referring to the figures , wherein like numerals indicate like or corresponding parts throughout the several views , fig1 and 2 illustrate typical hemoglobin spectra and calibrations curve utilized in the pulse oximetry measurements . the present invention provides a sensor for use in a reflection - mode or backscatter type pulse oximeter . the relative disposition of light source and detector in the reflection - mode pulse oximeter are illustrated in fig3 . [ 0072 ] fig4 shows light propagation in the reflection - mode pulse oximeter where , in addition to the optical absorption and reflection due to blood , the dc signal of the r and ir photoplethysmograms can be adversely affected by strong reflections from the bone . [ 0073 ] fig5 a and 5b illustrate a pulse oximeter reflectance sensor operating under , respectively , ideal and practical conditions . referring now to fig5 a , it is shown that , under ideal conditions , reflectance sensor measures light backscattered from a homogenous mixture of blood and bloodless tissue components . accordingly , the normalized r / ir ratio in dual - wavelength reflection type pulse oximeters , which relies on proportional changes in the ac and dc components in the photoplethysmograms , only reflect changes in arterial oxygen saturation . referring now to fig5 b , in practical situations , the sensor applications affect the distribution of blood in the superficial layers of the skin . accordingly , the r and ir dc signals measured by photodetectors contain a relatively larger proportion of light absorbed by and reflected from the bloodless tissue compartments . as such , the changes in dc signals depend not only on wavelength but also sensor geometry and thus cannot be eliminated completely by computing the normalized r / ir ratio , as is typically the case in dual - wavelength pulse oximeters . the result is large variations in the slope of the calibration curves , as illustrated in fig6 . referring now to fig6 graphs c 1 , c 2 and c 3 show three calibration curves , presenting the variation of the slope for oxygen saturation values between 50 % and 100 %. referring to fig7 there is illustrated an optical sensor 10 designed according to the invention aimed at minimizing some of the measurement inaccuracies in a reflectance pulse oximeter . the sensor 10 comprises such main constructional parts as a light source 12 composed of three closely spaced light emitting elements ( e . g ., leds or laser sources ) 12 a , 12 b and 12 c generating light of three different wavelengths , respectively ; an array of discrete detectors ( e . g ., photodiodes ), a “ far ” detector 16 and a “ near ” detector 18 , arranged in two concentric ring - like arrangements ( constituting closed paths ) surrounding the light emitting elements ; and a light shield 14 . in the present example , six photodiodes form each ring . all these elements are accommodated in a sensor housing 17 . the light shield 14 is positioned between the photodiodes and the light emitting elements , and prevents direct optical coupling between them , thereby maximizing the fraction of backscattered light passing through the arterially perfused vascular tissue in the detected light . it should be noted that more than three wavelengths can be utilized in the sensor . the actual numbers of wavelengths used as a light source and the number of photodetectors in each ring are not limited and depend only on the electronic circuitry inside the oximeter . the array of discrete photodiodes can be replaced by one or more continuous photodetector rings . in addition to the r and ir light emitting elements 12 a and 12 b as used in the conventional pulse oximeter sensors , the sensor 10 incorporates the third , reference , light emitting element 12 c , which emits light in the nir - ir spectrum . wavelength λ1 and λ2 of the r and ir light emitting elements 12 a and 12 b are centered , respectively , around the peak emission values of 660 nm and 940 nm , and wavelength λ3 of the third light emitting element 12 c has the peak emission value above 700 nm ( typically ranging between 800 nm and 900 nm ). in the description below , the light emitting elements 12 b and 12 c are referred to as two ir light emitting elements , and wavelengths λ2 and λ3 are referred to as two ir wavelengths . during the operation of the sensor 10 , different light emitting elements are selectively operated for illuminating a measurement location ( not shown ) with different wavelengths . each of the photodetectors detects reflected light of different wavelengths and generates data indicative of the intensity i of the detected light of different wavelengths . it should be noted that the sensor can be of a compact design utilizing an integrated circuit manufactured by cmos technology . this technique is disclosed in a co - pending application assigned to the assignee of the present application . according to this technique , the sensor comprises a package including the light source , a block of two tubular optical waveguides of different diameters concentrically dislocated one inside the other and surrounding the light source , and an integrated circuit plate comprising