I'llllllllllllllllllllllllllIllll||l|lllllllllllllllllllllllllllllllllllllll
`U5005297548A
`
`United States Patent
`Pologe
`
`[19]
`
`[11]
`
`Patent Number:
`
`5,297,548
`
`[45} Date of Patent: Mar. 29, 1994
`
`[54] ARTERIAL BLOOD MONITORING PROBE
`
`5,131.023
`
`8/1992 Mendelson er al‘ ................ 128/633
`
`[75]
`
`Inventor:
`
`Jones A. Foliage, Boulder. Colo.
`
`FOREIGN PATENT DOCUMENTS
`
`0303502 1/1939 Europa“ Fat or: -
`9004353 5/1990 PCT Int'l App].
`.
`9l1136
`31'1991 PCT lnt‘l App].
`.
`Primary Examiner—Angela D. Sykes
`Anal-My, Agent, or firm—Roger M. Rathbun; Larry R.
`C1952“; Jamal M- Graziano
`[57]
`ABSTRACT
`
`This arterial blood monitoring system takes advantage
`of the basic statistical property that arterial blood con-
`tains a plurality of dominant absorbers. whose measured
`fish! absorption spectra appear as a constant over a
`short interval of time. By measuring the transmitted
`light as it varies with arterial pulsation at selected wave-
`lengths of light, over a common light path, the relative
`amount of these dominant absorbers in the arterial
`blood can noninvasiver be determined. To ensure the
`common light path, a sandwich construction light de-
`"mm ‘5 “5°9-
`.
`IO Claims, 4 aning Sheets
`
`[73] Assignee: Ohmeda Inc, Murrayr Hill, NJ.
`
`[2]] Appl. No: 45,962
`[22] Filed:
`Apr. 12, 1993
`
`Related US. Application Data
`Continuation-impart ofSer. No. 832,551. Feb. 7, 1992.
`
`[63]
`
`A613 5/00'
`Int. Cl.5
`[51]
`[52] us. Cl. .................................... 128/633; 128/665;
`_
`355/41
`[53] Field Drawn:
`128/633—634.
`28/664467; 356/3941
`.
`Refemc's Clad
`U.S. PATENT DOCUMENTS
`.
`
`[59l
`
`123/666 x
`128/666 X
`125/633
`
`
`
`i'gg’gig ifiiggg
`4,367,165 9x19s9 Noller et al.
`5;0?8.136
`1/1992 Stone et a].
`5,!21.406
`7/1992 Yamaguehi
`
`.
`
`IOI
`
`“400
`
`
`s—e—l
`
`CURRENT
`
`T0
`
`
`v01.1mm:
`
`
`
`Iconven'ren
`
`
`SYNCHRONDUS
`MPLI' DEMODULA—
`F'ER
`TIDN
`'
`
`
`
`ETECTOR "3‘
`
`PROCESSOR
`
`
`
`
`DISPLAY
`DRIVER
`5to
`
`
`
`
`DIS PLAY MODULE
`
`PLETHYSMOGRAPH IC
`WAVEFOFIM
`
` HEART RATE
`'1 BPMl
`
`
`
`II5c
`
`0001
`
`Apple Inc.
`APLl 01 9
`
`US. Patent No. 8,929,965
`
`FITBIT, Ex. 1019
`
`Apple Inc.
`APL1019
`U.S. Patent No. 8,929,965
`
`0001
`
`FITBIT, Ex. 1019
`
`

`

`US. Patent
`
`Mar. 29, 1994
`
`Sheet 1 of 4
`
`5,297,548
`
`IOI
`
`“400
`
`SYNCHRONOUS
`DEMODULA—
`TION
`
`CUR RENT
`TO
`VOLTAGE
`CONVERTER
`
`
`
`
`
`105
`
`
`
`
`
`I360
`
`SCALING
`AMPLIFIERS
`
`.
`
`I35
`
`|36bfl
`
`fl:
`
`q 8
`53$53L)
`C3
`D:
`n.
`
`DISPLAY MODULE
`
` PLETHYSMOGRAPH I C
`WAVEFORM 7.8002
`
`HEART RATE
`(BPMI
`
`H4
`
`“Sb
`
`H5O
`
`0002
`
`HTBfiflEx1019
`
`0002
`
`FITBIT, Ex. 1019
`
`

