WO2006110488A9

WO2006110488A9 - High efficiency photoplethysmographic sensor with coupling gel

Info

Publication number: WO2006110488A9
Application number: PCT/US2006/012896
Authority: WO
Inventors: Eric W Starr; Bernie F Hete
Original assignee: Ric Investments Llc
Priority date: 2005-04-08
Filing date: 2006-04-07
Publication date: 2007-05-03
Also published as: WO2006110488A3; WO2006110488A2

Abstract

A pulse-oximeter photoplethysmographic sensor unit that includes a light source and a deformable pressure reducing relief layer. The pressure relief layer is disposed proximate to the light source between a subject's extremity and the light source. The relief layer is translucent to the light from the light source. The deformable relief layer is a removable, disposable gel pad attachable to the light source and may include opaque side walls increasing light transmission while reducing ambient light transmission. The relief layer may have an index of refraction between the index of refraction of the light source and that of a subject's extremity. The pulse-oximeter sensor unit may have an attachment mechanism such as a spring biased clam-shell clamp or a wrap around strip.

Description

HIGH EFFICIENCY PHOTOPLETHYSMOGRAPHIC SENSOR WITH COUPLING GEL

[01] Provisional U.S. patent application No. 60/565,708 filed April 26, 2004, entitled, "Medical Devices And Techniques For Rodent And Small Mammalian Based Research," is incorporated herein by reference.

BACKGROUND OF THE INVENTION

1. Field of the Invention

[02] The present invention pertains to a photoplethysmographic sensor, and, in particular, to a photoplethysmographic sensor utilizing an optical coupling material disposed in an optical path of the sensor to improve signal fidelity.

2. Description of the Related Art

[03] Photoplethysmographic sensor units are one type of non-invasive physiologic sensor, and are also called a pulse monitor or a plethysmograph. These devices are well known for use in humans and large mammals and are described in U.S. patent numbers 4,621,643; 4,700,708; and 4,830,014 which are incorporated herein by reference. These units typically incorporate an incandescent lamp or light emitting diode (LED) to trans-illuminate an area of the subject, e.g., an appendage, such as a finger, that contains a sufficient amount of blood flow. The light from the light source disperses throughout the appendage and is transmitted through the appendage. A light detector, such as a photodiode, is placed on the opposite side of the appendage to record the received light that is transmitted through the tissue and is received by a detector for generating a receiving signal. This signal may be broken down into non-pulsitile blood components, non-pulsitile arterial blood component, and pulsitile arterial blood component.

[04] Due to the absorption of light by the appendage's tissues and blood, the intensity of light received by the photodiode is less than the intensity of light transmitted by the light source. Of the light that is received, only a small portion (that effected by pulsitile arterial blood), usually only about two percent of the light received, behaves in a pulsitile fashion. The beating heart of the subject creates this pulsitile behavior. The "pulsitile portion light" is one of the signals of interest, and effectively forms the photo- plethysmograph. The absorption described above can be conceptualized as AC and DC components. The arterial vessels change in size with the beating of the heart. The change in arterial vessel size causes the path length of light to change from d_min to J_max .

This change in path length produces the AC signal on the photo-detector, I_L to I₁₁.

[05] When the light source includes two distinct wavelengths, e.g., two distinct

LED light sources having different frequencies, the blood oxygenation may be obtained as the detected signal amplitude changes in direct relation to the quantity of oxygenated hemoglobin within red blood cells that pass through the light beam. The light-absorptive characteristic of hemoglobin is different for different wavelengths of light, as well as, whether or not the hemoglobin is saturated with oxygen, and is quantified by a parameter known as the extinction coefficient. The extinction coefficient defines the light absorption of a material at a given wavelength for a given concentration of the material and a given path length through the material. These sensors are generally referred to as pulse oximeters. Within the meaning of this application, a photoplethysmographic pulse oximeter has at least two distinct wavelengths of light generally from two distinct light sources, such as two LED sources.

[06] Most pulse oximeters use both a red (typically 660 nm) and an infrared

(typically 940 nm) wavelength of light, because the ability of a detectable signal to be transmitted through the subject's appendage is important. The red and infrared wavelengths are used comparatively to conduct the oxygen saturation measurement. The wavelengths chosen are not arbitrary, but are based on a number of physical characteristics of the light measurement system. Ideally, a designer would desire to have the largest difference in absorption between saturated (oxy-hemoglobin) as compared with non-saturated or reduced hemoglobin for both wavelengths of light. However, this is not the only consideration. There is also the cost, availability, and performance of the photodiode receiver to the light source, typically an LED source. [07] As noted above, in a transmission photoplethysmographic system, light is generated by the light source, e.g., an LED source, it passes into the perfused biological medium, e.g., a finger, then it is received by a receiver, i.e., photodiode, on the other side of the medium. In addition to the extinction coefficient of the saturated and non- saturated hemoglobin, the bandwidth and power of the light source, the optical characteristics of the total transmission path, and the bandwidth of response of the receiver, are all important characteristics of the system. Existing pulse oximeters have a resolution of +/- 2%, with normal measurements ranging between about 88% and 100% oxygen saturation. Therefore, there is a need to improve the accuracy of such devices. There is a also need to improve the transmission of light from source to receiver and to improve the resulting signal.

