Methods of minimizing scattering and improving tissue sampling in non-invasive testing and imaging

An improved method and apparatus for use in optical testing of concentration in samples has been developed. The apparatus restricts the solid angle of illumination and the solid angle of detection to eliminate a high proportion of the scattered radiation while allowing the ballistic radiation and the snake-like radiation to be transmitted. In samples which contain multiple scattering centers, this allows less correction for variations in effective pathlength and allows easier calibration of the apparatus. The use of polarized radiation as a means of minimizing scattered radiation in the sample is also disclosed.

BACKGROUND OF THE INVENTION 
The present invention relates to improvements in optical systems and their 
uses for the measurement of concentration and temperature in scattering 
media, and the related discrimination of subsurface features. More 
particularly, the invention provides methods and apparatus which minimize 
the ratio of diffusely scattered radiation to directly transmitted 
radiation reaching the detector(s) in optical concentration measurement 
and imaging apparatus. The methods and apparatus of the invention have 
special applicability to non-invasive testing, particularly for 
concentration measurements of materials such as glucose and hemoglobin in 
blood. 
Recent literature is replete with articles describing attempts at 
performing noninvasive testing using optical measurements (e.g., infrared 
systems). Part of this expansion has been fueled by the spread of acquired 
immunodeficiency disease syndrome (AIDS), and the associated fear among 
public and health care personnel of AIDS. AIDS and other diseases such as 
hepatitis are born in the blood and can be spread by improper practice of 
invasive procedures. In addition, the diabetic population has also been 
anxiously awaiting non-invasive test instruments for many years. Many 
diabetics must test their blood glucose levels four or more time a day. 
The modem battery powered instruments for home use require a finger prick 
to obtain the sample. The extracted blood samples are placed on a 
chemically-treated carder which is inserted into the instrument to obtain 
a glucose reading. This finger prick is painful and can be objectionable 
when required frequently. In addition, although the price has dropped 
considerably on these instruments, the cost for the disposable and the 
discomfort and health risk associated with having open bleeding is 
undesirable. 
Accordingly, a number of groups have recently tried to make non-invasive 
instruments for testing a variety of analytes, particularly glucose. A 
recent trend in non-invasive testing has been to explore the use of the 
near infrared spectral region, primarily 700-1100 nm because this is the 
spectral response range of the silicon detectors typically used in the 
prior art. A wider wavelength range to .about.1800 nm can be accessed by 
the addition of germanium and/or InGaAs detectors, and useful measurements 
can be made into the 2500 nm range with InSb or other detectors. The 
region below .about.1400 nm is the most useful in transmission, as tissue 
is transparent enough there to allow high enough photon flux for accurate 
detection. Above 1400 nm, the strong absorption of water limits the 
penetration depth of tissue, so that useful measurements are typically 
made in reflectance geometry. Below 1100 nn, the penetration of the light 
is sufficient that the signal modulation during the arterial pulse can be 
measured comfortably in both transmission and reflectance geometries. 
Above 1400 nm, such pulsatile measurements are extremely difficult in 
transmission due to low intensity, and similarly difficult in reflectance 
because the light does not penetrate deeply enough to sample the pulsatile 
capillary beds. 
Most of the non-invasive testing work has been carried out using classic 
spectrophotometric methods, such as a set of narrow wavelengths sources, 
or scanning spectrophotometers which scan wavelength by wavelength across 
a broad spectrum. The data obtained from these methods are spectra which 
then require substantial data processing to eliminate background; 
accordingly, the papers are replete with data analysis techniques utilized 
to glean the pertinent information. Examples of this type of testing 
includes the work by Clarke, see U.S. Pat. No. 5,054,487; and primarily 
the work by Rosenthal et al., see e.g., U.S. Pat. No. 5,028,787. Although 
the Clarke work uses reflectance spectra and the Rosenthal work uses 
primarily transmission spectra, both rely on obtaining near infrared 
spectrophotometric data. 
The major successful application of noninvasive testing is the measurement 
of hemoglobin oxygen saturation with pulse oximetry. The most common 
method compares the percentage modulation of the intensity of light 
traversing a body part at two wavelengths chosen so that the ratio of 
their respective modulations is a relatively strong function of oxygen 
saturation. The observed change in this ratio is relatively large because 
the two hemoglobin species involved have both high enough concentrations 
and specific absorptions that they dominate the creation of the pulsatile 
signal components at the wavelengths of interest. As a result, the ratio 
of modulations can be attributed substantially to the two hemoglobins 
alone, and only needs to be measured to the order of 0.1% in order to 
achieve clinically significant detection limits with acceptable 
universality of calibration. 
The optical system in typical pulse oximeters have two or more LED emitters 
placed side-by-side on one side of a finger, and a single detector 
receiving the radiation on the other side of the finger. Some more recent 
systems have the detector on the same side of the tissue as the emitters, 
with baffles preventing the direct illumination of the detector by the 
sources. As the sources are physically small and optically displaced from 
each other and the detector, the light from each detector enters the 
tissue at slightly different locations, and therefore travel different 
paths through tissue to the detector. 
Despite its relatively large signal levels, pulse oximetry has well-known 
difficulties such as the selection of an adequately vascular sampling site 
on each individual and variability of the results with motion of the site 
and breathing by the patient, as well as sensitivity to changes in blood 
pressure, heart rate, temperature, and tissue hydration. Disturbances such 
as motion and breathing artifacts typically appear as statistical 
outriders, i.e., as measurements which fall well off the "average" 
calibration curve of the instrument obtained from a group of individuals 
breathing controlled gas mixtures to vary their oxygen saturation. 
The calibration of a pulse oximeter is subject to these same error sources; 
it is not uncommon to find site-to-site variations on the same individual, 
with results that suggest that the calibration curve even varies, for 
example, with the absolute magnitude of the pulsatile signal modulation. 
The effort to obtain a meaningful universal calibration is clearly at odds 
with intra- and inter-individual physiological variations. 
Despite recent efforts to improve the measurement S/N by increasing source 
intensities and lowering detector noise, as well as increasing the number 
of detectors, the frequency of outriders and the universality of 
calibration have not improved substantially. Thus it is clear that while 
the light traversing the tissue is being measured more precisely, the 
site- and physiologically-induced variability has not been improved 
significantly below the 0.1% level needed for the measurement of oxygen 
saturation. 
While these physical and physiological interferences are marginally 
acceptable for oxygen saturation measurements, they set a lower limit of 
detectivity that is too high for other clinical analytes such as glucose 
and cholesterol for which the combination of concentration and specific 
absorption requires optical measurements to be made 100-1000 times more 
precise than for the hemoglobins used in pulse oximetry. The hemoglobins, 
which in themselves are difficult to calibrate in the presence of these 
site- and physiologically specific limitations comprise a major background 
interference for the measurement of such trace constituents as glucose. 
The optical systems employed for these lower concentration analytes 
naturally drew on the experience of pulse oximetry, and typically employ 
similar arrangements of a plurality of slightly displaced LED's to extend 
the wavelengths sampled, or which use fiber optics to carry light to and 
from the sources and/or spectrometers which perform the separation of the 
signal into the different wavelengths employed. Displacement of the 
sources and wide numerical apertures for the light entering and leaving 
the tissue enhance the likelihood that different detected wavelengths will 
have sampled different portions of the medium. Many of the physiological 
interferences to accurate measurement are mediated by differences in the 
mean paths traced by light of different wavelength in traversing the 
intervening tissue between light sources! and detectors!. These path 
variations are produced, in part, by light scattering in the tissue, which 
varies with the wavelength of the light and which makes photons follow a 
jagged overall path from scattering to scattering. The detected signals 
are a complicated function of both the scattering and the total absorption 
of all constituents along the longer total path of the light. Thus, the 
present optical systems used for noninvasive measurement allow and perhaps 
even encourage light of different wavelengths to travel different paths 
through the tissue, sampling lateral and axial tissue inhomogeneities 
differently. 
