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Method and apparatus for determining blood parameters and vital signs of a patient - LEIDEN JEFFREY M.
United States Patent Application 20030212316
Leiden, Jeffrey M. (Glencoe, IL, US)
Khalil, Omar S. (Libertyville, IL, US)
Shain, Eric Brian (Glencoe, IL, US)
Kantor, Stanislaw (Buffalo Grove, IL, US)
Yeh, Shu-jen (Grayslake, IL, US)
Koziarz, James J. (Highland Park, IL, US)
Hanna, Charles F. (Libertyville, IL, US)
Wu, Xiaomao (Gurnee, IL, US)
Hohs, Ronald R. (Kenosha, WI, US)
10/144224
LEIDEN JEFFREY M.
KHALIL OMAR S.
SHAIN ERIC BRIAN
KANTOR STANISLAW
YEH SHU-JEN
KOZIARZ JAMES J.
HANNA CHARLES F.
WU XIAOMAO
HOHS RONALD R.
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20070197857 Method for making a blood pump and pumping blood August, 2007 Palmer
1. An apparatus for monitoring changes in blood parameters and vital signs of a patient, said apparatus comprising: a) means for illuminating a body part of a patient; b) means for collecting optical signals over a period of time from said body part; c) means for effecting pressure changes or temperature changes or both of the foregoing types of changes in said body part; d) means for measuring pressure changes or temperature changes or both types of the foregoing changes experienced by said body part; e) means for calculating at least one value of at least one blood parameter of said patient from the collected optical signals; f) means for determining at least one value of at least one vital sign of the patient from said collected optical signals; g) means for reporting said at least one value of said at least one blood parameter and said at least one value of said at least one vital sign; and h) means for providing an alarm when (1) said at least one value of said at least one vital sign crosses a specified cut-off value or (2) the rate of change in said at least one value of said at least one vital sign crosses a specified cut-off value or (3) said at least one value of said at least one blood parameter crosses a specified cut-off value or (4) the rate of change in the at least one value of the at least one blood parameter crosses a specified cut-off value.
7. A method for monitoring at least one blood parameter and at least one vital sign of a patient, said method comprising the steps of: a) collecting a set of optical measurements on a body part of a patient over a period of time; b) determining at least one value of at least one blood parameter of said patient; c) determining at least one value of at least one vital sign of said patient from said set of optical measurements; d) reporting a combination of said at least one value of said at least one blood parameter and said at least one value of said at least one vital sign; e) repeating steps a), b), c), and d) a sufficient number of times until a trend can be observed; and f) activating an alarm when (1) said at least one value of said at least one vital sign crosses a specified cut-off value or (2) the rate of change in said at least one value of said at least one vital sign crosses a specified cut-off value or (3) said at least one value of said at least one blood parameter crosses a specified cut-off value or (4) the rate of change in said at least one value of said at least one blood parameter crosses a specified cut-off value.
[0009] The concentration of hemoglobin and the ratio of oxygenated hemoglobin to total hemoglobin in blood are important parameters for indicating the anemic state and wellness of a patient. Hemoglobin is the protein that transports oxygen. It has a molecular weight of 65,500 Daltons; thus, 1 gram of hemoglobin is equivalent to 1.55×10−5 mole. The concentration of hemoglobin is expressed in g/dL. The hematocrit value is the ratio of volume of red blood cells to total blood volume, which comprises the volume of red blood cells and the volume of plasma. The hematocrit value is expressed as a percentage (i.e., percentage by volume of red cells in whole blood). While the measurement of concentration of hemoglobin provides an indication of the oxygen transport status of the patient, the measurement of the hematocrit value provides an indication of both red blood cells for transport of oxygen and plasma for transport of nutrients. The measurement of the hematocrit value is particularly important when a change in body hemodynamics is expected, such as during operations of long duration, such as, for example, cardiac and orthopedic surgery. Other applications of the measurement of the hematocrit value include the treatment of hemorrhage in accident victims and the monitoring of cancer patients undergoing chemotherapy. Yet another application of the measurement of the hematocrit value involves monitoring patients undergoing kidney dialysis to reduce the potential for incomplete dialysis or excessive dialysis of the patient. Incomplete dialysis leaves toxins behind. Excessive dialysis leads to shock.
