fig1 illustrates the invention in detail . a portion of bodily tissue 10 , such as the web of the hand between the thumb and index fingers , the forehead or any other portion of the skin , is illuminated by a light beam generated by two sources 61 and 62 of equal or unequal intensity and respective wavelengths λ1 and λ2 . both wavelengths are selected so as to be substantially insensitive to variations in oxygen saturation or any other blood constituent other than hematocrit or hemoglobin content . the wavelengths 815 nm and 950 nm are preferred examples . light sources l1 and l2 may comprise fixed or tuneable lasers or led &# 39 ; s . the light from sources l1 and l2 is time multiplexed in multiplexer 14 under control of timing switch s1 to produce a single beam of light comprising time multiplexed pulses λ1 and λ2 . the beam from multiplexer 14 may be butt - coupled to the skin 24 of tissue 10 or may be coupled to the skin by optical fiber f1 , as shown . the light beam undergoes scattering and absorption in the tissue according to certain path a or b determined by the constituents in the tissue . ultimately the light beam exits the tissue either by transmission , transflection or reflection and is detected by detectors d1 and d2 which are shown located for reflectance or transflectance measurement . by placing the two photodetectors d1 and d2 on or near the surface of the skin 24 at two different locations corresponding to a shorter pathlength a and a longer pathlength b , it is possible to detect two photoplethysmographic waveforms λ1 ( a ) & amp ; λ2 ( a ) and λ1 ( b ) & amp ; λ2 ( b ) corresponding to the light intensities emanating from the tissue 10 , respectively . note : if the detectors are not disposed on the skin , optical fibers f2 and f3 may be used to couple the received light to a respective detector d1 or d2 . the detected waveforms are demultiplexed in demultiplexer 26 which is comprised of an a - demultiplexer 26a and a b - demultiplexer 26b to which the respective diode detector outputs are coupled . the demultiplexed signals λ1 ( a ) and λ2 ( a ) travelling the a - path are coupled to signal processing a - channel . the demultiplexed signals travelling the b - path through tissue 10 ( λ1 ( b ) and λ2 ( b )) are coupled to signal processing b - channel . the waveforms demultiplexed in demultiplexers 26a and 26b are synchronized with multiplexer 14 by switch s1 . the relative intensities of the four photoplethysmographic waveforms λ1 ( a ) & amp ; λ2 ( a ) and λ1 ( b ) & amp ; λ2 ( b ) is proportional to the separation distances of the two photodetectors d1 and d2 , respectively , from the point of entry of the beam into the tissue . an optical shield 12 is placed between the light source 14 ( or fiber optic f1 ) and the two photodetectors d1 and d2 , ( or fiber optics f1 & amp ; f2 ), in order to eliminate the possibility of detecting light directly coupled between the light source and the photodetectors . the light intensity absorbed by the tissue 10 is modulated by the change in arterial blood volume each time the heart contracts . fig2 is a plot of the intensity detected by one of the detectors versus time after it has been time demultiplexed . when more blood is present in the tissue during peak systole ( point 2 ), more light will be absorbed and less light will be detected by the photodetector . likewise , during peak diastole ( point 3 ), less blood will be present in the tissue , and the photodetectors will detect a stronger signal . the relative magnitude of this signal , which is given by the peak - to - peak value of the photoplethysmographic waveform 4 , is a function of the incremental amount of blood that enters the tissue during systole with each heart beat . in addition to the time variant ( ac ) component of the photoplethysmogram , which is a function of the incremental increase in blood volume during systole , there is also a time - invariant ( dc ) component ( 5 ) in the photoplethysmogram . this dc component corresponds to the average amount of blood present in the tissue between heart beats and is also proportional to the amount of light io entering the tissue . the detected optical photoplethysmogram waveforms λ1 ( a ) & amp ; λ2 ( a ) and λ1 ( b ) & amp ; λ2 ( b ) after conversion into electrical intensity signals by respective detectors d1 and d2 and demultiplexing in 26a and 26b are separated into ac waveform components ac ( a ) and ac ( b ) by blocking capacitors c1 and c2 . these ac components are subsequently converted into respective dc voltages corresponding to the amplitude of the photoplethysmographic signals by respective rectifiers and ( if necessary amplifiers ) r1 and r2 . the dc components dc ( a ) and dc ( b ) are coupled directly to respective normalization circuits