`

`US. Patent
`
`'
`
`Mar. 29, 1994
`
`Sheet 2 of 4
`
`5,297,548
`
`ARTERIAL PULSAT ION
`ABSORPTION
`
`
`ABSORPTION
`
`
`ARTERIAL BLOOD
`ABSORPTION
`
`VENOUS BLOOD
`ABSORPTION
`
`OTHER TISSUE
`ABSORPTION
`
`TIME
`
`
`
`0003
`
`FITBIT, EX. 1019
`
`0003
`
`FITBIT, Ex. 1019
`
`

`

`US. Patent
`
`Mar. 29, 1994
`
`Sheet 3 of 4
`
`5,297,548
`
`RECEIVE DATA FROM PROBE AND
`
`410
`
`TRANSMIT DATA TO DISPLAY
`DRIVER TO PRODUCE
`
`STORE IN MEMORY
`WAVEFORM DISPLAY
`
`
`402x COMPUTE DIFFEREN‘RAL CHANGE IN
`
`ABSORPTION AT FIRST WAVELENGTH
`
`403
`
`COMPUTE DIFFERENTIAL CHANGE IN
`
`ABSORPTION AT SECOND WAVELENGTH
`
`
`
`
`
`COMPUTE RATIO OF COMPUTED DIFFERENTIAL
`CHANGE IN ABSORPION FOR FIRST WAVELENGTH
`TO COMPUTED DIFFERENTIAL CHANGE IN
`ABSORPTION FOR SECOND WAVELENGTH
`
`
`
`
`
`DRIVER TO PRODUCE NUMERIC OUTPUT
`
`TRANSMIT COMPUTED tHb TO DISPLAY
`
`‘FIGURE
`
`4
`
`0004
`
`FITBIT, EX. 1019
`
`0004
`
`FITBIT, Ex. 1019
`
`

`

`US. Patent
`
`Mar.29,1994
`
`Sheet 4 of 4
`
`5,297,548
`
`
` asczxv: DATA rnou p203: gun
`
`smog: In xenon!
`
`510
` mnaxsuxr DATA To nIsfiLn!
`
`
`DRIVER TO pnonuau
`
`unrnronx nxsrnnx
`
`
`
`
`COKPUTB BIrleENTIhL annual I“
`ABSORPTIOH IT BBCOHD IIVILBHG H
`
`
`
`
`COHPUTB DIPIERIHTILL CHINE! IN
`ABSORPTION 1T THIRD IIVILEHGTB
`
`
`
`
`
`
`
`COHPUT! tub IRDH TEE
`COMPUTB 8802 ’30! TH!
`IZLSURSD flIPPBRENTIlL
`MEASURED DIrPERBNTILL
`CHANGE IN IBBDRPTION
`CHANGE IH ABSORPTION
`IT ALL IIYELIHGTEB
`IT ALL IIVBLINGTHB
`
`OOKPUTB 02 at [30K tlb 33D 3102
`
`
`
`
`508 r-+—~““h‘
`
`mnanaurw tub, 3:02 nun 02 at Dawn
`TO DISPLAY DRIVER TO PRODUCE
`HUflIRIC OUTPUT
`
`FIGHJREZ
`
`S
`
`0005
`
`HTBfiflEx1019
`
`0005
`
`FITBIT, Ex. 1019
`
`