[0010] The standard method currently used for measuring the hematocrit value is an invasive method. Typically, a blood sample is obtained from a patient or a donor and centrifuged in a capillary tube to separate red blood cells from plasma. The length of the column in the capillary tube containing red blood cells and the total length of the column in the capillary tube containing both the red blood cells and the plasma are measured, and the ratio of these lengths is the hematocrit value (Hct). See, for example, Morris, M. W., and Davey, F. R., “Basic examination of blood”, in Clinical Diagnosis and Management by Laboratory Methods, Henry, J. B., ed., W. B. Saunders Company, Philadelphia, Pa. (1996), pages 549-559.
[0013] Methods for the non-invasive determination of the hematocrit value include pulse-based methods and direct current-based methods. Pulse-based methods, such as described by Schmitt et al., “Measurement of blood hematocrit by dual-wavelength near-IR photoplethysmography” SPIE Proceedings 1992; 1641:150-161, exhibit problems in the case of individuals having low peripheral perfusion.
[0014] Non-invasive measurement of hematocrit value was recently reported (Wu et al., “Non-invasive determination of hemoglobin and hematocrit using a temperature-controlled localized reflectance tissue photometer” Analytical Biochemistry 2000; 287:284-293, and Zhang et al., “Investigation of noninvasive in vivo blood hematocrit measurement using NIR reflectance spectroscopy and partial least squares regression” Applied Spectroscopy 2000: 54:294-299). Zhang et al. describes a method for determining the hematocrit value in vivo during cardiac bypass surgery. Zhang et al. reported that the temperature of the patient was found to change during surgery. A high number of wavelengths in the near-infrared region of the electromagnetic spectrum and a partial least squares regression analysis were used in an effort to minimize the effect of temperature change on the hematocrit value calculated by this method. Although the device and method described by Zhang et al. provide good calibration and prediction for a given patient during surgery, establishing a model to predict the hematocrit values across more than one patient was less successful. Systematic bias between patients was observed. Part of the observed variations was due to changes in subjects' skin temperature. A method for the determination of concentration of hemoglobin and the hematocrit value is described in WO 01/87151. Steuer et al., U.S. Pat. No. 6,266,546, describes an optical method for the determination of the hematocrit value that uses either the AC or the DC component of the signal at wavelengths of 805 nm and 1300 nm. The possibility of using the same device for determination of oxygen saturation at wavelengths of 660 nm and 805 nm is also disclosed.
[0042] FIG. 1 is a block diagram that describes the apparatus of this invention.
[0043] FIG. 2A is a perspective view of the component of the patient-interface module of the apparatus of this invention that contains the optical probe.
[0044] FIG. 2B is side view in elevation of a cross-section of the apparatus shown in FIG. 2A.
[0045] FIG. 2C is a perspective view, greatly enlarged, of the optical probe of the apparatus shown in FIG. 2A.
[0046] FIG. 3 is a flowchart depicting the steps for the determination of hematocrit and vital signs according to the method of this invention.
[0047] FIG. 4A shows the effect of a venous occlusion (130 mm Hg) on the change in optical signals, measured at a distance of 1.86 mm from the source of light. The source of light had wavelengths of 660 nm, 735 nm, 810 nm, and 890 nm.
[0048] FIG. 4B shows the effect of total occlusion (170 mm Hg) on the change in optical signals, measured at a distance of 1.86 mm from the source of light. The source of light had wavelengths of 660 nm, 735 nm, 810 nm, and 890 nm.
[0049] FIG. 5A is a graph showing the intensity of the reflected light from the forearm of a human subject and at a sampling distance of 1.86 mm as a function of time. The temperature of the skin was maintained at 41° C. The source of light had a wavelength of 590 nm. Signals were collected over a period of three minutes.
[0050] FIG. 5B is a graph showing an expanded portion of FIG. 5A, the portion extending from the 100-second point to the 150-second point.