SUMMARY OF THE INVENTION

[08] Accordingly, it is an object of the present invention to provide a photoplethysmographic sensor that overcomes the shortcomings of conventional photoplethysmographic sensors. This object is achieved according to one embodiment of the present invention by providing a photoplethysmographic sensor that includes a light member support with at least one light source, such as one or more LED light sources, operatively coupled to the light member support. A first layer of a coupling gel is disposed proximate to the light source. A sensor support with a light receiving sensor is operatively coupled to the sensor support. The photoplethysmographic sensor also includes an attachment mechanism that is adapted to attach the sensor to an extremity of a subject such that the first layer of the coupling gel is disposed between the light source and the subject's extremity.

[09] The photoplethysmographic sensor unit according to one aspect of the invention includes a removable release layer adjacent the first layer of coupling gel. The release layer is removed prior to attaching the sensor unit to a subject. In a further embodiment, the attachment mechanism is a clam-shell clamp having a biasing member adapted to bias the light member support, the sensor, or both support toward a subject's extremity. Alternatively, the attachment member is a strip configured to wrap around a subject's extremity. A further alternative is to use two independent pads, each of which is secured to the patient, as the attachment member.

[10] According to one aspect of the invention, the coupling gel has an index of refraction generally between, e.g., an average, the index of refraction of the light source covering, e.g., the clear plastic that covers the light emitting constituent, and the index of refraction of the skin or hair that covers the subject's extremity. The present invention also contemplates that the coupling gel has an index of refraction generally equal to either the index of refraction of a subject's extremity or of the light source. The photoplethysmographic sensor unit according to one aspect of the invention may further include a second layer of coupling gel disposed proximate to the light receiving sensor.

[11] The photoplethysmographic sensor unit according to one aspect of the invention may provide a light source, such as an LED, having a wavelength of less than 620 or even 600 nanometers. Further, the light source may have a wavelength between about 405 and about 572 nanometers. Specifically, the light source may have a wavelength selected from the group of about 405 nanometers, about 430 nanometers, about 463 nanometers, about 507 nanometers, about 524 nanometers, and about 572 nanometers, and wherein the light sources are selected to each have a significantly high HbO₂ extension coefficient as well as a significant separation in the extinction coefficients OfHbO₂ and Hb. The term "about" within the meaning of this application in referring to the above wavelengths, will be +/- 3 nanometers, with this variation allowing for commercial availability.

[12] In the photoplethysmographic sensor unit of the present invention, the gel forms a deformable pressure reducing relief layer that is disposed proximate to the light source, wherein the relief layer is translucent to the light from the light source and may be formed as a removable, disposable gel pad attachable to the light source. The gel pad may include opaque side walls to collimate the light and reduce ambient light transmission. The invention may further include a second gel pad or layer of coupling gel disposed proximate to the light receiving sensor.

[13] These and other objects, features, and characteristics of the present invention, as well as the methods of operation and functions of the related elements of structure and the combination of parts and economies of manufacture, will become more apparent upon consideration of the following description and the appended claims with reference to the accompanying drawings, all of which form a part of this specification, wherein like reference numerals designate corresponding parts in the various figures. It is to be expressly understood, however, that the drawings are for the purpose of illustration and description only and are not intended as a definition of the limits of the invention. As used in the specification and in the claims, the singular form of "a", "an", and "the" include plural referents unless the context clearly dictates otherwise.

BRIEF DESCRIPTION OF THE DRAWINGS

[14] FIG. Ia is a cross-sectional view of a photoplethysmographic sensor unit according to a first embodiment of the present invention attached to a subject's appendage; [15] FIG. Ib is a cross-sectional view of a photoplethysmographic sensor unit according to a second embodiment of the present invention attached to a subject's appendage; [16] FIG. 2 is a schematic plan view of the photoplethysmographic sensor unit of FIG. Ia; [17] FIG. 3 is a cross-sectional view of a replaceable, detachable, disposable, photoplethysmographic sensor unit gel pad attachment for the photoplethysmographic sensor unit of FIG. Ia; [18] FIG. 4 is a side view of the photoplethysmographic sensor unit of FIG. Ia shown prior to attachment to the subject with the gel pad attachment of FIG. 3 secured and a user release layer in place; [19] FIG. 5 is a schematic section, similar to FIG. Ia, of a photoplethysmographic sensor unit according to a third embodiment of the present invention attached to a subject's appendage; [20] FIG. 6 is a cross-sectional view of light incident on the interface between optically different materials; [21] FIG. 7 is a graph of the extinction coefficients of oxyhemoglobin and reduced hemoglobin verses light wavelength; and [22] FIG. 8 is a graph of the difference in the extinction coefficients shown in

FIG. 7 verses light wavelengths.