two ring - like areas of photodiodes positioned concentrically one inside the other . the integrated circuit is also provided with a plurality of printed contact areas and electric conductors intended for mounting the light source thereon , controlling the light source , and transmitting electric signals produced by the photodiodes areas for further processing . [ 0080 ] fig8 illustrates a block diagram of a pulse oximeter 20 utilizing the above - described sensor 10 . the pulse oximeter typically includes a control unit 21 , which is composed of an electronic block 22 including a / d and d / a converters connectable to the sensor 10 , a microprocessor 24 for analyzing measured data , and a display 26 for presenting measurement results . the measured data ( i . e ., electrical output of the sensor 10 indicative of the detected light ) is directly processed in the block 22 , and the converted signal is further processed by the microprocessor 24 . the microprocessor 24 is operated by a suitable software model for analyzing the measured data and utilizing reference data ( i . e ., calibration curve stored in a memory ) to compute the oxygen saturation value , which is then presented on the display 26 . the analysis of the measured data utilizes the determination of ac - and dc - components in the detected light for each wavelength , λ1 , λ2 , and λ3 , respectively , i . e ., i 1 ( ac ) , i 1 ( dc ) , i 2 ( ac ) , i 2 ( dc ) , i 3 ( ac ) , and i 3 ( dc ) , and the calculation of ac / dc ratio for each wavelength , namely , w 1 = i 1 ( ac )/ i 1 ( dc ) , w 2 = i 2 ( ac ) / i 2 ( dc ) , and w 3 = i 3 ( ac ) / i 3 ( dc ) , as will be described more specifically further below with reference to fig9 and 10a - 10 c . the pulse oximeter 20 with the sensor arrangement shown in fig7 provides the following three possible ratio values : w 1 / w 2 , w 1 / w 3 and w 2 / w 3 . it should be noted that w 1 / w 2 and w 1 / w 3 are the ratios that typically have the highest sensitivity to oxygen saturation . this is due to the fact that λ1 is chosen in the red region of the electromagnetic spectrum , where the changes in the absorption between hb and hbo 2 are the largest , as described above with reference to fig1 . therefore , in principle , the absorption ratios formed by either wavelength pair λ1 and λ2 or wavelength pair λ1 and λ3 can be used to compute the value of sao 2 . the inventor conducted extensive human and animal studies , and confirmed that either of the two ratios w 1 / w 2 and w 1 / w 3 can be affected not only by changes in arterial oxygen saturation , but also by sensor placement and by the amount of pressure applied by the sensor on the skin . any calculation of sao 2 based on either of the two ratios w 1 / w 2 and w 1 / w 3 alone ( as normally done in commercially available dual - wavelength pulse oximeters ) could result in significant errors . furthermore , since at least two wavelengths are necessary for the calculation of arterial oxygen saturation , it is not feasible to self - correct the calibration curve for variations due to contact pressure or site - to - site variations utilizing the same two wavelengths used already to compute sao 2 . the inventor has found that the third ratio w 2 / w 3 formed by the combination of the two ir wavelengths is mostly dependent on changes in contact pressure or site - to - site variations . furthermore , this ratio can depend , but to a much lesser degree , on variations in arterial oxygen saturation . the dependency on arterial oxygen saturation , however , is easily minimized or eliminated completely , for example by selection and matching of the peak emission wavelengths and spectral characteristics of the two ir light emitting elements 12 b and 12 c . generally , the two ir wavelengths λ2 and λ3 are selected to coincide with the region of the optical absorption curve where hbo 2 absorbs slightly more light than hb , but in the spectral region , respectively , where the extinction coefficients of both hb and hbo 2 are nearly equal and remain relatively constant as a function of wavelength . for example , at 940 nm and 880 nm , the optical extinction coefficients of hb and hbo 2 are approximately equal to 0 . 29 and 0 . 21 , respectively . therefore , ideally , the ratio of w2 / w3 should be close to 1 , except for situations when the ac / dc signals measured from λ2 and λ3 are affected unequally causing the ratio w2 / w3 to deviate from 1 . fortunately , variations in the ratio w2 / w3 mimic changes in the ratios w 1 / w 2 and w 1 / w 3 since these ratios are all affected by similar variations in sensor positioning or other uncontrollable factors that normally can cause large errors in the calibration curve from which oxygen saturation is typically derived . thus , by tracking in real - time changes in the ratio formed by wavelengths λ2 and λ3 , it is possible to automatically correct for errors in the normalized ratios obtained from wavelengths λ1 and λ2 , or from λ1 and λ3 . the use of an additional third wavelength in the sensor serves another important function ( not available in conventional dual - wavelength pulse oximeters ), which is associated with