[0051] FIG. 5C is a plot of the calculated Fourier Transform of the amplitude of the reflected light signal displayed in FIG. 5A.
[0052] FIG. 6A shows the optical signal collected at 1.86 mm from the light introduction site, recorded over a 10-second period. The pulse is superimposed on the low-frequency vasomotion and breathing frequency. Noise spikes are also noticeable.
[0053] FIG. 6B shows the same signal as shown in FIG. 6A, digitally filtered to eliminate the long-term motions and the noise spikes.
[0054] FIG. 6C shows the digitally filtered signal of FIG. 6B, but normalized by dividing the signal by the mean amplitude of the pulses.
[0055] FIG. 6D shows the identification of peaks and vallys for calculating the cardiac pulse rate.
[0056] FIG. 7A shows a plot of the calculated change of concentration of oxygenated hemoglobin, after a venous occlusion (130 mmHg) (upper trace) and release of pressure, and a total occlusion (170 mm Hg) (lower trace) of a human finger and release of pressure. Temperature of the skin was maintained at 38° C.
[0057] FIG. 7B shows a plot of the calculated change of concentration of deoxygenated hemoglobin, after a venous occlusion (130 mmHg) (upper trace) and release of pressure, and a total occlusion (170 mm Hg) (lower trace) of a human finger and release of pressure. Temperature of the skin was maintained at 38° C.
[0058] FIG. 7C shows a plot of the calculated change of concentration of total hemoglobin, after a venous occlusion (130 mmHg) (upper trace) and release of pressure, and a total occlusion (170 mm Hg) (lower trace) of a human finger and release of pressure. Temperature of the skin was maintained at 38° C.
[0059] FIG. 8A shows a plot of the optical signal as a function of time as cuff pressure is varied from 200 mm Hg to 50 mm Hg.
[0060] FIG. 8B shows a plot of the cuff pressure as a function of time as cuff pressure is varied from 200 mm Hg to 50 mm Hg.
[0087] FIG. 1 is a block diagram of an apparatus for carrying out the method of this invention. The components of the block diagram set forth the functions performed by the apparatus 10. The apparatus 10 comprises a patient interface module 12 and a control module 14. The patient interface module 12 comprises a pressure application module 16, an optical measurement module 18, and a plug-in bay 19. The patient interface module 12 has the function of providing points of contact of the pressure application module 16 and the optical measurement module 18 with a body part to obtain measurements of vital signs and optical signals. The control module 14 comprises a computational module 20, an alarm module 22, a communication module 24, and a plug-in bay 26. The control module 14 has the function of providing power and control signals to pressure application module 16 and the optical measurement module 18, pressure control elements, and temperature control elements and receiving signals collected from the optical measurement module 18. The pressure application module 16 performs the function of applying pressure of varying magnitudes to a body part to induce measurable changes in optical signals. A representative example of a pressure application module is an inflatable cuff that can be applied to an arm or a finger. The optical measurement module 18 is an integrated structure comprising at least one optical sensor. An embodiment of an optical sensor is shown in FIGS. 2A, 2B, and 2C and described later. The at least one optical sensor is capable of performing optical measurements of tissue, which measurements are used to calculate the concentration of hemoglobin, the hematocrit value, cardiac pulse rate, blood pressure, and other vital signs. The optical sensors in the optical measurement module 18 can also monitor changes in the hematocrit value and vital signs for patients who are at high risk of having postoperative complications. The pressure application module 16 and the optical measurement module 18 are supplied power through the plug-in bay 19 and are interconnected by means of the plug-in bay 19. The computational module 20 performs the function of performing calculations to compute the concentration of components of the blood and the values of the vital signs. A representative example of the computational module 20 is a personal computer or electronic boards that have microprocessors along with means having the ability to store data in electronic form and the means for communicating that data to other computational devices. The alarm module 22 performs the function of attracting the attention of a nurse or physician or other health care giver to changes in the patient's health status. Representative examples of the alarm module 22 include, but are not limited to, an audible signal or a blinking light. The communication module 24 performs the function of communicating data from the patient from the control module 14 to a nurse's station or a physician's office or to the location of some other health care giver. Representative examples of the communication module 24, include, but are not limited to, a wired connection, a fiber optic connection, or a wireless connection. The computational module 20, the alarm module 22, and the communication module 24 are supplied power through the plug-in bay 26 and are interconnected by means of the plug-in bay 26. The plug-in bay 19 and the plug-in bay 26 are also interconnected.