This situation violates a fundamental premise of all optical non-invasive 
measurement methods; namely, that the light intensity which is measured in 
the individual detection channels can be attributed to the analyte and not 
to any difference in tissue sampling. Tissue inhomogeneity produces 
wavelength-dependent spreading of the light which ultimately reaches the 
detectors, and in the extreme of high scattering and large inhomogeneity, 
the mixture of detector signals becomes an uncontrollable and 
uncalibratable average response to the physiological and biochemical 
conditions at the sampled site. 
In addition the existing noninvasive art has employed spectrophotometric 
methods which limit the intensity of light detected in the individual 
resolution elements, and which also apply the method in a way which uses 
the available spectral information inefficiently. These methods were 
conceived primarily for accurate determination of narrow band spectral 
structures rather than for discriminating the presence of weak broadband 
features in strong broadband backgrounds that characterize the noninvasive 
measurement problem for constituents such as glucose. The multivariate 
analysis mathematics required to separate the analyte signature from 
strongly overlapping interferent signatures also introduces an error 
propagation penalty that compounds the intensity limitation by increasing 
the impact of detector noise on the calculated measure of concentration. 
Improvements that enhance the solution of problems of interference in 
broadband spectra, by obtaining different raw data, are described in U.S. 
Pat. No. 5,321,265 (the "Block 265 Patent"). This patent sets forth a 
different approach in non-invasive testing as compared with the prior 
instruments and methods. As noted, substantially all workers in the 
non-invasive testing field prior to the Block '265 Patent were using 
classic spectrophotometric instrumentation and substantial processing in 
an attempt to resolve the low resolution features from the background. 
However, the spectra of analytes such as glucose in a human body are not 
discrete high resolution features which spectrophotometric instruments 
were originally designed to measure but rather have a few low resolutions 
features with much of the information contained in subtle variations of 
the detected intensity as a function of wavelength. As such, these spectra 
appear more like reflectance spectra of colored objects in the visible 
region. The Block '265 Patent teaches the use of an analog of human color 
perception to obtain meaningful data by means of methods and apparatus 
which utilize overlapping, broadbeam detectors to mimic the spectral 
response characteristics of the human retinal cones, but translated into 
the near-infrared. This approach, which is radically different than 
classic spectrophotometric measurements, provides advantageous effects in 
determining the concentration of glucose and other similar materials in an 
aqueous solution and is particularly advantageous for use with scattering 
media such as tissue where it also provide the added advantage of higher 
light flux at the detectors so that the intrinsic shotnoise limitation as 
a percentage of the total signal intensities is reduced. 
U.S. patent applications Ser. Nos. 08/130,257, 08/182,572 and 08/333,758, 
the disclosures of which are incorporated herein by reference, all 
disclose improvements in the basic techniques and apparatus described in 
the Block '265 Patent. These improvements include the concepts of 
congruent illumination and detection of light emerging from the sampled 
tissue site, pulsatile processing, modulation of illumination sources as a 
means of eliminating unwanted radiation, the use of non-overlapping broad 
beam radiation as well as overlapping radiation, and a number of 
variations thereon. These applications make it clear, in part, that a 
variety of techniques are useful (and in some instances may be necessary) 
to deal with the problems encountered in non-invasive measurement of 
analyte concentration in tissue or other scattering media. Many of these 
problems arise from the fact that scattering media exhibit higher 
effective path lengths than their physical dimensions because of 
scattering by the samples themselves. In fact, the samples, such as human 
tissue, act as if they are formed of a plurality of scattering sites or 
centers in the sample. Techniques such as the congruent illumination and 
congruent detection described in these patent applications equalize the 
acceptance angles and distances traveled by light of different wavelengths 
outside the scattering media. Technically, this is achieved by locating 
all the illumination sources and/or detectors so that the path lengths and 
angles between the media and the detector(s)/source(s) are equal, so that 
the detectors or radiation sources act as if they were optically 
superimposed. 
However, the desired congruency of detected light is degraded within the 
observed media because the multiple scatterings of light spread the light 
beam to adjacent regions in a way which is strongly wavelength dependent. 
If the scattering media is inhomogeneous, the result of this spreading is 
to mix light from these adjacent structures in relative amounts which are 
dependent on wavelength. One object of the present invention is to reduce 
this disturbing effect by refining the launch and detection optics to 
limit their angular acceptance ranges. 
It has long been known that a certain portion of the illuminating radiation 
survives transit across a turbid sample without being either scattered or 
absorbed, while a much larger portion is scattered in all directions. The 
more scattering a particular photon undergoes the longer the integrated 
path it follows, and the longer the time that elapses before it emerges 
from of the sample. Some groups have attempted to reduce the deleterious 
effects of scattering by using pulsed sources and time gating the 
detection so as to view the sample only in light which has undergone few 
scatterings. What is measured is a "snapshot" of the sample in light 
starting at the time of flight for an unscattered beam, and extending long 
enough in time to obtain sufficient signal for the desired analysis 
without including much scattered radiation. When the time gate is short, 
"ballistic" or "snakelike" photons which have undergone no or few 
scatterings along their path are selected, and shadowgram images similar 
to those commonly obtained with x-ray's can be obtained. 
This approach, however, requires complicated apparatus, and in addition to 
the intensity limitation from the short time-gate after each pulse of the 
light source, adds a further limitation on the number of detected photons 
because the duty-cycle of the pulsed source is low compared to the 
continuous source of the present invention. Other research groups such as 
Wist et al., IEEE Transactions on Medical Imaging, 12 (4) 751-757 (1993), 
have demonstrated that shadowgram-type images can be obtained by severely 
restricting the angular acceptance range of detected photons about the 
forward direction, essentially demonstrating that doing so limits the 
detection to "ballistic" or "snakelike" photons. The Wist et al. apparatus 
generates a geometrically narrow beam which is raster-scanned across the 
sample, at a first wavelength, and then generates new images at changed 
wavelengths. The work of this group, however, also demonstrates a severe 
limitation on the total flux of transmitted photons which make it 
inapplicable to the detection of trace constituents in scattering media. 
Other workers such as Schmitt et. al., SPIE 1641, 150-161, (1992), have 
demonstrated advantages for using collimated input and output light on in 
vitro phantoms that simulate some of the light scattering properties of 
turbid media, but the transmitted intensity limitation of their system 
when it was applied to real in vivo measurements made it necessary to 
change the system design away from this collimated approach. One 
difficulty appears to be that their in vitro system was designed to 
"approximate the plane-parallel conditions under which the theoretical! 
photon diffusion model was derived," rather than addressing the 
characteristics of the in vivo sample. Schmitt's collimated system was 
designed to approximate a "collimated beam of infinite extent" by 
establishing a finite incident beam of light traversing tissues and 
confining the collimated detection to a small central region on the exit 
side, apparently in order to eliminate unwanted edge effects. In addition, 
the narrowband sources and detector used limited the transmitted 
intensity. 