DETAILED DESCRIPTION OF THE EXEMPLARY EMBODIMENTS

[23] FIG. 1 a is a cross-sectional view of a photoplethysmographic sensor unit

10 according to a first embodiment of the present invention showing the sensor unit attached to a subject's appendage, namely a finger 12. Photoplethysmographic sensor unit 10 includes a light member support 14 with one or more light sources 16, such as LED elements, operatively coupled to the light member support. Light member support 14 may be formed integral with light sources 16, such as in commercially available LED source assemblies in which the light member support 14 is essentially the housing for the light sources 16. The present specification separates the light source and the light member support by their respective function, one being for mechanical or structural support and the other providing the illumination function. Wires 18 extend to light source 16, as known in the art. A first layer of a coupling gel, also called a gel pad, 20 is disposed proximate to the light source 16. Gel pad 20 is disposed in the optical path of the sensor unit, so that light emitted from the source must pass through the gel pad. The present invention contemplates that gel pad 20 includes an opaque side wall 22, as described below.

[24] In an exemplary embodiment of the present invention, the gel pad is formed from the gel material manufactured by Dow Corning under part # 7-9800, which is a silicon-based gel. It is to be understood, however, that the present invention contemplates using any suitable gel material. Of course, a biocompatible material would be perhaps best suited for this application so as not to irritate the skin of the user.

[25] Sensor unit 10 further includes a sensor support 24 with one or more light receiving sensors 26, e.g., photodiodes, operatively coupled to the sensor support. Sensor support 24 may be formed integral with light sensors 26, such as in commercially available detectors in which the sensor support is essentially the housing for the light receiving sensor. This specification separates the sensor support and the light receiving sensors by their respective function, one being for mechanical support and the other functioning as the electronic detector. Wires 28 extend to light sensor 26 as known in the art. Sensor unit 10 further includes an attachment mechanism, such as wrap 30, adapted to attach the sensor unit to an extremity of a subject, e.g., finger 12, such that the first layer of coupling gel 20 is disposed between light source 16 and the subject's extremity 12. In the illustrated exemplary embodiment, an optional second layer of a coupling gel, also called a gel pad, 40 is disposed proximate to light sensor 26. The present invention contemplates that gel pad 40 includes an opaque side wall 42, as described below.

[26] Wrap 30, shown in plan view in FIG.2, encircles finger 12, to secure sensor unit 10 in place on the user. An alternative to wrap 30, not shown in the drawings, is use two independent pads. Such independent pads may be formed by severing wrap 30 between gel pads 20 and 40. Each pad is secured to the surface of the user using any conventional securing technique. A further simplification of the sensor unit shown in FIG. Ia is to use only gel pads 20 and 40 without the opaque sides 22 and 42, wherein the backing element 30 (no longer called a wrap as it does not encircle the appendage) can be formed as an opaque cloth to block ambient light. Gel pads 20 and 40, without the opaque sides 22 and 42, may be used with wrap 30 essentially as shown in FIG. Ia.

[27] FIG. Ib is a schematic section of a further modified photoplethysmographic sensor unit 10' according to the present invention with the sensor unit 10 attached to a subject's finger 12. Photoplethysmographic sensor unit 10' of FIG. Ib is essentially the same as that shown in FIG. Ia and includes one or more light sources 16, wires 18, a first layer of a coupling gel 20 disposed proximate to the light source 16, one or more light receiving sensors 26, wires 28, wrap 30, and an optional second layer of a coupling gel 40. The sensor unit of FIG. Ib differs from sensor unit 10 of FIG. Ia in that the support structure (items 14 and 26 in FIG. Ia) is formed or molded integrally with the opaque portions 22 and 42. The operation of the device is unchanged.