[0088] A representative embodiment of the apparatus of this invention is illustrated in FIGS. 2A, 2B, and 2C. The apparatus 100 is in the form of a clamp that is capable of surrounding and securely attaching to a finger, designated in FIG. 2B by the letter “F”. The lower part 102 of the apparatus 100 and the upper part 104 of the apparatus 100 are connected by a hinge 106. Handles 108 and 110 can be used to move the lower part 102 of the apparatus 100 toward the upper part 104 of the apparatus 100 or to move the lower part 102 of the apparatus 100 away from the upper part 104 of the apparatus 100. It is preferred that both the interior surface 112 of the lower part 102 of the apparatus 100 and the interior surface 114 of the upper part 104 of the apparatus 100 be concave to easily accommodate a finger. An optical probe 116 is fixed onto the interior surface 112 of the lower part 102 of the apparatus 100. The optical probe 116 is substantially similar to the optical probe described in WO 99/59464, incorporated herein by reference. It is preferred that the lower part 102 of the apparatus 100 be biased toward the upper part 104 of the apparatus 100 by a biasing means (not shown) so that contact between the finger and the optical probe 116 can be securely maintained as optical measurements are performed. A biasing means suitable for this purpose is a spring or a strip of elastic material. A detector 118 for detecting light transmitted through the finger and detection electronics (not shown) are fixed onto the interior surface 114 of the upper part 104 of the apparatus 100.
[0089] The optical probe 116 comprises a light introduction fiber 120 for introducing light to the finger from a source of light (not shown). The source of light must be capable of generating light at at least two wavelengths. Light that is suitable for use in the apparatus 100 of this invention has wavelengths ranging from about 500 nm to about 2000 nm, inclusive. Light is introduced into the tissue of the finger, and light reflected from the tissue of the finger is collected by a plurality of light collection fibers 122, 124, and 126. Each of the light collection fibers 122, 124, and 126 is positioned at a specified distance from the light introduction fiber 120. The light introduction fiber 120 is connected to the source of light, which is preferably housed in the lower part 102 of the apparatus 100. The light collection fibers 122, 124, and 126 are connected to a set of detectors, amplifiers, and a signal processing board, all of which are also preferably housed in the lower part 102 of the apparatus 100. The set of detectors, amplifiers, and signal processing boards can be housed at a location remote from the apparatus 100. The power input to the apparatus 100 and the signal put out by the apparatus 100 are routed through cables (not shown) to the control unit (not shown). The light introduction fiber 120 and the light collection fibers 122, 124, and 126, sources of light, and detectors housed in the lower part 102 of the apparatus 100 are used to perform optical measurements to obtain data needed to calculate, in a continuous manner, the oxygenation status of blood in the tissue of the finger, the concentration of the different components of hemoglobin, and the change in the concentration of hemoglobin. The components of hemoglobin are oxygenated hemoglobin (HbO2), deoxygenated hemoglobin (RHb), and total hemoglobin, which is the sum of the oxygenated hemoglobin and deoxygenated hemoglobin. Other parameters, such as oxygen consumption in the tissue, can also be calculated from the data collected by means of the optical probe 116. The hematocrit value can be calculated from the measured concentration of hemoglobin or the change in concentration of hemoglobin by a commonly used multiplication factor.
[0090] In a preferred embodiment of this invention, the optical probe 116 is set in a metal disc 128, the temperature of which can be controlled, to allow optical measurements to be carried out at different cutaneous temperatures. The optical probe 116 will sample tissue layers to a depth of approximately 2 mm, when the separation between the light introduction fiber 120 and one of the light collection fibers 122, 124, and 126 is approximately two mm.