`

`1
`
`5,297,548
`
`ARTERIAL BLOOD MONITORING PROBE
`
`CROSS-REFERENCE TO RELATED
`APPLICATIONS
`
`_
`
`5
`
`This application is a continuation-in-part of us. pa— .
`tent application Ser. No. 07/832,551. titled “improved
`Arterial Blood Monitoring System", filed Feb. 7. 1992,
`pending
`
`10
`
`FIELD OF THE INVENTION
`
`This invention relates to non-invasive photoplethys-
`mographic measurement of blood analytes and, in par-
`ticular. to a probe for use in an arterial blood monitoring
`system to more accurately measure the change in inten-
`sity of the light transmitted through the arterial blood of
`a patient.
`
`15
`
`PROBLEM
`
`2
`tor. The greater the departure of the light beams from a
`common light path, the more significant the opportu-
`nity for the introduction of errors into the resultant
`measurements. This is especially true if multiple inde-
`pendent discrete light sources and multiple discrete
`light detectors are used in the probe, resulting in sepa-
`rate light
`transmission paths through the intervening
`appendage. The use of multiple light detectors, each
`Sensitive to different wavelength regions, becomes a
`necessity if the wavelengths of light selected are far
`apart in wavelength, since there does not exist a single
`light detector device that can detect a wide bandwidth
`of light with significant speed. sensitivity and an accept-
`ably flat response. Therefore, existing probe designs can
`introduce errors into the measurements by their inabil-
`ity to transmit a plurality of light beams substantially
`along a common light path through the arteriolar bed of
`the appendage being monitored.
`SOLUTION
`
`25
`
`30
`
`35
`
`It is a problem in the field medical monitoring equip- 2°
`ment to accurately measure various parameters of arte-
`rial blood in a noninvasive manner. For example. the
`oxygen saturation (Sn 0;} of the hoglobin in arterial
`blood is determined by the relative proportions of city-
`genated hemoglobin and reduced hemoglobin in the
`arterial blood. A. pulse oximeter system noninvasively
`determines the oxygen saturation of the hemoglobin by
`measuring the dili‘erence in the light absorption of these
`two forms of hemoglobin. Reduced hemoglobin absorbs
`more light in the red band (600-800 am) than does oxy-
`hemoglobin while oxyhemoglohin absorbs more light in
`the near infrared band (800—1000 nm) than does reduced
`hemoglobin.
`The pulse oximeter includes a probe that is placed in
`contact with the skin, either on a flat surface in the case
`of reflectance probes or across some appendage in the
`came of a transmission probe. The probe contains two
`light emitting diodes. each of which emits a beam of
`light at a specific wavelength, one in the red band and
`one in the infrared band. The magnitude of red and 4-0
`infrared light transmitted through the intervening ap-
`pendage contains a non-pulsatile component which is
`influenced by the absorbency of tissue, venous blood,
`capillary blood, non—pulsatile arterial blood, and the
`intensity of the light source. The pulsatile component of 4-5
`the received signals is an indication of the expansion of
`the arteriolar bed in the appendage with arterial blood.
`The effects of diil'erent tissue thicknesses and skin pig-
`mentation in the appendage can be removed from the
`received signals by normalizing the change in intensity
`of the received signal by the absolute intensity of the
`received signal. Taking the ratio of the mathematically
`processed and normalized red and infrared signals re-
`sults in a number which is theoretically a functiOn of
`only the concentration of oxyhemoglobin and reduced
`hemoglobin in the arterial blood. This assumes that
`oxyhemcglobin and reduced hemoglobin are the only
`substantial absorbers in the arterial blood.
`The amplitude of the pulsatile component is a very
`small percentage of the total signal amplitude and de-
`pends on the blood volume change per pulse and the
`oxygen saturation (8,02) of the arterial blood. The
`received red and infrared signals have an exponential
`relationship to the path length of the arterial blood. The
`photoplethysmographic measurement of these analytes
`is predicated on the assumption that the light beams
`from the two light sources follow identical paths
`through the intervening appendage to the light detec—
`
`55
`
`65
`
`The above described problems are solved and a tech-
`nical advance achieved in the field by the probe for an
`arterial blood monitoring system that creates a single
`light path through an appendage to noninvasiver mea-
`sure and calculate characteristics of arterial blood. This
`arterial blood monitoring system probe takes advantage
`of the basic statistical property that arterial blood con-
`tains a plurality of dominant absorbers, whose measured
`light absorption spectra appear as a constant over a
`short interval of time. The arterial blood characteristics
`to be measured are empirically related to the changes in
`the measured light transmission through the plurality of
`dominant absorbers as a function of the changes in arte-
`rial blood volume at the probe site. By measuring the
`transmitted light as it varies with arterial pulsation at a
`plurality of selected wavelengths of light, over a single
`common light path, the relative amount of these domi-
`nant absorbers in the arterial blood can noninvasiver
`be determined.
`By selecting one wavelength of light around 1270 nm,
`where water has a measurable extinction and second
`and third wavelengths at about 660 nm and 940 nm, a
`direct relationship between the transmitted intensities at
`these three wavelengths and the arterial hemoglobin
`concentration exists and can be calculated. The accu-
`rate detection of these three wavelengths of light is
`accomplished by the use of two different light detec-
`tors. To avoid the problem of different light paths
`through the intervening appendage. a sandwich or lay-
`ered detector design is used in the probe. The light
`detector consists of a multiple layer element that con-
`tains a germanium photodiode placed under, and coinci-
`dent with, a silicon phatodiode. For the wavelengths of
`light shorter than approximately 1000 nm. the silicon
`photodiode receives the incident light and produces a
`signal indicative of the intensity of the received light.
`Above this wavelength. the silicon photodiode becomes
`transparent and the germanium photodiode picks up the
`incident
`light. Thus, the light from the three light
`sources is transmitted through the tissue along substan-
`tially identical light paths to be detected by the coinci-
`dent light detectors at exactly the same "exit area",
`regardless of wavelength. By constraining the detected
`light to traverse one path through the tissue, regardless
`of wavelength, this apparatus avoids the inaccuracies
`caused by sampling different cross-sections of tissue, as
`
`0006
`
`FITBIT, EX. 1019
`
`0006
`
`FITBIT, Ex. 1019
`
`