the following . reflectance pulse oximeters have to be capable of detecting and relying on the processing of relatively low quality photoplethysmographic signals . accordingly , electronic or optical noise can cause large inaccuracies in the final computation of sao 2 . although the amount of electronic or optical noise pickup from the sensor can be minimized to some extent , it is impossible to render the signals measured by the pulse oximeter completely noise free . therefore , pulse oximeters rely on the assumption that any noise picked up during the measurement would be cancelled by calculating the ratio between the r - and ir - light intensities measured by the photodetector . practically , however , the amount of noise that is superimposed on the r - and ir - photoplethysmograms cannot be cancelled completely and , thus , can lead to significant errors in the final computation of sao 2 which , in dual - wavelength pulse oximeters , is based only on the ratio between two wavelengths . by utilizing a third wavelength , the invention has the added ability to compute sao 2 based on the ratio formed from either w 1 / w 2 or w 1 / w 3 . an algorithm utilized in the pulse oximeter according to the invention has the ability to track and compare in real - time changes between w 1 / w 2 and w 1 / w 3 to determine which ratio produces a more stable or less noisy signal and selectively choose the best ratio for calculating sao 2 . the method according to the invention utilizes the so - called “ selection process ” as part of the signal processing technique based on the measured data obtained with the multiple photodetectors . the main steps of the selection process are shown in fig9 in a self - explanatory manner . here , the symbol i corresponds to a single photodetector element in the array of multiple discrete photodetector elements , the term “ 1st ” signifies the last photodetector element in the array , and the term “ data ” signify three ratios ( ac / dc ) computed separately for each of the three wavelengths , namely , w 1 , w 2 and w 3 . the selection process is associated with the following : practically , each time one of the light emitting elements is in its operative position ( i . e ., switched on ), all of the photodetectors in the sensor receiving backscattered light from the skin . however , the intensity of the backscattered light measured by each photodetector may be different from that measured by the other photodetectors , depending on the anatomical structures underneath the sensor and its orientation relative to these structures . thus , the selection process is used to discriminate between photodetectors receiving “ good ” signals ( i . e ., “ good ” signal meaning that the calculation of spo 2 from the pulsating portion of the electro - optic signal ( ac ) and the constant portion ( dc ) would result in accurate value ) and “ bad ” signals ( i . e ., having ac and dc values which would result in inaccurate calculations of spo 2 ). accordingly , each data point ( i . e ., ratio w 1i , w 2i or w 3i detected at the corresponding if detector ) is either accepted , if it meets a certain criteria based for example on a certain ratio of ac to dc values ( e . g ., such that the intensity of ac signal is about 0 . 05 - 2 . 0 % of the intensity of dc signal ), or rejected . all of the accepted data points ( data from accepted detection locations ) are then used to calculate the ratios w 1 / w 2 , w 1 / w 3 and w 2 / w 3 , and to calculate the spo 2 value , in conjunction with the signal processing technique , as will be described further below with reference to fig1 a - 10 c . besides the use of the third ir - wavelength to compensate for changes in the internal calibration curve of the pulse oximeter , the pulse oximeter utilizing the sensor according to the invention provides a unique new method to compensate for errors due to sensor positioning and pressure variability . this method is based on multiple photodetector elements , instead of the conventional approach that relies on a single photodetector . while optical sensors with multiple photodetectors for application in reflectance pulse oximetry have been described before , their main limitation relates to the way the information derived from these photodetectors is processed . although the primary purpose of utilizing multiple photodetectors is to collect a larger portion of the backscattered light from the skin , practically , summing the individual intensities of each photodetector and using the resulting value to compute sao 2 can introduce large errors into the calculations . these errors can be caused , for example , by situations where the sensor is placed over inhomogeneous tissue structures such as when the sensor is mounted on the chest . the case may be such that , when using a continuous photodetector ring to collect the backscattered light , a portion of the photodetector ring lies over a rib , which acts as a strongly reflecting structure that contributes to a strong dc component , and the remaining part of the photodetector is positioned over the intercostals space , where the dc signal is much smaller . in this case , the final