n1 and n2 along with the outputs from r1 and r2 . the magnitude of each photoplethysmogram is normalized by dividing its ac component by its dc component in the normalization circuits n1 and n2 . this cancels out the effect of variations in the output light intensity of the light sources used to illuminate the tissue 10 . by calculating the ratios between the normalized magnitudes of the two photoplethysmograms detected by the two differently spaced photodetectors in ratio circuits 16a and 16b , a relationship is obtained which provides information on the total hemoglobin content present in the tissue and hematocrit . the ratio between a 1 and b 1 ( one pair of outputs of the normalization circuit n 1 and n 2 ) is computed by the ratio circuit 16a and this value , which is computed for λ1 ( or 850 nm ), is empirically correlated against hemoglobin content to provide a mathematical relationship for predicting hemoglobin content values . in like manner , the de - multiplexed λ2 ( a ) and λ2 ( b ) signals for the second wavelength λ2 are then processed in the same manner in channels a and b , respectively , to produce a second set of normalized photoplethysmograms a2 and b2 . the second wavelength , λ2 ( for example 950 nm ) is not only sensitive to variations in hemoglobin , but is also sensitive to variations in the plasma content in the tissue . this wavelength may be used to obtain hematocrit information , as indicated above . the scattered and absorbed light from λ2 is also detected in detectors d1 and d2 and split into dc and ac components ac ( b ) and dc ( b ) by rectifying amplifying and filtering , as above in r1 and c1 and r2 and c2 . this process provides information on the optical absorption of tissue as a function of the total light absorbed by both hemoglobin and plasma . the ratio of the normalized ac / dc components a2 / b2 of each photoplethysmogram detected at respective wavelength λ2 is empirically correlated against known hematocrit values to find a mathematical relationship for predicting hematocrit values . it is thus possible to obtain a quantitative measure of hemocrit by computing the relative concentrations of hemoglobin content to plasma content . data processor 20 has stored in it the aforesaid mathematical relationship to provide an absolute indication of blood hematocrit and hemoglobin content noninvasively . these values may be displayed on display 22 . an example of a mathematical relationship which could be stored in the data processor 20 is as follows : ## equ1 ## where k 0 , k 1 , k 2 , k 3 , k 4 and k 5 are six empirically determined regression coefficients . these coefficients can be determined by a calibration study in which ## equ2 ## are measured in different patients having different known hemoglobin content and hematocrit values . these values may have been obtained by prior in vivo or in vitro tests . after these six coefficients are determined , they are permanently programmed into the data processor ( 20 ) and used to predict hemoglobin content and hematocrit by measuring the values of ## equ3 ## in the patient under test . those skilled in the art will recognize , or be able to ascertain , using no more than routine experimentation , many equivalents to the specific embodiments of the invention described herein . the invention is not limited to the two wavelengths ( 815 and 950 nm ). it is only necessary that one wavelength be used that is not absorbed by the plasma but is absorbed by the hemoglobin ( this can be done by the 815 nm wavelength ) for computing hemoglobin content . for computing hematocrit , a second wavelength is required that is absorbed by both hemoglobin and plasma ( such as 950 nm ). an alternate embodiment of a single wavelength embodiment invention is shown in fig3 wherein components common to fig1 retain the same numerical designation . in this embodiment only a single wavelength λ1 is needed in order to generate the two ratio signals a1 / b1 and a2 / b2 processed in processor 20 . otherwise , the function and operation are as previously discussed in connection with fig3 . rather than multiplexing and demultiplexing the illuminating signals at wavelengths λ1 and λ2 , two sets of input optics and processing optics and detectors and processor channels may be operated in parallel . these and all other equivalents are intended to be encompassed by the following claims .
the present invention relates to the noninvasive measurement of blood hematocrit and hemoglobin content using differential optical absorption of two or more wavelengths of light during blood volume changes . the method is also useful for noninvasive measurements of other blood analytes , such as glucose , where variations in hematocrit or blood hemoglobin concentration cause errors in the measurement .