[0092] The upper part 104 of the apparatus 100 has a single detector 118, such as a silicon photodiode, for the measurement of light transmitted through the finger. Light transmitted through the finger can be used to calculate arterial oxygen saturation and the cardiac pulse rate. While the optical probe 116 will sample tissue layers to a depth of approximately 2 mm, the signal collected and detected in a transmission mode will have passed through the entire vascular bed of the finger, and thus, will have a larger change in magnitude upon change in blood volume during the pulse than would be expected in the reflectance mode. The same source of light as is used for reflectance measurements can be used for measurement of transmitted light. Measurements in the reflectance mode and measurements in the transmission mode can be carried out simultaneously, if desired.
[0093] The apparatus 100 of this invention can be used to monitor fast periodic actions, such as the cardiac pulse, and slow periodic actions, such as breathing rate and the periodic motion resulting from the collective oscillations in the cutaneous vascular system. Both types of motions, which lead to periodic changes in the optical signal, can be detected and measured by the apparatus 100 of this invention.
[0102] FIG. 3 is a flowchart depicting the steps for the determination of the hematocrit value and vital signs, or the change in the concentration of hemoglobin and vital signs, according to the method previously described. FIG. 3 also shows the preliminary steps of (1) initially calibrating the apparatus and (2) allowing the temperature of the body part to equilibrate.
[0105] The intensity of the light reflected or transmitted through tissue can be expressed by Beer's law, where OD=−log10 I/Io, where Io represents the intensity of the light introduced into the tissue and I represents the intensity of the light reflected from or transmitted through the tissue.
OD=εHbO2(1.6165×10−3)+εRHb(0.5489×10−3) (1)
[0107] where εHbO2 represents the molar extinction coefficient in M−1 cm−1 for oxygenated hemoglobin and ERHb represents the molar extinction coefficient in M−1 cm−1 for deoxygenated hemoglobin, and the number 1.6165×10−3 is the molar concentration for oxygenated hemoglobin and the number 0.5489×10−3 is the molar concentration for deoxygenated hemoglobin, as calculated for a concentration of 14.6 gm/dL hemoglobin, with the assumption that oxygenated hemoglobin comprises 75% of total hemoglobin and deoxygenated hemoglobin comprises 25% of total hemoglobin.
[0108] The blood content of the tissue will change over a short period of time as a result of occlusion, bleeding, or hemodialysis. The change in the optical density from the initial concentration of hemoglobin (or the initial hematocrit value) to that at a subsequent value (ΔOD)t, as a result of occlusion, bleeding, or hemodialysis, at any wavelength, can be expressed as:
[0110] Δ(OD)t represents the difference in the measured optical density at a given wavelength and at time t,
[0111] εHbO2 represents the molar extinction coefficient of oxygenated hemoglobin at the same wavelength,
[0112] Δ[HbO2])t represents the change in the concentration of oxygenated hemoglobin at time t,
[0113] εRHb represents the molar extinction coefficient of reduced hemoglobin (deoxygenated hemoglobin) at the same wavelength, and
[0114] (Δ[RHb])t represents the change in the concentration of reduced hemoglobin (deoxygenated hemoglobin) at time t, wherein the change in concentration of hemoglobin results from occlusion, bleeding, or the effect of hemodialysis.