The failure of Schmitt's design was that insufficient photon flux was 
available at the detector, so that this system was abandoned for his in 
vivo work. Instead, Schmitt's in vivo apparatus employed a fiber optic 
that launched light into the tissue at its large (.about.50 degree) 
numerical aperture, and an integrating detector on the opposite side of 
the tissue receiving light through almost the whole hemisphere. Even then, 
as noted in his article, the system had inadequate light intensity for the 
measurement he was attempting. His work thus vividly illustrates the light 
transmission limitations of real tissue that characterizes the prior art. 
Thus, it is a specific object of the present invention to balance the light 
collection efficiency and spatial resolution of the optical sampling 
system viewing scattering media to simultaneously achieve high detected 
light intensity and equality of response, as a function of wavelength, to 
inhomogeneous inclusions within the media. This is accomplished by 
selecting optical configurations of sources, detectors, and intervening 
optical elements to minimize the effect of tissue inhomogeneities on the 
relative changes in signal strengths in each of the different detectors 
due to the presence of analyte. 
It is a further object of the invention to achieve this balance in a way 
which improves the repeatability of the measurements from site-to-site on 
a given individual in the presence of disturbances such as motion, 
breathing, hydration, and the like, with the Ultimate objective to achieve 
universal calibratability of the measurement across subject in the 
presence of such disturbances. 
A related object of the invention is to provide a method of non-invasive 
concentration measurement in a scattering media which increases the ratio 
of direct collimated radiation to diffusely scattered radiation reaching 
the detector, while maintained high integrated light intensity at the 
detectors. 
Another object in the invention is provide an apparatus for non-invasive 
concentration measurements which maximizes the ratio of direct collimated 
radiation to diffusely scattered radiation while maintaining high 
integrated light intensity. 
A further object of the invention is to facilitate the use of tighter 
collimation by increasing the number of photons received in the individual 
detector resolution elements through broadening their wavelength 
acceptance range. 
A similar objective of the invention is to facilitate the use of tighter 
collimation by increasing the number of photons received by individual 
detector resolution elements through increasing their surface area while 
maintaining their congruency. 
Yet another object of the invention is to further facilitate the use of 
tighter collimation by the use of overlapping broadband detector 
resolution elements in an analog of human color perception to combine 
increased photon flux with more efficient separation of similar broad 
analyte and interferent spectral features. 
Consequently, it is a specific object of this invention to select optical 
configurations of sources, detectors, and intervening optical elements to 
minimize the effect of tissue inhomogeneities on the relative changes in 
signal strengths in each of the different detectors due to the presence of 
analyte. 
It is a still further object of this invention to adjust the optical 
interface to take maximum advantage of the natural spreading 
characteristics of the light distribution patterns in tissue in maximizing 
the S/N of the determination. 
These and other objects which features the invention will be apparent from 
the detailed description and the drawing. 
SUMMARY OF THE INVENTION 
The present invention features methods and apparatus for measuring 
concentration in a sample which contains a plurality of radiation 
scattering sites, and for measuring the distribution of concentration 
and/or temperature within a sample when employed with imaging detectors. 
The methods and apparatus can also be utilized for discrimination of 
subsurface features through shadowgram generation. This procedure is also 
useful for detection of temperature inhomogeneities. The methods and 
apparatus of the invention employ means for restricting the solid angle of 
illumination and/or collection, e.g., by collimation of the radiation to 
minimize the amount of scattered radiation collected, employ polarization 
of the illuminating radiation to differentiate scattered from unscattered 
radiation, or a combination of the two. The methods and apparatus of the 
invention provide more reproducible measurements on scattering media and 
are particularly well suited to non-invasive testing of tissue for 
materials such as glucose and drugs of abuse. 
More particularly, the present invention provides a method of measuring the 
concentration in a sample of a selected substance which absorbs radiation 
in a particular region of the spectrum. The sample containing the 
substance of interest also contains a plurality of sites which scatter 
radiation in the same particular region of the spectrum. The method has 
the steps of illuminating the sample with broad geometric area 
illumination within a particular region of the spectrum (preferably using 
broad spectrum radiation) where the substance of interest has absorption, 
with the illumination and detection solid angles restricted, and with both 
said solid angles extending over a geometrically wide surface 
cross-section. The term "broad spectrum illumination" as used herein means 
and implies that the wavelength of the illumination covers a substantial 
portion of the region of the spectrum in which there is absorption by the 
selected substance. Normally, the illumination is greater than 50 nm wide, 
and if the substance of interest has absorbance at several wavelengths in 
the particular region, it preferably is wide enough to cover all of the 
absorption bands. 
After leaving the sample, radiation which is transmitted or reflected from 
the sample is collected with a detector, the detector being selected and 
located such that the each resolution element of the detector collects 
radiation only from a limited solid angle extending over a relatively wide 
area of the viewed surface. The restricted solid angle illumination also 
extends over a relatively wide area of the illuminated surface. The term 
"wide" as used herein implies a beam width that is comparable or larger in 
size than the thickness of the tissue being viewed. The term "comparable" 
means and implies that the width of the beam or viewing area is at least 
half the thickness of the sample or tissue. That thickness is itself 
preferably restricted to be not much deeper than several "natural" 1/e 
penetration lengths, the depth over which the diffuse radiation photon 
density falls to a few percent of its maximum value near the entrance 
surface of the medium. For typical tissues viewed in the 700-1400 nm 
wavelength range, these preferred thicknesses are of the order of a few 
mm, and the surface areas through which the light enters and exits are 
both in the 5-10 mm range. 
The terms "restricted solid angle" and "limited solid angle" as used herein 
imply that the type of detector or illuminating radiation, which may 
include some form of filtering and/beam focusing apparatus, limits the 
angle over which the illumination or detection occurs. Preferably, the 
illumination and or detection is restricted to a solid angle of about 
10.degree. or less from the central illumination beam axis. This 
configuration maximizes the ratio of directly transmitted radiation to 
scattered radiation collected by the detector from the sample. Preferably, 
the restricted solid angle of illumination is achieved by collimating the 
radiation from the radiation source prior to illuminating the sample, most 
preferably with collimating optics such as described herein. 
Alternatively, a laser or another source which provides restricted 
divergence illumination may be used without the necessity of some type of 
collimating optics. 
Similarly, the preferred detector limits the solid angle of the detected 
radiation transmitted from said sample by excluding uncollimated radiation 
prior to collection by the detector. Again, collimating optics, such as a 
combination of lenses and/or apertures, can be used. Alternative 
collimating optics known in the art, such as a channel plate or a 
honeycomb collimator, can be used as well. Some of the benefits of the 
method may be achieved by comparing the results for two or more narrow 
angular acceptance ranges at different angles relative to the central axis 
of the illumination beam. In fact, using a second, off axis detector can 
assist in identifying the contribution of diffuse radiation and assist in 
correction for motion and other artifacts. In another embodiment, it is 
advantageous to have congruent detectors viewing the scattering media on 
the same side as the illumination, in what is commonly referred to as a 
reflectance or transflectance measurement geometry. Certain tissue sites, 
such as the forehead, may be particularly advantageous in this geometry 
because the vascular tissue between the skin and the bone has a thickness 
comparable to the "natural" penetration depths, and the bone serves as a 
relatively inert backing that isolates the overlying tissue from other 
analyte containing deeper tissue. 