[28] FIG. 3 is a schematic, cross-sectional view illustrating that gel pads 20 and

40 may be formed as a replaceable, detachable, disposable, photoplethysmographic sensor unit gel pad attachment 43 for the photoplethysmographic sensor unit of FIGS. Ia or Ib. Attachment 43 is separate from the other components of the sensor unit until the time of use, where it is assembled with light member support 14 and sensor support 24. In the illustrated exemplary embodiment, gel pad attachment 43 includes two release layers, a first unit release layer 44 that can be removed so that gel pads 20 and 40 can be secured to light member support 14 and sensor support 24, respectively, as shown in FIG. 4. FIG. 4 is a side view of the photoplethysmographic sensor unit 10 of FIG. Ia shown prior to the attachment to a subject, with gel pad attachment 43 of FIG. 3 secured to light member support 14 and sensor support 24 and with the second release layer, also called a user release layer 46, still in place. Release layer 46 is removed and the sensor unit is attached to the subject's appendage, i.e., finger 12 as shown in FIG. Ia.

[29] FIG. 5 is a schematic section of a photoplethysmographic sensor unit 10" according to a second embodiment of the present invention shown attached to a subject's appendage, specifically finger 12. This embodiment differs from the embodiment of FIGS. Ia and Ib only in that the attachment mechanism is a clam shell 32 type of mechanism rather than a wrap. The arms or sides 33a and 33b of clam shell 32 are biased by a spring 34 into engagement with the finger. The clam shell construction is believed to be well known in the art and no further details are needed.

[30] A first advantage of sensor unit 10 of the present invention is the reduction of the optical impedance of the light path from source 16 to receiver 26. Whenever a wave front of any type passes between two different media with a planar interface, some of the wave front is reflected, and the remainder is transmitted. A schematic of this appears in FIG. 6, where the variables H]₁ and η₂ represent the refractive indices of the two materials. The ratio of reflection to transmission is a function of the difference in magnitude of the impedance to transmission of each medium for the type of wave. For light, the optical impedance is a function of the index of refraction of the materials. It has further been shown that the transmitted intensity of light (electromagnetic radiation) at a planar interface between materials of different optical impedance is given as:

[31] Expression (1) above evidences that the transmitted light intensity is a function of both the angle of incidence of the light, as well as the magnitude of the difference between the indexes of refraction of the two materials. Thus, as the angle of incidence moves further away from the surface normal, less light is transmitted. Also, the greater the difference between the indexes of refraction between the materials, the less light will be transmitted. This latter point is easier to see if equation (1) is re-written for normal incidence, as follows:

¹ T 4η

(2)

It is notable that the difference in the angles θ and φ is also a function of relative indexes of refraction. This relation is known as Snell's Law, and is given as: sin(^) = ^sin(ø) . (3)

To understand the nature of index of refraction, it is helpful to point out that it is defined as the ratio of the speed of light in a vacuum, to the speed of light in the given material or

η =- - (4) v

[32] From equation (4), the index of refraction of a vacuum can be shown as

1.00000. Air has an index of refraction that is approximately 1.00. Meanwhile, most materials have indexes of refraction less than 2.5. However, the index of refraction of skin and hair is in the neighborhood of 1.50. These numbers in equation (2) show that if there is an air gap between the light source 16, i.e., the LED, and the tissue 12, the transmitted radiation is 64% of the incident radiation, which represents a loss of 36% of the incident radiation. If it is further assumed that light member support 14, also called a housing, containing light source 16 also has a refractive index of nearly 1.50, then there is another loss of 36%, which portends a total loss of 59% (1.00-0.64²) of the incident light transmitted into the tissues 12. This loss in transmission is due to reflection and it is certain that some of the light is re-transmitted due to multiple reflections, so the loss mentioned above is a maximum. However, it is still significant.

[33] Sensor unit 10 provides a solution to this problem by providing a coupling medium through gel pads 20 and 40 that will have an index of refraction that is matched more closely with the skin (or hair) of the appendage, and of the light source 16 as well, wherein coupling gel pads 20 and 40 have an index of refraction generally between the index of refraction of the light source and the index of refraction of a subject's extremity. In one embodiment of the invention, coupling gel pad 20 has an index of refraction generally equal to the index of refraction of a subject's extremity, generally about 1.50. Even if gel pads 20 and 40 cannot be selected to have an index of refraction generally between the index of refraction of the light source and the index of refraction of a subject's extremity, the gel pads will certainly provide a closer match than that from air.

[34] Even where ultrasound gel was used to form gel pads 20 and 40 in experimental sensor units 10, significantly greater light transmission was observed than through air (no gel pads). Ultrasonic coupling gel is just one implementation for gel pads 20 and 40 to obtain this advantage. A viscous liquid could also be used. However, the preferred material is a very soft visco-elastic translucent material, such as that used in the ventilation gel mask sold by Respironics, Inc. In experimental testing of sensor unit 10 with gel pads 20 formed of visco-elastic translucent material, such as that used in the ventilation gel mask sold by Respironics, Inc., the quantity of light that reaches the skin surface is greatly enhanced, hi fact, it is very difficult to see any light illuminating the skin. The present inventors concluded from these experiments that matching the optical impedance along the light pathway from source to tissue to receiver can greatly increase the conservation of the light that passes through the skin.