[0115] The coefficients in the expression are the values of extinction coefficients of oxygenated hemoglobin and deoxygenated hemoglobin and are available in the literature. These coefficients vary as a function of wavelength according to the following relationships: 1Δ(OD)t at 590 nm=14×103(Δ[HbO2])t+28×103(Δ[RHb])t(3)Δ(OD)t at 660 nm=0.32×103(Δ[HbO2])t+3.2×103(Δ[RHb])t(4)Δ(OD)t at 735 nm=0.41×103(Δ[HbO2])t+1.10×103(Δ[RHb])t(5)Δ(OD)t at 810 nm=0.86×103(Δ[HbO2])t+0.72×103(Δ[RHb])t(6)Δ(OD)t at 890 nm=1.2×103(Δ[HbO2])t+0.74×103(Δ[RHb])t(7)Δ(OD)t at 935 nm=1.2×103(Δ[HbO2])t+0.73×103(Δ[RHb])t(8)
[0116] The value of the change in the concentration of oxygenated hemoglobin (Δ[HbO2]) and the value of the change in the concentration of deoxygenated hemoglobin (Δ[RHb]) can be obtained by solving any two of the foregoing equations, (3) through (8). The change in the concentration of total hemoglobin resulting from occlusion, bleeding, or changes during dialysis, can be determined by the equation:
Δ[Total Hb]t=Δ[HbO2]+Δ[RHb]t (9)
[Total Hb]t=Initial[Total Hb]±Δ[Total Hb]t (10)
[0118] The Δ(OD)t values, which are determined at several time intervals, starting from the onset of occlusion, the beginning of surgery, the beginning of post-operative care, or the beginning of a hemodialysis session, are used to calculate the change in concentration of total hemoglobin (±Δ[Total Hb]t) resulting from occlusion, bleeding, or changes during dialysis by means of equation (9). The value of the concentration of total hemoglobin at the end of any other time interval, starting from the onset of occlusion, start of surgery, start of post-operative care, or start of a hemodialysis session, can then be determined by using equation (10).
[0120] FIG. 4A shows the effect of occlusion on the optical signal under the following conditions: 130 mm Hg pressure, wavelengths of 660 nm, 735 nm, 810 nm, and 890 nm, light collected at a site at a distance of 1.86 mm from the light introduction site. FIG. 4B shows the effect of occlusion on the optical signal under the following conditions: 170 mm Hg pressure, wavelengths of 660 nm, 735 nm, 810 nm, and 890 nm, light collected at a site at a distance of 1.86 mm from the light introduction site. In this study, a blood pressure cuff was placed on the arm of a subject who was sitting in a clinical chair, the subject's left arm resting on the arm of the chair. The subject's index finger was placed in contact with the optical probe. The temperature in the aluminum disc was maintained at 38° C. The temperature of the finger was allowed to equilibrate with the disc for two minutes before measurements were begun. Data, i.e., optical signals, were collected for three minutes at the rate of 22 measurements of data per second. The data are presented as a plot of optical density (OD) vs. time in seconds. The pressure in the cuff was maintained at zero mm Hg for the first 60 seconds. The pressure was increased to 130 mm Hg, which was higher than the diastolic pressure and lower than the systolic pressure for the subject, and maintained at this pressure for 60 seconds. The pressure was released instantaneously, and data were collected for the remainder of the 180-second duration of the experiment. During the experiment the cardiac pulse rate, the oxygen saturation value, and a perfusion parameter were recorded by means of a Hewlett-Packard vital signs monitor having a plethysmographic sensor attached to the subject's middle finger. The measurement was repeated several times at different pressures, ranging from below the diastolic pressure to above the systolic pressure. The systolic pressure was defined as the pressure at which the pulse disappeared. For applied pressures below the systolic blood pressure, the back-flow of venous blood to the heart is stopped as a result of closing the venous path that returns blood to the heart, thus leading to a state of venous occlusion (FIG. 4A). The intensity of the reflected light decreased, i.e., the measured optical density increased (because pooled blood increases light absorption) until the optical density reached a plateau. As the pressure was reduced, the optical density returned to approximately the initial value of the optical density, i.e., the value prior to occlusion. When the arm of a subject is occluded at a pressure above the systolic blood pressure, arterial blood flow from the heart to the limb is stopped because of closing of the artery, and venous blood flow back to the heart is stopped because of closure of the return venous path, thus leading to a state of total ischemia, which is the state of total occlusion (FIG. 4B). The optical density measured at the finger decreases and reaches a plateau as a result of occlusion. As the pressure is released subsequent to total occlusion, the optical density increases (because pooling of blood increases light absorption) and returns to approximately its initial value, i.e., the value prior to occlusion. Similar response curves are observed at other wavelengths. It is possible to use equations (3) through (8) to calculate the change in the concentrations of oxygenated hemoglobin and deoxygenated hemoglobin and hence, the change in concentration of total hemoglobin at the finger when the arm is occluded.