In addition to the strictly collimated beam, a combination of selected 
slightly converging or diverging beams generated by a combination of 
lenses and/or apertures can be used in the present invention. In one 
aspect of the invention, a first converging lens is selected and located 
on the illumination side of the sample such that its focal point for the 
illuminating radiation is located on the detection side of the sample and 
the first converging lens. In this embodiment, limiting the solid angle 
viewed by the detectors may be achieved by means of a second converging 
lens between the sample and the detector which has a limiting aperture 
mask at its focal distance so that the size of a central hole in that mask 
defines the angular acceptance range. This same type of second converging 
lens is used on the illumination side. After passing through this 
aperture, the light beam is then expanded once again and passed through 
the congruent beamsplitter arrangement described in U.S. patent 
applications Ser. Nos. 08/130,257 and 08/182,572 and in the following 
examples. If this configuration is used to collect light in other than the 
forward direction, the entire lens/aperture/detector assembly would be 
rotated about an axis centered preferably beneath the surface of the 
viewed sample. 
The optics of the system and methods described herein are such that they 
eliminate much of the scattered radiation reaching the detector, 
specifically all the scattered radiation outside the limited solid angle 
viewed by the detector, while allowing the use of geometrically wide 
radiation beams. This enhances the ratio of directly transmitted to 
diffuse radiation which reaches the detector. This is particularly 
advantageous with a heavily scattering media, e.g., the sample is a 
portion of an mammalian body, such as a human body. If a human body is 
used as a sample, preferably a thin region of tissue is selected so as to 
minimize person-to-person variation in tissue thickness. Once again the 
preferred thickness would lie in the range of a few (2-10) mm, comparable 
to the distances over which the internal light density falls to a few 
percent of its maximum value. Thinner tissues can be used provided that 
they are sufficiently vascular to provide good pulsatile signals. One 
possible area for use is a finger web which can be clamped to provide a 
substantially standardized tissue bed thickness. Another alternative is 
the eyelid. Thinner tissues such as these can also be backed by reflective 
surfaces to achieve a "double-pass" effect, or with an absorbing backing 
that also isolates the tissue from underlying structures. 
The geometrical width of the illumination and detection areas on both sides 
of the sample are designed to be wide enough to average over multiple 
internal structures such as capillary beds. This reduces sensitivity of 
the results to the exact positioning of the optical system at the chosen 
site. This approach achieves significantly higher integrated light 
intensity at the detectors, while avoiding edge effects at the extremes of 
the illumination beam. Thus, the large geometric beam area facilitates 
passage of a large number of photons, and the solid angle restrictions 
limit differences in the lateral spreading of the light within the tissue 
at different wavelengths. This geometry produces relatively higher 
detected light intensity with more consistent sampling across tissue 
inhomogeneities as a function of wavelength than does previously described 
apparatus and methods. The intensity is high enough that the fundamental 
shot-noise detection limit falls below the precision needed for the trace 
analytes of interest, and the improved sampling reduces the 
non-linearities imposed on the measurement calibration by scattering. 
The preferred working range for the method and apparatus invention is from 
about 700 to about 2500 nm. This region covers the absorbance of some of 
the most preferred substances, including glucose and its identifying 
substances, hemoglobin, deoxyhemoglobin and various drugs of abuse. The 
method can be used to determine the hemocrit or to derive the oxygen 
saturation level in the blood. This method and apparatus can be made 
sufficiently rapid to measure arterial pulse data, thereby eliminating 
another source of potential error. The present improvements can also be 
utilized with modulated sources, which are particularly helpful in 
eliminating radiation generated from sources other than what is 
transmitted by the sample. 
In an other embodiment of the invention, polarized light is used as the 
illuminating radiation. The detector which collects the radiation includes 
an analyzer or filter for transmitting polarized light before, or as part 
of, the detector, while excluding depolarized light transmitted from the 
sample. Since scattering of polarized radiation by the scattering sites in 
the sample will depolarize the scattered radiation, the use of the 
polarizer in conjunction with the detector will maximize the ratio of 
polarized radiation to depolarized radiation collected by the detector. 
The preferred polarization system has the restricted solid angles of 
illumination and limited solid angle of detection previously described. 
The reason for this is, in part, that using this technique not only 
eliminates scattered radiation but also eliminate radiation from other 
non-polarized background light sources and provides the highest ratio of 
desired to undesired radiation while providing sufficient signal. 
If the restricted solid angle geometry is utilized for imaging, an array of 
detector units forming the detector, e.g., a CCD array, can be used. This 
array, and/or illumination source, can be scanned across the sample using 
any standard mechanical stage or raster scanner, to generate a series of 
shadowgrams which can be combined to form a larger shadowgram showing 
tissue inhomogeneities. Temperature changes across the tissue can also 
show up as an inhomogeneity. 
As noted, the present invention also provides a device or apparatus for 
carrying out the method of the invention. All the aspects previously 
described with respect to the methods can be utilized in the apparatus. 
The invention is further illustrated by the detailed description of the 
invention and the drawing.

DETAILED DESCRIPTION OF THE INVENTION 
The present invention is directed, in part, to methods of improving 
consistency of measurement and, therefore, calibration of samples which 
have scattering properties. The methods and apparatus of the invention are 
particularly well suited to non-invasive testing of human tissue. The 
invention is based, to a large extent, on the recognition that if 
sufficient signal is generated in a non-invasive test system, limiting the 
solid angle such that much of the scattered radiation would be outside the 
solid angle of detection would minimize changes in the effective 
pathlength caused by unequal scattering in the sample. 
One of the advantages described in the Block '265 Patent, as well as the 
other abovecited United States patent applications, is that the use of 
broad spectrum illumination and detection provides greater signal 
intensity than in a conventional spectrophotometric systems. High detected 
intensity is vital for the detection of the specific absorption of trace 
analytes, because the changes they make in the detected signals are very 
small. The signal changes produced by clinical concentrations of glucose, 
for example, as so small that they are difficult to separate in aqueous 
solution from changes due to the displacement of water and slight 
alterations in solution temperature. As a result, there are no reliable 
publications of the specific absorption spectrum of glucose in solution in 
the near infrared. Nevertheless, the scale of glucose absorption can be 
approximated on the assumption that the hydroxyl groups of glucose create 
an average band whose size and shape are similar to the corresponding band 
of pure water, but whose location is shifted by interactions with adjacent 
groups on the molecule to have a peak slightly above 1010 nm. At 5.5 
mmol/L (100 mg/dL), these glucosebound hydroxyl groups are about 2000 
times less concentrated than those of bulk water. By scaling from the 0.2 
OD absorption band of pure water at 960 nm, the shifted band-peak from 
glucose at 5.5 mmol/L concentration would have a magnitude of about 
10.sup.-4 OD at 1 cm path, which would create a transmission change of 
2.3.times.10.sup.-4. This result agrees with the general level observed 
for this shifted OH band in higher concentration measurements reported by 
Koashi in U.S. Pat. No. 4,883,953 for the chemically similar saccharose. 
To be clinically useful, this glucose concentration should be measured with 
a precision roughly 20 times smaller, which translates to a measurement 
resolution or precision at the one part in 100,000 level in the case when 
the tissue viewed has an effective pathlength of 1 cm. If pulsatile 
measurement is employed, the time-varying components of the signals, which 
typically comprise 2-10 percent of the total signal, must also be measured 
to sufficient precision. Here, however, the effective pathlength 
generating the observed pulsations is much shorter than 1 cm. For example, 
in the 950 nm range, the dominant absorber in the blood is oxyhemoglobin, 
with total absorbance of about 3 OD for a midrange hemoglobin 
concentration. Such an absorbing solution creates a 10 percent change in 
transmission at a pathlength of only 1/60th cm. As a result, the pulsatile 
measurement of specific glucose absorption requires precision better than 
one part-per-million. 