[35] A second advantage of sensor unit 10 according to the present invention is that the gel pads disposed between the skin and the sensor wrap acts as a strain relief. Many transmittance photoplethysmographic sensors in existence today use some sort of adhesive wrap, such as wrap 30, or biased clam shell, such as clam shell 34, to secure the emitter and detector on an appendage, such as finger 12. As wrap 30 or clam shell 34 is applied tighter and tighter, the arterial pulsitile behavior within the appendage is attenuated. This can be significant in infants and children, as the total flow volume is lower. Essentially, the gel material used in gel pads 20 and 40 placed between wrapping material 30, or clam shell 34, and the skin deforms under pressure.

[36] The deformation (strain) of gel pads 20 and 40 reduces the pressure applied to the arteries and thus reduces the attenuation of the pulsitile signal. A further advantage of sensor unit 10 having gel pads 20 and 40 formed of gel, such as that in the gel mask, is that because the gel is sticky, it could be used as an adhesive, rendering the need for contact pressure of the source and receiver against the skin moot. An additional benefit of better light transmission and use of a coupling material, such as gel pads 20 and 40, is that the power input to the light source can be reduced, providing the ability to have more dynamic range available for transmission paths of greater length or higher opacity. Further, the design requirement for dynamic range of the light source can be narrower, since the transmitted light would be more controlled. A coupling gel, such as gel pads 20 and 40, may also allow the use of skin lubricants that would prevent sores that might result from lengthy skin contact. A coupling gel medium can also conserve the light within its material boundary through the use of opaque sides 22. The index of refraction matching discussed above will also apply to receiver 26 side, as well as on the source side.

[37] Another advantage of sensor unit 10 is that by utilizing a colored gel material in gel pads 20 and 40, the system will reduce ambient light transmission as will opaque side walls 22 and 42. Ambient light noise is a common problem associated with existing pulse-oximetry. One advantage of gel pads 20 and 40 is the selection of gels that that only pass light in the wavelengths used in the system, i.e., the wavelength of the light sources. The easier approach to minimizing ambient light transmission is to cover patient interface with opaque material 22 and 42 to block ambient light.

[38] Returning to the disposable gel pad attachment shown in FIG. 3, in a further embodiment of this design, the present invention contemplates designing the shaped-aperture gel pad as a disposable part that attaches via snaps or other suitable attaching technique into housing 14 or 24 containing light source 16 or receiver 26. This embodiment allows light source 16 and receiver 26 to be more easily reusable, while maintaining the disposable patient interface. Another embodiment of the present invention contemplates a disposable gel pad in which the gel is a material that becomes solid, or at least less liquid, at cooler temperatures, such as temperatures present in conventional refrigerators (or even room temperature), but becomes liquid either at room temperature or at skin temperature. This makes the pad easier to use, because the liquid gel could not flow until it was warmed on the skin.

[39] One consideration in the design of sensor unit 10 is to improve the signal to noise (S/N) ratio by using a wavelength for light source 16 that has a greater separation of extinction coefficients between oxy hemoglobin and reduced hemoglobin. As an example, from FIG. 7 (Extinction coefficient units are L/mmol/cm.) it can be seen that the largest spread appears to be around a wavelength of 660 nm, the wavelength commonly used in pulse oximetry systems. However, the commonly used graph in FIG. 7 (Extinction coefficient units are L/mmol/cm.) is a logarithmic plot. Thus, the difference in wavelength is not that large as a matter of linear magnitude. This is demonstrated in the graph of FIG. 8, in which the extinction coefficients of oxyhemoglobin and reduced hemoglobin have been subtracted and are plotted linearly versus wavelength.

[40] It is apparent from FIG. 8 that the largest separation occurs in the high blue region near 420 nm. A sensor using this band would provide a much greater S/N than would be seen in the red region. Sensor unit 10 of the present invention, when used as a pulse oximeter, requires two distinct light sources, as two distinct wavelengths are always required to obtain oxygenation levels. The present invention contemplates that for at least small mammal applications, at least one light source has a wavelength of less than 620 or even 600 nanometers. More preferably, one light source has a wavelength between about 405 and about 572 nanometers, and most preferably, at least one light source has a wavelength selected from the group of about 405 nanometers, about 430 nanometers, about 463 nanometers, about 507 nanometers, about 524 nanometers, and about 572 nanometers.