[0121] The cardiac pulse rate can be determined from the optical signals collected from a body part. The optical signals collected from a human body part over a period of time is a composite of several periodic signals that includes signals arising from the cardiac pulse rate, breathing rate, and vasomotion, as shown in FIG. 5A. FIG. 5A is a graph showing the intensity of the reflected light from the forearm of a human subject at 590 nm and at a sampling distance of 1.86 mm as a function of time. Signals were collected over a three-minute period. The temperature of the skin was maintained at 41° C. FIG. 5B is a graph showing a the portion of FIG. 5A from the point of time of 100 seconds to the point of time of 150 seconds. FIG. 5C is a plot of the calculated Fourier Transform of the amplitude of the reflected light signal shown in FIG. 5A.
[0122] The cardiac pulse rate can be determined by recording the output of the optical probe over several pulses, over a given period of time. By expanding a portion of FIG. 5A (see FIG. 5B), it can be seen that the cardiac pulse rate is superimposed over other pulses having lower frequency. By performing a Fourier Transform, a plot of the power spectrum can be constructed, which plot shows the cardiac pulse rate at 1.18 Hz (see FIG. 5C) and several low frequency pulses indicative of other oscillations in the vascular system of the skin.
[0123] Alternatively, the cardiac pulse rate can be calculated from the filtered signal. See FIG. 6C. The cardiac pulse rate can be reported as pulses per second by counting the number of peaks or valleys of the filtered pulses over a period of time and calculating the cardiac pulse rate in pulses per minute. This calculation is shown in Example 2.
[0130] 5) determining the intensity of the signal at each peak (Ip) and at each valley (Iv) of each digitally filtered pulse, where
[0131] Ip represents the intensity of the signal at the peak of a cardiac pulse wave, and
[0132] Iv represents the intensity of the signal at the valley of a cardiac pulse wave;
(OD)v=−log (Iv/Io)
(OD)p=−log (Ip/Io)
[0138] 10) calculating the concentration of oxygenated hemoglobin ([HbO2]) and the concentration of deoxygenated hemoglobin ([RHb]) by means of simultaneous equations, where
Δ(OD)λ1=a[HbO2]+b[RHb]
Δ(OD)λ2=c[HbO2]+d[RHb]
[0140] 11) calculating the value of oxygen saturation according to the following equation: 2Arterial oxygen saturation=[HbO2]([HbO2]+[RHb])×100 %
[0141] The coefficients a, b, c, and d in step 10) are the values of the extinction coefficients at the maximum wavelength of the particular LED. The approximation does not take into consideration the finite bandwidth of the LED or the skew of the intensity distribution over the bandwidth. 1
[0146] Referring now to FIGS. 2A, 2B, and 2C, an optical probe suitable for carrying out the method of this invention comprises a set of light emitting diodes (LEDs) that emit light at wavelengths 590 nm, 660 nm, 890 nm, and 935 nm. The output of the LEDs is focused on a light introduction fiber 120 that transmits light from the LEDs to the skin at a light introduction site. Each light emitting diode (LED) can be operated in a modulated current mode by modulating the current input to each LED at a fixed frequency. Alternatively, LEDs can be operated in a constant current mode.
[0147] In this example, LED 1 emits light having a wavelength of 660 nm, a modulation frequency of 1024 Hz, and a half bandwidth of 15 nm. LED 2 emits light having a wavelength of 590 nm, a modulation frequency of 819 Hz and a half bandwidth of 15 nm. LED 3 emits light having a wavelength of 935 nm, a modulation frequency of 585 Hz, and a half bandwidth of 25 nm. LED 4 emits light having a wavelength of 890 nm, a modulation frequency of 455 Hz, and a half bandwidth of 25 nm.
[0150] The optical probe 116 located in the lower part 102 of the apparatus 100 illustrated in FIG. 2A was used in different set-ups to illustrate the ability of the apparatus to perform cardiac pulse rate measurements, arterial blood oxygen saturation measurement, and response of cutaneous blood vessels to occlusion.