calculation of sao 2 would be inaccurate , if the current produced by this photodetector is used indiscriminately to compute the dc value before the final computation of sao 2 is performed . therefore , in addition to automatically correcting errors in the calibration curve as outlined above using three different leds ( one r and two different ir wavelengths ), the sensor 10 has the optional ability to track automatically and compare changes in the r / ir ratios obtained from each of the discrete photodiodes individually . for example , if some of either the near or the far photodetectors in the two concentrically arranged arrays detect larger than normal dc signals during the operation of one of the photodiodes compared to the other photodiodes in the sensor , it could be indicative of one of the following situations : the sensor is positioned unevenly , the sensor is partially covering a bony structure , or uneven pressure is exerted by the sensor on the skin causing partial skin “ blanching ” and therefore the blood - to - bloodless tissue ratio might be too high to allow accurate determination of sao 2 . if such a situation is detected , the oximeter has the ability to selectively disregard the readings obtained from the corresponding photodetectors . otherwise , if the dc and ac signals measured from each photodetector in the array are similar in magnitude , which is an indication that the sensor is positioned over a homogeneous area on the skin , the final computation of sao 2 can be based on equal contributions from every photodetector in the array . turning now to fig1 a , 10b and 10 c , there are illustrated three main steps of the signal processing technique utilized in the present invention here , th 1 and th 2 are two different threshold values ( determined experimentally ) related respectively to w 2 / w 3 and ( w 1 / w 2 − w 1 / w 3 ). during step 1 ( fig1 a ), measured data generated by the “ near ” and “ far ” photodetectors indicative of the detected ( backscattered ) light of wavelength λ2 and λ3 is analyzed to calculate the two ratios w 2 / w 3 ( far and near ). if one of the calculated ratios ( far or near ) is not in the range of 1 ± th 1 ( th 1 is for example 0 . 1 ), then this data point is rejected from the spo 2 calculation , but if both of them are not in the mentioned range , a corresponding alarm is generated indicative of that the sensor position should be adjusted . only if there are calculated ratios which are in the range of 1 ± th1 , they are accepted and the process ( data analysis ) proceeds by performing step 2 . step 2 ( fig1 b ) consists of determining whether the quality of each photoplethysmogram is acceptable or not . the quality determination is based on the relative magnitude of each ac component compared to its corresponding dc component . if the quality is not acceptable ( e . g ., the signal shape detected by any detector varies within a time frame of the measurement session , which may for example be 3 . 5 sec ), the data point is rejected and a corresponding alarm signal is generated . if the ac / dc ratio of w 1 , w 2 and w 3 are within an acceptable range , the respective data point is accepted , and the process proceeds through performing step 3 . in step 3 ( fig1 c ), the measured data is analyzed to calculate ratios w 1 / w 2 and w 1 / w 3 from data generated by far and near photodetectors , and to calculate the differences ( w 1 / w 2 − w 1 / w 3 ). in a perfect situation , w 1 / w 2 ( far ) is very close to w 1 / w 3 ( far ), and w 1 / w 2 ( near ) is very close to w 1 / w 3 ( near ). in a practical situation , this condition is not precisely satisfied , but all the ratios are close to each other if the measurement situation is “ good ”. then , the calculated differences are analyzed to determine the values ( corresponding to far and near photodetectors ) that are accepted and to use them in the spo 2 calculation . for each detector that satisfied the condition abs ( w 1 / w 2 − w 1 / w 3 )& lt ; th 2 ), where abs signifies the absolute value , its respective data point is accepted and used to calculate the oxygen saturation value that will be displayed . if the condition is not satisfied , the data point is rejected . if all data points are rejected , another measurement session is carried out . it should be noted that , although the steps 1 - 3 above are exemplified with respect to signal detection by both near and far photodetectors , each of these steps can be implemented by utilizing only one array of detection locations along the closed path . the provision of two such arrays , however , provides higher accuracy of measurements .
a sensor for use in an optical measurement device and a method for non - invasive measurement of a blood parameter . the sensor includes sensor housing , a source of radiation coupled to the housing , and a detector assembly coupled to the housing . the source of radiation is adapted to emit radiation at predetermined frequencies . the detector assembly is adapted to detect reflected radiation at least one predetermined frequency and to generate respective signals . the signals are used to determine the parameter of the blood .