[41] The light sources are, preferably, selected to each have a significantly high

HbO₂ extension coefficient as well as a significant separation in the extinction coefficients OfHbO₂ and Hb. As noted above, the term "about" within the meaning of this application contemplates a tolerance of +/- 3 nanometers, with this variation allowing for commercial availability. The other light source may have any wavelength, including those greater than 620 or 600 nanometers, provided that the other wavelength is selected to have different signals than the first wavelength, i.e., two distinct lights. Preferably both light sources have wavelengths less than 620 nanometers.

[42] A supposition of the present invention is that sensor unit 10 using wavelengths of light with higher HbO₂ extinction coefficients will result in bigger pulsitile signals for thin appendages, such as the paws of mice and rats, the web of an adult human hand, or a human earlobe, but not for the thick appendages such as the fingers of adult humans. The preferred wavelengths discussed herein may be beneficial for neo-natal photoplethysmograph sensor units 10 as well. Notice that because the HbO₂ extinction coefficient is a relatively small number, the AC (pulsitile) signal is relatively small for a given DC (mean). It can be demonstrated that by using a light wavelength with a bigger HbO₂ extinction coefficient, the resulting pulsitile signal may be increased. However, when the absorption rate increases, not enough light is transmitted through large subjects (adult human fingers).

[43] In computer modeling, even after increasing the input light to makeup for a factor of 100 times an increase in absorption, the signal could not be resolved. Again, this is because of the extremely high absorption of a relatively thick appendage. This is what occurs when wavelengths of light with really high extinction coefficients are tried on thick appendages, such as an adult human finger. However, significantly reducing the total path length of the light, such as by a factor often, a resolvable signal can be obtained.

[44] A mouse paw is at least one tenth the thickness of a human finger and thus, the arterial vessels within a mouse paw should be about one tenth the size of the arterial vessels of within a human finger. The present invention relating to wavelength considerations has particular application to small mammal research, as well as sensors designed for low birth weight humans. Effectively, a very big, resolvable pulsitile signal can be achieved by using a wavelength with a very big extinction coefficient only if the specimen is very thin, such as the paw of a mouse or rat or finger of an infant or web of a human hand, etc. Today's devices have been designed for adult human fingers, which are much thicker specimens than neo-natal human fingers or the paws of mice and rats.

[45] Experimental results have shown that 405 nm U.V. light, 430 nm blue light, 463 nm blue light, aqua light, green light, and the yellow lights all have wavelengths that produce results better than 3 times that produced by red and infra red light, hi other words, all of these wavelengths appear to be better choices for pulse oximetry. Green and aqua light appear to be the better choices for pulse oximetry and the 463 nm, 430 nm blue, and 405 ntn aqua lights appear to be the best choices for pulse oximetry. When experiments were conducted using the pinky finger of a seven year old child, pulsitile measurements could not be obtained for any of the LED's with wavelengths in the blue and U. V. range. This is most likely because the absorption of blue and U. V. light is so high that, for the amount of light that was put into the child's finger, a distinguishable amount of light could not pass.

[46] Interestingly, when this experiment was conducted using web of the child's hand, measurements using 463 nm and 405 nm could be obtained, but measurements at 430 nm were still not obtainable. This is most likely because the absorption of 430 nm light is at least 3,000 times higher than the absorption of 463 nm and 465 nm light. During this experiment, where the earlobe of a child was used, measurements using 463 nm, 405 nm and even 430 nm light could be obtained. This is most likely because even though the absorption of these wavelengths is so high, a child's earlobe is so thin that enough light could pass to make measurements possible at all three wavelengths. Consequently where the path length can be decreased (from an adult or even child's finger as in a prenatal finger), the decreased wavelengths can achieve superior results. In adults, the path lengths can be reduced by not taking measurements at the patient's finger. The patients ear lobe, nose, web of the hand have all proven to be sufficiently perfuse, e.g., enough blood flow, with a short enough path length to achieved superior signal strength.

[47] In order to address whether order-of-niagnitude or linear differential is more important, the following analysis is provided. The transmitted intensity of light traveling through a medium containing oxy-hemoglobin can be written as

T _ T -ε&MOztø (c\

¹HbO₁ ~ ^I0^κ > \^JJ where I₀ is the incident light intensity, ε(.λ_mθ2) is the extinction coefficient for oxygenated hemoglobin, c is the concentration of oxygenated hemoglobin, and d is the path length through the medium.

The same equation for reduced hemoglobin is expressed as: ¹Hh — -O^e • \°)

[48] In these equations, it is assumed that the light intensity, the concentration, and path length are the same for both conditions. From a design standpoint, it is the ratio of the transmitted intensity that is of concern, which is obtained simply by dividing one equation by the other (the order is arbitrary). Dividing equation (5) by (6), yields

^HbO₂ _ -[B (λ_mθ2 )-ε (λ_m)]cd ,-..

