, blood supply to the lower trunk is supplied via the ascending aorta Differential plethysmographs between the upper and lower extremities may not exist if the ductus is widely patent. If the ductus closes, however, this condition can be detected by the stereo pulse oximeter as a reduced amplitude and phase delay between the plethysmographs measured at a lower body site with respect to an upper body site.  
         [0113]     The stereo pulse oximeter has been disclosed in detail in connection with various embodiments of the present invention. These embodiments are disclosed by way of examples only and are not to limit the scope of the present invention, which is defined by the claims that follow. One of ordinary skill in the art will appreciate many variations and modifications within the scope of this invention.