`

`3
`with two or three discrete light detectors mounted side
`by side.
`-
`BRIEF DESCRIPTION OF THE DRAWING
`
`FIG. 1 illustrates in block diagram form the overall
`architecture of the arterial blood monitoring system and
`the probe of the present invention;
`FIG. 2 illustrates in graphical form the various com-
`ponents of' the input signal from the probe;
`FIG. 3 illustrates a cross-section view of the light
`detector used in the probe of the present invention;
`FIG. 4 illustrates in flow diagram form the opera-
`tional steps taken by a two wavelength arterial blood
`monitoring system to measure selected components in
`arterial blood; and
`FIG. 5 illustrates in flow diagram form the opera-.
`tional steps taken by a three wavelength arterial blood
`monitoring system to measure selected components in
`arterial blood.
`
`DETAILED DESCRIPTION
`
`An arterial blood monitoring system takes advantage
`of the basic statistical property that arterial blood con-
`tains a plurality of dominant absorbers. whose measured
`light absorption spectra appear as a constant over a
`short interval of time. The arterial blood characteristics
`to be measured are empirically related to the changes in
`the measured light transmission through the plurality of
`dominant absorbers as a function of the changes in the
`arterial blood volume at the probe site. Therefore, by
`measuring the transmitted light as it varies with arterial
`pulsation, at selected wavelengths, the relative amount
`of these dominant absorbers in the arterial blood can
`noninvasiver be determined. A single probe can be
`used to generate the plurality of wavelengths of light.
`therefore simplifying the arterial blood monitoring sys-
`tern.
`
`Definition of Terms
`
`10 =The intensity of the beam of light at a given wave-
`length incident on the tissue-under—test, where the
`wavelength is denoted by the subscript.
`l=The instantaneous value of the intensity of the light
`received by the detector. The light is at a given wave-
`length, which wavelength is indicated by a subscript.
`e=The extinction coefficient of light by a given sub-
`stance (indicated by a superscript) at a given wave-
`length (indicated by a subscript).
`C=The concentration of a given substance (indicated
`by a superscript).
`L=Thc pathlength of a given substance (indicated by a
`superscript).
`tHb=Total hemoglobin measured in arterial blood.
`Usually expressed in terms of grams per deciliter.
`0:Used as a superscript to represent oxyhemoglobin.
`R=Used as a superscript to represent reduced hemo-
`globin.
`W=Used as a superscript to represent water.
`t=Used as a superscript to represent the combination of
`oxyhemoglobin and reduced hemoglobin.
`
`System Architecture
`FIG. 1 illustrates in block diagram form the overall
`architecture of the arterial blood monitoring system 100
`and the probe 101 of the present invention. The arterial
`blood monitoring system 100 consists of a probe 101
`connected to probe interface circuit 102 by means of a
`set of electrical conductors 103 and connector 1030.
`
`10'
`
`IS
`
`20
`
`25
`
`_
`30
`
`.35
`
`45
`
`50
`
`55
`
`60
`
`65
`
`5,297,548
`
`4
`The probe 101 consists of an exterior housing 104 that
`applies the active elements of the probe 101 to the tissue
`under test, such as a finger 105, containing an arterial
`blood flow that is to be monitored. Included within