This is significantly below the repeatability achievable with the present 
state-of-the-art in pulse oximetry The photodetector shot noise limit 
alone requires that there be more than 10.sup.12 photoelectrons detected 
in the integration time of the measurement to reach the part-per-million 
resolution range. Allowing for the fact that only a portion of each 
arterial pulse is useful, the shot noise limit translates to a requirement 
for intensity levels at the photodetectors which lie in the 10-100 
microwatt range, after passage through the tissue. With thick tissues 
and/or inefficient use of intensity, this requirement in turn can lead to 
a need for 10-100 milliwatts, or even watts, of radiation launched into 
the tissue. The intensity which can be used has a practical upper limit 
set, at least, by the levels which become uncomfortable for the test 
subject due to heating by the absorbed radiation. 
Sensitivity at the part-per-million range also enables the measurement of 
the temperature of subsurface features with milli-degree.degree. C. 
sensitivity, or lower. This follows from the fact that the 960 nm water 
band changes its distribution with wavelength when the temperature varies, 
thereby allowing the temperature to be visualized by imaging detectors 
just as if it were another absorbing constituent. Our measurements with 
four 70 nm wide overlapping filters positioned near the 960 nm absorption 
band of water have shown that a 1.degree. C. temperature deviation 
produces signal deviations that are on the same approximate scale as those 
due to the intrinsic absorption bands of glucose at a concentration of 
about 110 mmol/L (2000 mg/dL), but with a different distribution between 
the four detector signals. When the signals from imaging detectors are 
analyzed to display the mean subsurface temperature along the collimated 
path through the tissue, the resultant "shadowgram" images allow the 
localization, for example, of regions of high metabolism, such as tumors. 
It is known that light transmitted through tissue or other scattering media 
can pass through in a variety of ways. FIG. 1 shows a sampling of these 
modes of transmission, in the form of a series of ray traces. More 
particularly, ray trace A on FIG. 1 shows a "snake like" ray which only 
has minor, glancing scattering collisions through the tissue. This type of 
scattering comprises a forward-scattered ray. Ray trace B shows a 
"ballistic" ray which travels directly through the tissue without being 
scattered or absorbed; that is, it passes directly through the tissue, 
being transmitted forward as if there were no scattering centers in the 
sample. Differentiation between "snake-like" and "ballistic" rays is 
difficult, and those who attempted time gating normally include both types 
of rays in their collected data. Ray trace C on FIG. 1 is a multiply 
scattered ray that emerges within the selected angular distribution range 
accepted by the detector. This type of ray will be deemed a forward 
scattered ray by the apparatus and it will be measured or included as part 
of the presumed "unscattered" radiation. Ray trace D is a multiply 
scattered ray which scatters outside the selected angular detection region 
viewed by the detector. For convenience, this multiply scattered ray is 
shown as being only slightly outside the solid angle viewed by the 
detector but actually multiply scattered rays outside the solid angle 
viewed by the detector would constitute the majority of the radiation 
hitting the sample. 
In the visible and near infrared, the relative abundance of these various 
ray types is a very strong function of the thickness of the tissue, the 
magnitudes of the scattering and absorption cross sections for the 
selected tissue sites, and on the so-called single scattering phase 
function (sometimes called the angular scattering distribution). The cross 
sections and phase functions vary strongly between tissue types, and 
within given tissues depend strongly on the state of hydration and 
patient-dependent mixture of physiological sub-structures. 
There has been much attention paid recently in the literature on methods to 
measure, interpret, and extrapolate from these fundamental measurements to 
obtain single scattering parameters and bulk absorption scattering 
coefficients for the media. The theoretical problem, however, is complex 
and does not lead to closed-form solutions unless severely restrictive 
assumptions are made. The situation is further complicated by the fact 
that the coefficients themselves are strongly tied to the underlying 
theoretical model, so that much care must be taken even to compare 
coefficients obtained by different workers to be sure that they are 
defined and measured in the same way. In general, the theoretical models 
have not yet been able to deal well with nonhomogeneous tissue, and 
provide only general guidance on the scale and functional dependencies to 
be expected in real measurements. 
As a result of the complexity of theoretical models, the Monte Carlo 
calculational method is often employed. Here, the paths of many photons 
such as those indicated in FIG. 1 are followed through their transit 
across the scattering medium. The frequency of interaction and the angle 
of each scattering are selected randomly by the computer so as to be 
consistent, on the average of all interactions, with the assumed 
fundamental cross sections and phase functions which form the input data 
for the calculation. While this method readily accommodates nonhomogeneous 
media, it is very ponderous for the solution of the inverse scattering 
problem in which the cross sections and phase functions are determined 
from the observed experimental measurements. Here again, the results 
provide only general guidance as to how the measurements will vary with 
experimentally controllable parameters. 
There is agreement in the literature, however, that in the limit of small 
angular acceptance for a parallel initial beam incident on a semi-infinite 
plane slab of tissue or other scattering media, the transmission of light 
in the forward direction follows the classic form of Beer's law: 
EQU T=e.sup.- (.mu.ss.sup.+.mu. sa.sup.)X 1! 
where .mu..sub.s designates the coefficient of scattering, .mu..sub.a is 
the total absorption coefficient made up of the sum .mu..sub.a 
=.mu..sub.a1 +.mu..sup.a2 +.mu..sup.a3 + . . . of the absorption 
coefficients of the individual constituents present in the scattering 
media, and x is the thickness of the slab. While .mu..sub.a is relatively 
independent of the experimental setup, .mu..sub.s depends strongly on how 
small a solid angle the detector subtends in the forward direction (i.e., 
the smaller the angle, the easier it is for a scattering event to remove a 
photon from the detected beam, and the larger .mu..sub.s becomes). When 
the detector solid angle and/or the thickness are small enough that 
equation 1! is followed, scattering affects the transmission with the 
identical functional dependence as any of the absorbing constituents. 
Further, the contributions of layered media can be readily calculated by 
dividing x into a sum over the individual layer thicknesses. Note that 
with this exponential dependence the effect of a strongly absorbing layer 
on the transmission is independent of the layer depth within the composite 
medium; such a layer produces the same percentage change in the 
transmission wherever it lies along the optical path. 
There are many different theoretical models in use to describe the 
transport of light across scattering media, most of which are based on the 
available closed form solutions to a diffusion equation. To couple these 
available solutions to the real scattering problems, these models 
typically assume a point source of light, completely diffuse illumination, 
or a rapid transition from a parallel input beam of light to a highly 
diffused beam. Once such diffuse illumination is established, it typically 
propagates through the tissue with a relatively stable angular 
distribution that rapidly approaches an isotropic one whose magnitude 
decays with distance following a functionalform similar to: 
##EQU1## 
Here the prime (') indicates that the coefficients in equations 2! are 
not the same as in equation 1!. The magnitudes of the coefficients are in 
fact quite model dependent, while the functional form is generally model 
invariant. 
Equation 2! is only applicable for one dimensional models in which the 
angular distribution of the light reaching the detector is ignored. 
Theoretical models with three dimensional geometries produce equations 
similar to equation 2!, with T replaced by the photon density, x replaced 
by the radius r from the coordinate origin, and with additional factors of 
1/r.sup.2 present to account for the loss in absolute photon density as 
the approximately isotropic beam spreads radially. 