— e . {/)

¹Hb

[49] As shown in equation (7), the attenuation of the light intensity is related directly to a linear difference between the extinction coefficients of the two hemoglobin states. Thus, a greater signal level is obtained at lower wavelengths where the extinction coefficient difference is linearly much greater. The biggest separation occurs at 420 run as noted in FIG. 8. Although some of the wavelengths of light less than 600 nm have a greater separation of extinction coefficients between oxyhemoglobin and reduced hemoglobin than a wavelength of 660 nm, the typical photodiode receiver response is greatly reduced. If the photodiode response is greatly reduced, then the signal to noise ratio of the photoplethysmograph is greatly reduced. This may explain why pulse oximeters have historically employed a wavelength of 660 nm to for measuring oxygen saturation.

[50] Although some wavelengths that are less than 660 nm provide more separation between extinction coefficients, the photodiodes that have been employed historically they do not respond well to them. In order to minimize this photodiode response problem, the common silicon photodiode can be replaced with what's known as a "blue or U. V. enhanced photodiode" for the sensor 26. At all wavelengths from approximately 400 nm to approximately 730 nm, a "blue or U. V. enhanced" photodiode has a much greater response than a common or normal photodiode. At approximately 500 nm the blue enhanced photodiode responds approximately equal to the normal photodiode's response to 660 nm of light.

[51] As noted above, any wavelength less than 620 or 600 nm will have a larger pulse gain than that at the 660 nm and 900 nm which are most commonly used in pulse oximeters. Regarding differential SpO₂ gain, there is a preferential pairing of wavelengths to improve the oximeter results. As shown in FIGS. 7 and 8, the SpO₂ gain at red 660 nm is larger than that at infrared 900 nm, but it is also larger than many values OfSpO₂ gain at smaller wavelengths. Additionally, as was stated above, the SpO₂ gain by itself is not as important as the difference in SpO₂ gain between the two chosen wavelengths. Thus, we must look at the combinations of any two wavelengths. [52] hi order to do this effectively, the differential SpO₂ gain of a wavelength pair may be compared relative to that between red 660 and IR 900. The method involves first calculating the SpO₂ gain of any wavelength, then normalizing it to the SpO₂ gain of red 660. This ratio, R , may be written as:

_R Sp₁GaIn_x S _pO₂Gcιin_red660

[53] Equation (8) can also be evaluated at the wavelength IR 900 as follows:

[54] The differential SpO₂ gain is effectively defined as the difference between the SpO₂ gains of any two wavelengths. Denoting these wavelengths as x and y , the relative SpO₂ gains may be written as R_x and R_y . Expanding R_x above, and can write R_y as follows:

Having written these, the differential SpO₂ gain may be written as:

AR_x>y = abs(R_x -R_y). (11)

[55] Finally, to compare the gains of two selected wavelengths, the present inventors have calculated the difference in R values for the specific pairs as shown in the following chart:

Any pair of wavelengths has an improved differential SpO₂ gain over that of red 660 and IR 900 if KR_{x y} > M₆₆₀₉₀₀.

[56] As noted in the above chart, all the tested pairing exhibited superior gains for the SpO₂ signal than conventional 660, 900 nm sensors except (507, 524), (507, 630), (507, 660), (524, 630), (524, 660) and (630,660). Reiterating the above discussions, any wavelength less than 620 nm (or about 600 nm) will have a larger pulse gain than that at 660 nm and 900 nm, as commonly used in the current art, where path length is not a great concern. Any pair of wavelengths has an improved differential SpO₂ gain over that of red 660 and IR 900 if AR_x > AR₆₆₀₉₀₀ . The best pairs of LED' s for the present pulse oximeter should be the pairs that provide the largest differential SpO₂ gain and also provide the largest pulse gain for each of the two wave length LED's that comprise the pair. The top five choices of oximeter LED pairs are: 1^st choice = 430 nm & 405 nm wavelength LED's; 2^nd choice = 572 nm & 430 nm wavelength LED's; 3^rd choice = 463 nm & 430 nm wavelength LED's; 4^th choice = 524 nm & 430 nm wavelength LED's; and 5^th choice = 507 nm & 430 nm wavelength LED's. Other good pairs, as shown in the table, would be of utility in situations in which tissue attenuation is greater. [57] Those skilled in the are can appreciate that the photoplethysmographic sensor unit of the present invention achieves the following advantages: (a) increases the light transmission; (b) increases the ratio of the AC signal to DC signal, and (c) increases the resolution of the SpO₂ measurement. Various other objectives and advantages of the present invention become apparent to those skilled in the art. [58] Although the invention has been described in detail for the purpose of illustration based on what is currently considered to be the most practical and preferred embodiments, it is to be understood that such detail is solely for that purpose and that the invention is not limited to the disclosed embodiments, but, on the contrary, is intended to cover modificatipns and equivalent arrangements that are within the spirit and scope of the appended claims.