`housing 104 is a plurality (at least two) of light emitting
`devices 111, 112 and at least one corresponding light
`detector 113.
`Emitter driver circuit 131 produces the analog drive
`signals to activate light emitting devices 111, 112 in
`probe 101. These analog drive signals are carried over
`cable 103 to probe 101. To measure the concentration of
`total hemoglobin (ti-lb), oxygen saturation ($402), or
`other blood analytes, in arterial blood, the concentra-
`tion of several dominant absorbers contained in the
`arterial blood must be measured. In particular. for the
`measurement of total hemoglobin (ll-1b), concentration
`of the water and hemoglobin components of the arterial
`blood must be measured. The light emitting devices 111.
`112 each produce an output light beam of predeter-
`mined wavelength which is directed at the finger 105
`enclosed by housing 104.
`In this embodiment,
`light
`emitting device 111 is selected to produce a beam of
`light at approximately 310 nm. which wavelength is
`substantially isobestic to the oxygenated and deoxygen-
`ated components of the hemoglobin in the arterial blood
`(that is, the extinction coefficients of the oxygenated
`and deoxygenated hemoglobin are substantially identi-
`cal). Light emitting device 112 is selected to produce a
`beam of light at approximately [270 nm. The selection
`of these two wavelengths is such that water is transpar-
`ent at the first wavelength oflight (810 run) but detected
`at the second (longer) wavelength of light (1270 run). In
`addition, these wavelengths are such that the extinction
`coefficients of the two components (water and hemo-
`globin) differ at the first wavelength oflight. Further, at
`both wavelengths the two species of hemoglobin are
`substantially isobestic in extinction but not transparent.
`The light detector 113 monitors the level of light that
`is transmitted through or reflected from finger 105. The
`analog data signals produced by light detector 113 in
`response to the received beams of light are received
`from probe 101 over conductors 103 and filtered by
`, analog hardware 132—134 in probe interface circuit 102.
`The input analog data from probe 101 may be decom-
`posed into its non-pulsatile and pulsatile sub-elements in
`probe interface circuit 102 in order to provide accurate,
`high resolution, measurements of these components.
`The pulsatilc component typically represents anywhere
`from 0.05% to 20% of the total input signal and the
`decomposition of the input signal
`into pulsatile and
`non-pulsatile components permits accurate analog to
`digital c0nversion of even the smallest of these pulsatile
`components.
`In order to distinguish between the light beams pro-
`duced by first 111 and second 112 light emitting de-
`vices. these light emitting devices 111, 112 are modu-
`lated in a manner to allow the output of the light detec-
`tor 113 to be synchronously demodulated. Ambient
`light, being unmodulated, is easily eliminated by the
`demodulator process.
`
`Signal Components
`FIG. 2 illustrates in graphical form (not to scale) the
`various components of the total absorption produced by
`finger 105. The light detector output signal, high where
`absorption is low and visa versa, consists of a large
`magnitude non-pulsatiie component and a small magni—
`tude pulsatile component. The non-pulsatile component
`
`0007
`
`FITBIT, EX. 1019
`
`0007
`
`FITBIT, Ex. 1019
`
`