A similar form also develops in the well-known Kubelka-Munk scattering 
theory, with the characteristic exponential decay constant 3 .mu..sub.a 
'(.mu..sub.s '+.mu..sub.a ') of equation 2! replaced in more complicated 
combinations of exponentials by K(2S+K), where K and S represent bulk 
absorption and scattering coefficients of the media defined specifically 
under the constraint of perfectly diffused incident and transmitted (or 
reflected) light distributions. The slight difference in the functional 
form of these exponential terms results from the details of the 
differential equation whose closed form solutions are used to approximate 
reality. Thus, in the Kubelka-Munk formulation a one dimensional 
differential equation is assumed with cross-coupling constants K and S 
between two isotropically distributed light beams progressing in the 
forward and backward directions. All the three dimensional information is 
carded in the assumption of well-maintained-isotropy. Despite their 
differential genesis, it is clear that suitable renormalization can carry 
the K-S form into the .mu..sub.s '-.mu..sub.s ' form. 
Over most of the 700-1400 nm range, .mu..sub.a ' is much smaller than gs 
(see Wilson et al., IEEE J Quant Elec 1990; 26,2186-99 for a review), and 
the second term under the square root in equation 2! is often dropped in 
the theoretical analyses. Above 1400 um, .mu..sub.a ' undergoes a tenfold 
increase because of increased absorption by water. The exponent of the 
transmission in equation 1! is often referred to as the "transport mean 
free path (mfp)," with .mu..sub.s also including a correction for the 
average cosine of the scattering phase function. For typical tissues, the 
mfp usually lies in the range of 0.05-0.2 mm. Similarly the exponent of 
the transmission in equation 2! is called the penetration depth, i.e., 
the distance at which the diffuse radiation levels fall to 1/e. For 
typical combinations of coefficients and phase functions, this is 10 to 20 
times larger than the mfp, and ranges up to a few mm. 
Monte Carlo models, such as those of Flock et al., IEEE Trans Biomed Eng 
1989; 36, 1162-8, provide some guidance on the relative size of collimated 
and diffuse radiation beams. FIG. 6 in the Flock et al. paper is a polar 
plot of intensities for "ten bins of equal solid angle." While the most 
forward bin shown in this figure appears to have a smaller solid angle 
than the others, the figure does indicate that the collimated and 
diffusely transmitted photons will have roughly equal intensity at depths 
of about 20 transport mean free paths, comparable to the natural 
penetration depth of the diffuse radiation. The distance for such 
equivalence is clearly a strong function of the angles and fundamental 
coefficients and those of the site. 
The important result is that the scattering and absorption effects are 
mixed together, so that neither acts alone; an increase in one of them 
with wavelength or sampling site enhances the impact of the other on the 
transmission. When scattering dominates over absorption the dependence of 
the exponent on .mu..sub.a in equation 2! changes from a linear one 
similar to equation 1! to a square root dependence, and the usual 
logarithmic transformation of Beer's law that linearizes equation 1! can 
no longer linearize equation 2!. It is worth noting that equation 2! 
does not converge smoothly to equation 1! as the scattering coefficient 
goes to zero because of the extra factor of 3 in the equation 2! 
exponent, again highlighting the model dependent normalization 
differences. 
The difficulties inherent in the calibration of diffuse transmission are 
illustrated by the work of Cope et al., SPIE 1431,251-62 (1991), who 
coalesce the square root and 3.mu..sub.s ' factors in the exponent of 
equation 2! into a new variable they call the differential pathlength 
factor ("DPF").The DPF is a locally linear approximation to the slope of 
the square root dependence after logarithmic transformation which must be 
calibrated separately at each wavelength of interest with the implicit 
assumption that the bulk coefficients .mu..sub.a ' and .mu..sub.s ' and 
the thickness x' will remain constant between the calibration and analysis 
samples. Cope et at. found that use of such a wavelength dependent DPF 
with these assumptions "significantly improved the residuals generated by 
multilinear regression analysis." At the same time, they also note that 
the calibration of the DPF is impractical in vivo, where .mu..sub.a ' can 
not be varied independently, and instead propose an additional measurement 
of the mean time of flight of detected photons as a measure of the DPF. 
Their calibration method is clearly highly complex even for this case in 
which the analyte itself contributes roughly 1/10 of the total bulk 
absorption coefficient .mu..sub.a '. This is 1,000-10,000 time higher than 
the equivalent relative contribution of glucose to the total absorption in 
the present invention. 
As noted in the U.S. patent application No. 08/333,758, the small size of 
the analyte absorption allows a convenient alternative calibration method 
in which the effect of the analyte is to create small linear perturbations 
of the signal intensities from a set of reference intensities which are 
defined by the absorption of major constituents, such as the hemoglobins, 
and the thickness and scattering characteristics of the sampled site. 
These reference intensities are themselves highly non-linear, comprising a 
mixture of the functional dependencies of equations 1! and 2!. The 
preferred methods described in the Block '265 Patent and the related 
application take advantage of the broad and shallow wavelength variations 
of the major absorbers and the scattering characteristics of the tissue to 
facilitate the accurate determination of the reference intensities, and do 
so dynamically on each individual measurement so as to adjust 
automatically for the inevitable physiological changes in the selected 
sites from day-to-day. The use of restricted solid angle, particularly in 
the form of collimation carries the functional behavior closer to that of 
equation 1! as the effects of scattering, absorption, and thickness 
become decoupled so that the estimation of the reference intensities 
become more reliable and more easily calibrated. This trend applies across 
the whole spectrum of calibration methodology disclosed in the related 
patent and applications, including the simultaneous intercomparison of 
measurements on multiple sites and/or with multiple detector 
configurations with different wavelength responses. 
Most importantly, neither the theoretical models outlined above nor the 
experimental measurement systems employed using classical methods deal 
well with the inhomogeneities in tissue and other layered scattering 
media. Closed-form solutions only appear for extremely limited assumptions 
of the mutual variation of .mu..sub.a ' and .mu..sub.s ' with depth into 
the scattering medium. The Monte Carlo approach can handle tissue 
inhomogeneities somewhat better, but also require input values for the 
coefficients and the angular distributions of single scattering events 
that are very difficult to obtain reliably. Even the early experimental 
work of Kubelka, J Opt Soc America (1954) 44,330-5, demonstrated that in 
diffuse light, the impact of an inhomogeneous layer is very strongly 
dependent on its depth when viewed in reflectance, and somewhat less 
strongly when viewed in transmission. His transmission results also showed 
a remarkable symmetry of the effect of a strongly absorbing layer in which 
it has the identical impact on the total transmission in two locations 
that are symmetrically offset from the midpoint of the medium. 
Little theoretical work has been done as yet on the impact of lateral 
inhomogeneities within individual layers in the scattering medium. These 
are on the one hand equivalent to classical wedging errors in 
spectrophotometry, and on the other hand potentially more complicated 
because of the non-linearities inherent in the admixture of diffuse light. 
At the same time, the available sites for noninvasive measurement of 
glucose and other trace constituents can not be expected to be homogeneous 
at the part-per-million level within a given individual, let alone across 
a population of individuals. 
One preferred embodiment of the present invention, which is designed to 
address the simultaneous need for high detected intensity and tolerance of 
inhomogeneity in the sampled tissue, is shown in FIG. 2. This optical 
system represents an improvement over those taught in the '265 Patent and 
related applications, in that additional optical elements have been added 
to decrease the angular range of light leaving the tissue which will be 
accepted into the detector elements. The attendant loss of intensity at 
the detectors is compensated by increased size of the detectors, 
beamsplitters, and the geometrical area of the incident and detected light 
beams. 