Claims

What is claimed is:

1. A photoplethysmographic sensor unit comprising: a light member support; a light source operatively coupled to the light member support; a first layer of a coupling gel disposed proximate to the light source; a sensor support; a light receiving sensor operatively coupled to the sensor support; and an attachment mechanism adapted to attach the sensor unit to an extremity of a subject such that the first layer of coupling gel is disposed between the light source and such a subject's extremity.

2. The sensor unit of claim 1, further including a removable release layer adjacent the first layer of coupling gel, wherein the release layer is removed prior to attaching the sensor unit to a subject.

3. The sensor unit of claim 1, wherein the attachment mechanism is a clam-shell clamp comprising a biasing member adapted to bias the light member support, the sensor support, or both toward a subject's extremity.

4. The sensor unit of claim 1, wherein the attachment member is a strip configured to wrap around a subject's extremity.

5. The sensor unit of claim 1, wherein the coupling gel has an index of refraction generally between the index of refraction of the light source and the index of refraction of a subject's extremity.

6. The sensor unit of claim 5, wherein the coupling gel has an index of refraction generally equal to the index of refraction of a subject's extremity.

7. The sensor unit of claim 1, further including a second layer of coupling gel disposed proximate to the light receiving sensor.

8. The sensor unit of claim 1, wherein the light source has a wavelength of less than 620 nanometers.

9. A photoplethysmographic sensor unit comprising: a light member support; a light source operatively coupled to the light member support; a deformable pressure reducing relief layer disposed proximate to the light source, wherein the relief layer is translucent to light from the light source; a sensor support; a light receiving sensor operatively coupled to the sensor support; and an attachment mechanism adapted to attach the sensor unit to an extremity of a subject such that the deformable pressure reducing layer is positioned proximate to a subject's extremity substantially eliminating air gaps therebetween.

10. The sensor unit of claim 9, wherein the deformable relief layer is formed as a removable, disposable gel pad attachable to the light source.

11. The sensor unit of claim 10, further including a removable release layer disposed proximate to the gel pad, wherein the release layer is removed prior to attaching the sensor unit to a subject.

12. The sensor unit of claim 9, wherein the attachment mechanism is a clam-shell clamp with a biasing member for biasing the supports toward a subject's extremity, a wrap around strip configured to wrap around a subject's extremity, or a combination thereof.

13. The sensor unit of claim 9, wherein the relief layer has an index of refraction between the index of refraction of the light source and the index of refraction of a subject's extremity.

14. The sensor unit of claim 9, wherein the relief layer is formed as a layer of coupling gel and further including a second layer of coupling gel disposed proximate to the light receiving sensor.

15. The sensor unit of claim 9, wherein the light source has a wave length of less than 620 nanometers.

15. The sensor unit of claim 9, wherein each deformable relief layer is formed as a removable, disposable gel pad attachable to the light source and the gel pad includes opaque side walls to increase light transmission and reduce ambient light transmission.

16. A pulse-oximeter photoplethysmographic sensor unit comprising: a light member support; a light source operatively coupled to the light member support wherein the light source has a wave length of less than 620 nanometers; a deformable pressure reducing relief layer disposed proximate to the light source, wherein the relief layer is translucent to the light from the light source and adapted to be positioned between a subject's extremity and the light source; a sensor support; a light receiving sensor operatively coupled to the sensor support; and an attachment mechanism for attaching the sensor unit to an extremity of the subject.

17. The pulse-oximeter sensor unit of claim 16, wherein the deformable relief layer is formed as a removable, disposable gel pad attachable to the light source and the gel pad includes opaque side walls to increase light transmission and reduce ambient light transmission.

18. The pulse-oximeter sensor unit of claim 16, wherein the attachment mechanism is a clam-shell clamp with a biasing member for biasing the supports toward a subject's extremity, a wrap around strip configured to wrap around a subject's extremity, or a combination thereof.

19. The pulse-oximeter sensor unit of claim 16, wherein the relief layer has an index of refraction between the index of refraction of the light source and the index of refraction of a subject's extremity.

20. The pulse-oximeter sensor unit of claim 16, wherein the relief layer is formed of a layer of coupling gel and further including a second layer of coupling gel disposed proximate to the light receiving sensor.