`

`5,297,548
`
`5
`represents light remaining after absorption due to a
`combination of venous blood. cutaneous tissue. bone.
`and constant arterial blood while the small pulsatile
`component is caused by the light absorption due to
`pulsatile arterial blood flow that is to be measured.
`Following synchronous demodulation. the data signals
`produced by light detector 113 and transmitted to probe
`interface circuit 102 consist of a series of data points that
`are digitized and stored in memory 106. Since the first
`111 and second 112 light emitting devices are sampled
`simultaneously and in rapid succession, these digitized
`data points consist of a plurality of sets of measure—
`ments, with one set corresponding to samples of the
`light beam intensity at a first wavelength, the other set
`15
`corresponding to samples of the light beam intensity at
`a second wavelength, and, in some schemes, a third set -
`corresponding to the intensity of the ambient light.
`Ideally, in pulse oximeter systems red and infrared
`wavelengths of light are used and the ratio of the nor-
`malized derivative (01' logarithm) of the red intensity to
`the normalized derivative (or logarithm) of the infrared
`intensity is a constant. This constant is indicative of the
`partial oxygenation (5402) of the hemoglobin in the
`arterial blood flow. It is obvious that this ratio changes
`as 5.10; changes but. for a short interval with rapid
`enough sampling rate, the ratio remains constant.
`Probe Interface Circuit
`
`10
`
`20
`
`25
`
`6
`scaling amplifiers 135 such that they can be converted,
`with optimal resolution,
`to a digital equivalent. All
`channels output by sealing amplifiers 135 are then si-
`multaneously sampled by the sample/hold circuitry
`136a. 136b, . . . 13611. The sampled data is passed a chan-
`nel at a time via multiplexer 137 to the analog to digital
`converter 138. From there the data, now in digital form,
`is sent on to data processing circuit 107 where it is
`stored in memory 106 for processing. The digital data
`represents the substantially simultaneously sampled
`amplitudes of the received light intensities from each of
`the wavelengths used at a sampling frequency of typi-
`cally 30 He or greater. These data values are referred to
`as ll, 1;, .
`.
`. IN, where the subscript indicates the given
`wavelength. 1,, then indicates the received light inten~
`sity at any given wavelength.
`
`Data Processing Circuit
`In a two wavelength system, data processing circuit
`10‘! computes a ratio from the digital amplitude data
`measured at each wavelength of light. In particular. this
`process used by data processing circuit 107 is illustrated
`in flow diagram form in FIG. 4. At step 401, data pro-
`cessing circuit 107 receives a set of digital input data
`indicative of the measured intensity of light at both
`wavelengths, as received by light detector 113. Data
`processing circuit 101 at step 410 transmits the received
`set of data to diaplay driver 109 for display in graphical
`form on display 114. The displayed waveform repre-
`sents the pulsatile component of the arterial blood. Data
`processing circuit 101r also stores the received set of
`data in memory 106 and uses this set of data and the last
`most recently received set of data to compute at steps
`402 and 403 the differential change in absorption of the
`arterial blood in finger 105 at the first and second se-
`lected wavelengths of light. re5pectively. The differen-
`tial change in absorption at wavelength n is computed
`by data processing circuit 107 as:
`
`Hf
`at“ _ __”
`
`(it
`
`‘Because din is a mathematical construct. it is approxi-
`mated in arterial blood monitoring system 100 by AI",
`where M... is the difference between two consecutively
`received 1,. values. Only AI values that are caused by a
`small but non zero change in path length through finger
`105 are used and therefore AIR can also be a longer
`interval of time if necessary to obtain a sufficient change
`in rccaived intensity of the beam of light. The 1,. value
`used in equation 1 is the average of the two successively
`received I" values used to compute A15.
`In a two wavelength system, a final ratio is then cal-
`culated by data processing circuit 107 at step ‘04 as:
`
`M;
`R = TA:
`
`(2)
`
`where the data values used to compute dA1are from the
`same points in time as the data values used to compute
`dAz.
`This ratio is then used in a calibration equation by
`data processing circuit 107 at step ms to relate the R
`value to a specific blood analyte value. For example,
`when measuring total hemoglobin, the calibration equa-
`tion is approximated by a second order polynomial of
`the form:
`
`The actual analog data received by the probe inter-
`face circuit 102 can include a fairly significant noise
`component which is caused by a number of sources
`including motion of finger 105, the introduction of am-
`bient light into housing 104. and various sources of
`electrical noise. These noise components skew the val-
`ues of either or both of the magnitudes measured in each
`set of data points destroying the correct relationship
`between the red and infrared signals. Existing pulse
`oximeter circuits make use of various filtering tech-
`niques to minimize the impact of noise on the Sat):
`value measured by the system. This filtering circuitry
`and software/algorithms are analogous to that used in
`the arterial blood monitoring system 100 and are there
`fore not described in detail herein.
`Probe interface circuit 102 includes emitter driver
`circuit 131 that is capable of driving light emitting de-
`vices 111, 112 such that the light beams produced tra-
`verse finger 105 and sufficient light intensity is incident
`on light detector 113 to produce data indicative of the
`light absorption of the dominant absorbers in arterial
`blood. The data produced by light detector 113 (voltage
`equivalent of the received light
`intensities) at each
`wavelength is kept distinct and can be processed inde-
`pendently. This can be done by any of the many
`schemes presently in use for pulse crimetry, such as
`time division multiplexing, or frequency division multi-
`plexing.
`The light received from finger 105 is converted to an
`equivalent current signal by the photodiodes of light
`detector 113, and then canverted to a voltage signal by
`the current to voltage corrverter 132. The data is then
`amplified by amplifier 13, and demultiplerted via syn-
`chronous demodulation circuit 134. The demultiplerted
`data comprises analog voltage signals applied to leads
`CHAN 1, CHAN 2 .
`.
`. CHAN 11 representative of the
`intensity of the received light at each of the wave-
`lengths of light produced by light emitting devices 111,
`112, respectively. The voltage signals on leads CHAN
`1, CHAN 2 are then sealed (further amplification) by
`
`30
`
`35
`
`45
`
`55
`
`65
`
`0008
`
`FITBIT, EX. 1019
`
`0008
`
`FITBIT, Ex. 1019
`
`