This optical system employs collimating optics for both illumination and 
detection, with the detector having a plurality of detector units placed 
such that they achieve congruent sampling. Radiation source 10 is selected 
so that it provides broad spectrum illumination, e.g., 700-2500 um 
illumination. Radiation from radiation source 10 passed through 
collimating lens 12 before striking tissue 20. Optional aperture 14 is 
shown which helps define the collimation optics in conjunction with 
collimating lens 12. Tissue sample 20 may be any sample but a thin, 
repeatable sample such as a finger web is preferred. For this type of thin 
sample, the amount of scattering is minimized and it is possible to hold 
the web and compress it such that a standardized thickness may be 
achieved. The more standardized the thickness is, the more likely 
universal calibration (or calibration to a limited group of thicknesses) 
can be. If universal calibration can be achieved, the calibration can set 
at the factory and corrections for effective pathlength can be made in the 
instrument itself. If such universal calibration is not achieved, some 
calibration measurements may be required before meaningful data can be 
obtained. 
Once the radiation has traversed tissue sample 20 and exits the tissue 
through area defining aperture 25, it passes through detector collimating 
optics 30 formed of converging lens 32, aperture 34 and recollimating lens 
36. This type of collimating optics is conventionally used in telescopes 
and other devices where collimation of light is desired. The collimated 
beam exiting collimation optics 30, specifically recollimating lens 36, 
then goes through a series of beam splitters 42A, 42B and 42C and onto 
four detector units 44A, 44B, 44C and 44D. The beam splitters and detector 
units are arranged such that the entire detection unit 40 provides 
congruent sampling of the beam. More particularly, the beam splitters and 
detector units are arranged such that the pathlength and angles from 
recollimating lens 36 to any of detector units 44 are equal and each of 
detector units 44 are optically superimposable upon the other. More 
details of this type of congruent sampling is set forth in U.S. patent 
application Ser. No. 08/130,257. Although four detector units are shown, 
the exact number may be varied. 
FIG. 2 also shows an additional testing or detector unit 50 (shown as a 
"black box") which is not in line with the collimation or angle 
restricting optics. In some circumstances, the scattered radiation may 
provide information in addition to or supplementing that obtained from the 
unscattered radiation. Additional detector unit 50, which may actually be 
a plurality of detector units, can be used at different positions and 
angles, thereby testing the scattered radiation and providing additional 
information. One particularly valuable embodiment uses this detector unit 
in reflectance or transflectance mode, that is on the same side of the 
tissue as the illuminating optics. This additional detection unit can be 
used with any of the embodiments shown herein. 
The optical system of FIG. 2 preserves spatial information, in that there 
is a one-to-one correspondence between a location on the viewed tissue 
surface area and a location on the active area of each detector. With 
adjustment of lens-aperture-lens-detector distances, a reasonably sharp 
focus can be achieved, particularly as the accepted viewing solid angle 
shrinks towards perfect collimation. As a result, replacement of the 
detectors shown in FIG. 2 by imaging devices such as CCD's or other array 
detectors creates an imaging system that produces "shadowgrams" using 
"ballistic" or "snakelike" photons, with the added advantage of 
simultaneous congruent imaging in all pixels in the detector arrays. The 
simultaneity facilitates the real-time processing of the signals to form 
tuned images of subsurface structures in the different analytes, including 
temperature. Since certain tumors are known to have a different 
temperature than surrounding tissue, this imaging system has uses for 
detection of rumors and other anomalies. The illumination source and 
detector array can be scanned across a large tissue sample and combined to 
form a larger shadowgram. 
The essence of the present invention relative to the prior art is 
illustrated in the sequence of FIGS. 3A-E. FIG. 3A is an expanded scale 
version of FIG. 2, showing the way in which the collecting lens and angle 
defining aperture function to limit the acceptance cones of the light 
emerging from the tissue at different point on the exit area. Light 
emerging from the tissue at angles outside the acceptance cone from 
anywhere in that exit area, defined by aperture 25, strike the angle 
defining aperture outside its central opening. As the relative magnitudes 
of the forward collimated light intensity and the diffuse scattered light 
intensity are not know exactly from either theory or experiment, as 
explained above, the aperture size and the resultant angular acceptance 
cone must be tailored to the particular observation sites selected. 
FIG. 3B illustrates the difference between the present invention (top) and 
prior art (bottom) systems using fiber optics to introduce incident light 
to the scattering media, and then transport the transmitted light to 
analysis means. The angular ranges over which incident and detected light 
are launched or received by the optics are indicated by the arrows. These 
angular range for fiber optics cover their full numerical aperture, which 
typically comprises an .about.50 degree full angle cone, while the present 
invention illuminates the tissue with nearly parallel light, and detects 
light emerging in only a small angle cone whose full angle is less than 20 
degrees, and preferably less than 10 degrees. Light entering along any of 
the rays shown in the figure will spread sideways about that ray due to 
scattering, and the amount of spread will vary with the wavelength of the 
light. Light which has spread laterally has a much higher probability of 
reaching the detector in the prior art arrangement shown because the 
receiving fiber optic will accept light over a much higher solid angle. 
The arrangement shown for the present invention configuration provides 
benefit even for the detection of diffuse radiation because the 
contributions of inhomogeneities near the edges of the observed areas will 
be more evenly sampled at all wavelengths. In addition, to the extent that 
the acceptance cones, tissue thickness, and tissue scattering parameters 
can be chosen to include predominantly unscattered collimated light 
reaching the detectors, the variation of that detected light with analyte 
concentration will trend from the inherently complex form in equation 2! 
to the inherently more readily calibratable form of equation 1!. 
FIG. 3C illustrates the optical geometry employed in most commercial pulse 
oximeters. Here, LEDs (50) at the selected wavelengths are placed adjacent 
to each other on one side of the tissue, and emit their radiation into the 
tissue at angles similar to those of the previously described fiber 
optics. The detector is placed on the far side of the monitored tissue, 
with an acceptance cone for light leaving the tissue that can approach the 
full hemisphere. Again, scattering spreads the light sideways about each 
incident ray, and light which the detectors receive will have sampled a 
region much wider than the spacing between the two detectors. This light 
includes contributions from tissue inhomogeneities which lie at the outer 
edges of the indicated rays, and since the degree of spreading is 
wavelength dependent, these inhomogeneities make different contributions 
to the signal at each wavelength. This effect is a major contributor to 
the high sensitivity of pulse oximetry results to motion of the site, 
compression of the site by the instrument, and to small changes in tissue 
thickness, all of which are greatly reduced in the present invention. To 
meet the present invention, detector 54 is broadened in area as shown by 
the dotted lines to be at least comparable in size to the sample 
thickness. 
FIG. 3D illustrate the way in which the optical system of the present 
invention can be employed to detect radiation reflected or backscattered 
from the monitored site. This arrangement is particularly advantageous for 
sites such as the forehead or eyelid, where it is impractical to place a 
detector on the far side of the tissue. Here, the incident and emerging 
light beams preferably cross within the tissue. Once again, the advantage 
of this optical arrangement, even for detection of diffuse as opposed to 
singly scattered radiation, is that the relative contribution of tissue 
inhomogeneities near the edge of the illuminated area are rendered more 
equal between detectors at different wavelengths. The particular angles 
shown should be taken as illustrative rather than restrictive; there are 
numerous alternative arrangements well known in the art which provide 
additional advantages. One worthy of mention is the case in which the 
incident radiation would enter normal to the surface, with the restricted 
angle acceptance cones lying along (all or) part of an annulus centered 
about the incident beam. 