[0151] FIGS. 5A, 5B, 5C, and 5D show the results of the steps carried out to calculate the cardiac pulse rate from optical signals. These steps were as follows:
[0157] The cardiac pulse rate was calculated from the first section of the curve at zero occlusion, at both 38° C. and 22° C. The data for a subject with normal perfusion condition are shown in Table 2 (38° C.) and in Table 3 (22° C.). 2
[0158] The average cardiac pulse rate at all separations of the light introduction site from the light collection site and at all wavelengths was 73 pulses per minute at 38° C. 3
[0168] 5) determining the intensity of the signal at each peak (Ip) and at each valley (Iv) of each digitally filtered pulse, where
[0169] Ip represents the intensity of the signal at the peak of a cardiac pulse wave, and
[0170] Iv represents the intensity of the signal at the valley of a cardiac pulse wave;
[0175] 10) calculating the concentration of oxygenated hemoglobin ([HbO2]) and the concentration of deoxygenated hemoglobin ([RHb]) by means of simultaneous equations, where
[0177] 11) calculating the value of oxygen saturation according to the following equation: 3Arterial oxygen saturation=[HbO2]([HbO2]+[RHb])×100 %
[0179] The wavelength pairs 660 nm/810 nm and 660 nm/890 nm yielded oxygen saturation values between 91 and 100 at all separations of the light introduction site from the light collection site. The calculated oxygen saturation (O2 sat) values for a normal subject are shown in Table 4. 4
[0182] The initial hematocrit value is determined either by an invasive method or by a non-invasive method. The optical density at the measurement site is determined at the time the concentration of hemoglobin or the hematocrit value is measured by contacting the optical probe with the body part. The optical density of the tissue of the finger (OD) is determined at the wavelengths 660 nm, 735 nm, 810 nm, and 890 nm at another time t, after the initial measurement. At least two of the following four linear equations are solved to obtain the values of A[HbO2] and A[RHb]. 4Δ(OD)t at 660 nm=0.32×103(Δ[HbO2])t+3.2×103(Δ[RHb])t(4)Δ(OD)t at 735 nm=0.41×103(Δ[HbO2])t+1.10×103(Δ[RHb])t(5)Δ(OD)t at 810 nm=0.86×103(Δ[HbO2])t+0.72×103(Δ[RHb])t(6)Δ(OD)t at 890 nm=1.2×103(Δ[HbO2])t+0.74×103(Δ[RHb])t(7)
Δ[Total Hb]=Δ[HbO2]+Δ[RHb] (9)
[Total Hb]=Initial[Total Hb]±Δ[Total Hb] (10)
[0186] FIG. 7A and FIG. 7B show the change in concentration of oxygenated hemoglobin and the change in concentration of deoxygenated hemoglobin as a result of venous or arterial occlusion. The calculated change in concentration of hemoglobin is shown in FIG. 7C.
[0189] The optical probe of this invention is capable of montoring the systolic blood pressure of a patient and the change in blood pressure as a function of time. FIG. 8A shows a tracing of the optical signal versus time for a human finger as the pressure in a cuff was rapidly increased to 200 mm Hg and slowly decreased to 50 mm Hg over a period of 180 seconds. Pressure was applied to the left arm at the 60-seconds point in time to bring about occlusion. The cuff pressure was increased from zero to 200 mm Hg within approximately two seconds. The pressure was then slowly reduced. A plot of the cuff pressure versus time is shown in FIG. 8B. The rate of pressure reduction was 0.833 mm Hg per second. Upon occlusion of the blood vessels in the left arm, the optical signal increased and remained at a plateau until a pressure of 141 mm Hg was reached and then sharply decreased as the pressure fell below 130 mm Hg. The blood pressure of the subject was measured on the right arm immediately before the study, and the systolic pressure was 134±5 mm Hg. Thus, the inflection point in the plot of optical signal versus time (deflation time) lies at the systolic blood pressure of the subject. Accordingly, it is possible to track the systolic blood pressure of a person using the apparatus and method of this invention.
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