`

`5,297,548
`
`er= .431 +3)! + r:
`
`(3)
`
`Where A, B. and C are constants that depend on the
`specific wavelengths of light used.
`The tHb value is then output by data processing cir-
`cuit 107 at step two to display driver 109 (and/or hard-
`oopy) to display in human-readable form an diSplay 115
`a numeric value of the concentration of total hemoglo-
`bin in the arterial blood of finger 105. Processing then
`returns to step 401.
`
`blood monitoring system 100 to read the differential
`change in absorption (dA values) as accurately as possi-
`ble. This leaves only the values of the differential path
`lengths dL as unknowns. With two equations. the two
`dL values can be uniquely and simply solved for.
`Writing the proportion of hemoglobin in the arterial
`blood as:
`
`ID
`
`d'L‘
`mm!“ m z w + at."
`
`(Bl
`
`Theory
`This device is based on the theory that:
`
`Mnf- en’C‘d "
`
`15
`
`(‘1
`
`A differential change in absorption at a given wave-
`length n to a given substance (dAnS), is equal to the
`extinction ofthat substance (ens) times the concentratiOn
`(C1) of that substance times the differential change in
`pathlength of that substance (dL’).
`Further the differential change in absorption can be
`defined as:
`
`20
`
`
`d1...
`L:
`
`“a:
`
`15
`
`s
`(l
`
`Note that no measurement of the incident light inten-
`sity, Io. is required to measure the differential change in
`absorption dA. However. samples of 1.. must be taken
`sufficiently close in time so that A1,. represents a good
`mathematical approximation of dln.
`To determine the relative proportions of two domi-
`nant absorbers, in this case water and hemoglobin, one
`chooses two wavelengths of light at which the two
`absorbers have extinctions. such that the following set
`of simultaneous equations has a unique solution for all
`possible concentrations and pathlengths of the two ab-
`sorbers.
`
`dA1=el‘dL‘+ (1'61"
`
`dA::ez'n‘L'+Ez”dL”
`
`(6)
`
`(7}
`
`In this system of equations it is assumed that the only
`components which change in pathlength are those of
`the arterial blood. Further it is assumed that the primary
`absorbers are those of water and hemoglobin where the
`hemoglobin species in the blood are essentially only
`those of onyhemoglobin and reduced hemoglobin.
`Choosing a. wavelength of light
`that represents an
`isobestic point for the two species of hemoglobin. such
`as 804 manometers, minimizes the effects of changes in
`oxygen saturation on the total hemoglobin readings.
`Notice that in equations 6 and 7 above. the concentra-
`tion term expressed in equation 4 has been eliminated.
`By viewing the optical system as compartmentalized.
`that is. looking at the tissue under test as one in which
`the light first passes through 100% skin tissue, followed
`by l00% venous blood. followed by 100% arterial he-
`moglobin. followed by 100% water. and so on. the
`commutation terms expressed by equation 4 are actu-
`ally constants. Thus, beginning with equation 6, the
`extinction coefficients are meant to represent the combi-
`nation of the actual extinction coefficient and the actual
`concentration for 100% of any given absorber.
`In the system of equations (Equations 6. 7) the extinc-
`tion e values are constants and it is the job of the arterial
`
`30
`
`35
`
`40
`
`45
`
`55
`
`65
`
`While this proportion is not directly the tHb, it is di-
`rectly related to it. And while this relationship could be
`theoretically derived. an empirical relationship (as de-
`fined in equation 3) is measured instead. This is neces-
`sary due to several ways in which the true optical sys-
`tem of living tissues and realistic optical elements devi’
`ate from the exact theoretical model developed here.
`Equation 3 is therefore referred to as the calibration
`equation and its coefficients A. B. and C. are experimen-
`tally derived via clinical testing. The coefficients are
`then installed in the arterial blood monitoring system
`software. It should be noted that these coefficients dif-
`fer for different wavelength emitters.
`The wavelengths of light produced by light emitting
`devices 111, 112 are also selected so as to optimize the
`performance of the entire electro Optical system: low
`enough light absorption so that sufficient optical signal
`is received by light detector 113 and high enough light
`absorption so that there is an appreciable change in light
`absorption over the range of physiological changes in
`pathlength caused by the pulsation of arterial blood.
`Typical wavelengths of light selected for a realization
`of this system are 810 nm and 1270 not. however many
`wavelength combinations meeting the above criteria
`can be used.
`
`Probe
`
`Probe 101 contains a minimum of two lights emitting '
`devices 111, 112, each of which produces a beam of
`light centered about a selected wavelength (810 run and
`1270 nm. respectively). Probe 101 also contains light
`detector 113 capable of receiving the emitted wave-
`lengths of light. In the present implementation, the light
`detector 113 consists of a multiple layer clement. shown
`in additional detail in FIG. 3, that contains a germanium
`photodiodc 1131.1 placed under a silicon photodiode
`1130. For the Wavelengths of light shorter than approxi-
`mately 1000 nut, the silicon photodiode 113a receives
`the incident light. Above this wavelength. the silicon
`photodiode 113w:r becomes transparent and the germs-
`nium photodiocle 1135 picks up the incident light. Probe
`101 includes a cable 103 and connector 1113a for trans-
`mitting and receiving signals between probe 101 and
`probe interface circuit 102. Probe 101 is positioned on
`the tissue either in the transmission mode: light emitting
`devices 111, 112 on one side and light detector 113 on
`the other side of finger 105. earlobe. toe or other appro-
`priate site through which light can be received by the
`light detector 113 at acceptable signal levels; or in the
`reflectance mode: in which the light emitting device
`111, 112 and light detector 113 are placed on the same
`side of the tissue under test. such as the forehead or
`forearm.
`
`0009
`
`FITBIT, EX. 1019
`
`0009
`
`FITBIT, Ex. 1019
`
`