FIG. 3E illustrates two optical configuration often employed in the prior 
art for the measurement of reflected light. The fiber optic (60/62) 
configuration shown suffers from the same limitations as that in FIG. 3B, 
with the wavelength dependent differences in scattering and hence 
penetration to deep inhomogeneities skewing the relative signal between 
detectors even more strongly. The 2nd configuration in FIG. 3E, with LEDs 
(70) and the detector (74) on the same side of the monitored site, has 
still greater potential for unequal sampling of both lateral and axial 
inhomogeneities because of the larger angular acceptance of the detector. 
Although FIG. 2 shows a collimated beam for illuminating the tissue, in 
some circumstances it may be better for the illumination beam to be not 
perfectly collimated FIG. 4 shows two variations which implement such a 
non-collimated beam condition; FIG. 4A shows a converging beam for tissue 
illumination while FIG. 4B shows a slightly diverging beam for tissue 
illumination. More particularly, the apparatus of FIG. 4A has a radiation 
source 110 which generates a radiation beam that passes through a 
converging lens 112 before striking tissue 120. The focal point of 
converging lens 112 is on the opposite side of tissue 120 from radiation 
source 110. The radiation transmitted through tissue 120 passes through 
converging lens 132, preferably an aperture 134, and a recollimating lens 
136. 
FIG. 4B shows substantially the same system as FIG. 4A except lens 212 is a 
slightly diverging lens as a compared with the slightly converging lens 
112 in FIG. 4A. Radiation source 210, lens 232, aperture 234 and lens 236 
are the substantial equivalent of their corresponding numbered parts (110, 
132, 134 and 136, respectively) in FIG. 4A. While FIGS. 4A and 4B 
illustrate an output beam from the optical system which is collimated, it 
may be that not all of optics 130 in FIG. 4A or optics 230 in FIG. 4B is 
necessary since a slight variation from collimation on the output beam may 
be desirable to get the best ratio of snakelike/ballistic to scattered 
rays. The configuration chosen is the one which empirically optimizes the 
transmitted photon intensity while maintaining insensitivity to internal 
inhomogeneities in the tissue. 
FIG. 5 shows another variation on the optical system of the invention, one 
with a more highly converging lens 312 which has a focal point on the far 
side of tissue 320. An aperture 334 is placed at or near the focal point 
of lens 312 and a recollimating lens 336 is placed near aperture 334. By 
placing aperture 334 at the focal point of lens 312, a pinhole camera-type 
system is created whereby the image of tissue 320 is reversed but formed 
directly upon the detector. Again, this optical system may have 
advantageous properties depending on the type of tissue or other sample 
measured. 
FIG. 6 shows a variation of FIG. 2, whereby instead of the beam splitter 
apparatus 40 shown in FIG. 2, a bifurcated optical bundle which is split 
into four parts, each leading to a different detector unit, is 
substituted. If the detector units are located such that the length of the 
optical fiber leading to the particular detector unit is identical, this 
system provides an approximation of the congruent sampling shown in FIG. 
2. More details concerning this type of bifurcated optical bundle is 
described in U.S. patent application Ser. No. 08/130,257. 
FIG. 7 shows a system using the beam splitter array of FIG. 2 reversed for 
congruent illumination rather than congruent sampling. Four radiation 
sources 710A, 710B, 710C and 710D, are used to illuminate the tissue 
sample. The radiation issuing from each of the radiation sources goes 
through a collimating lens (712A, 712B, 712C and 712D, respectively) and 
then is redirected by one of the beam splitters 716A, 716B or 716C to 
illuminate tissue 720. The radiation transmitted by tissue 720 passes 
through converging lens 732 and aperture 734 before striking detector 744. 
Optionally, an additional lens 736 (not shown) could be used to 
recollimate the transmitted radiation before it strikes detector 744. The 
radiation sources, collimating lenses and beam splitters are arranged to 
provide congruent illumination and each separate radiation source may have 
an associated modulator to provide a different modulation to the radiation 
issuing from that radiation source. This type of modulation apparatus, and 
its advantages, is described in more detail in U.S. patent application 
Ser. No. 08/182,572. Briefly, using a plurality of modulators each 
providing a different modulation to the associated radiation issuing 
therefrom, and using a form of modulation differentiation at the detector 
(such as electronically separating the signals based on modulation 
frequency) provides a method which allows differentiation at the detector 
of the source of the illuminating radiation, and accordingly allows 
additional information to be generated from a single detector. For 
example, if the radiation sources cover different wavelengths, a single 
detector can differentiate the intensity of the transmitted radiation at 
each wavelength range by using the modulation to determine the wavelength 
range. This can eliminate the requirement of the system illustrated above 
which requires a plurality of detector units. For improved results, both 
the congruent illumination shown in FIG. 7 and the congruent sampling 
shown in FIG. 2 may be used in the same device. 
FIG. 8 shows a different embodiment of the invention whereby a polarizer 
and an analyzer are used in conjunction with the optical system in the 
invention to provide signal differentiation. More particularly, radiation 
source 810 emits radiation (shown with the polarization distribution 812) 
which then passes through polarizer 815. Only polarized radiation (see 
816) illuminates tissue 820 and since the act of diffuse scattering in 
tissue 820 will depolarize the scattered light, only unscattered (or 
forward scattered) radiation is transmitted from tissue 820 as polarized 
radiation. The radiation is then transmitted through analyzer 850, which 
is a polarizer that passes only radiation which has the same polarization 
as polarizer 815, and detector 844 (not shown) detects this polarized 
radiation to yield a signal. 
One advantage of using this polarizer/analyzer system is that it 
immediately segregates scattered from unscattered radiation, since 
scattered radiation is depolarized and cannot pass through analyzer 850 to 
detector 844. Accordingly, although collimation optics and the restricted 
solid angle can be used with the apparatus of FIG. 8, these are not 
absolutely necessary because the polarizer/analyzer pair will eliminate 
the off-angle scattered radiation in any case. 
In all of the foregoing embodiments, there is an advantage to using 
geometrically broad beam radiation as opposed to narrow beam radiation. 
The use of such broad beam radiation provides a higher input signal which, 
when constrained by the solid angle restrictions of the present apparatus, 
still provides sufficient illumination and signal to meet the precision 
requirements for the analysis of trace constituents. Further, once the 
signals have been received, the processing described in the Block '265 
Patent and the previously cited applications may be applied to 
differentiate signal from background and obtain meaningful data. 
The advantages of the present invention apply to spectrophotometric systems 
such as those employed in pulse oximetry. While the shot-noise constraints 
on the detected intensity are lower because the absorption of the 
hemoglobins are so much larger the acceptance angle restrictions provide 
greater linearity and improved calibratability, as well as reduction in 
the severity of motion and breathing artifacts, and other limitations on 
universality of calibration. 
For the analysis of trace constituents where the high photon flux 
requirement is critical, the present invention is particularly 
advantageous when combined with the use of broadband and broadband 
overlapping detectors, as taught in the Block '265 Patent and the parent 
applications. 
The foregoing description is meant to be explanatory only and is not 
intended to be limiting as to the scope of the invention. The invention is 
defined by the following claims.