`

`9
`
`5,297,548
`
`10
`(021%). reduced hemoglobin (RHb), and water (H20).
`
`a!“
`also
`m " dL°+ ark + + at.”
`
`“03
`
`Note that this equation shows only that the total hemo-
`globin is preportional
`to this ratio of path length
`changes, not equal to it. This is due. at least in part, to
`the fact the tHb is measured in terms of grams/deciliter
`of whole blood and this equation is a ratio of path
`lengths. There is a one to one correspondence between
`this ratio of path lengths and the tHb which is deter-
`mined, and curve fit, experimentally. The empirical
`curve fit also compensates for the differences between
`the theoretical models and the actual optical systems.
`
`s
`‘02
`
`410
`=-—--—-—--
`dL°+dLR
`
`(l. l)
`
`For a three wavelength system, with the subscripts 1, 2,
`and 3 indicating the specific wavelengths used we can
`write following system of equations
`
`Combination tHb Monitor and Pulse Oximeter
`
`The methodologies for pulse oximetry are well
`known. The method of obtaining tHb noninvasively
`and in real time has been disclosed above. The arterial
`blood monitoring system of the present invention can
`combine the two technologies to create a device for
`measurement of both parameters. tI-lb is an interfering
`substance in the measurement of SRO: by the present
`technologies. By ‘interfering substance" it is meant that
`variations in tHb cause variations in the 5,0; as ready
`by a pulse oximeter. These variations in 59028.1: corre-
`lated to, but not corrected for, the till) level. A device
`capable of measuring tHb can therefor provide a means
`for eliminating the error it causes in determining 5,02.
`The same holds true in terms of 5502 being an interfer- ,
`ing substance in the measurement of tI-Ib. The solution
`to this problem lies in a combination device capable of
`reading both parameters. Such a device can be simply
`obtained by using two wavelengths to derive the 5,02.
`and two more as described above for obtaining tHb.
`The resulting values for $402 and Thb can then be used
`to correct the readings of the other. A more sophisti-
`cated system uses a three wavelength system, where the
`practical realization of this system utilizes the standard
`onimetry wavelengths of 660 nm and 940 mu produced
`by two light emitting devices Illa, 111b, along with a
`wavelength of 1270 um produced by a light emitting
`device 112. (Once again, any three wavelengths that
`meet the criteria stated above for a standalone tl-l'b
`system can be used.) In addition, the two Segment light
`detector 113 is activated in a manner to reflect the use of
`three wavelengths of light. Silicon photodiode 113::
`detects both of the light beam