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Description  |
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FIELD OF INVENTION
The present invention concerns the photometric determination of glucose in
the bloodstream or tissues of patients suspected to have developed
diabetes. This determination is carried out by measuring the optical near
infrared absorption of glucose in regions of the spectrum where typical
glucose absorption bands exist and computing the measured values with
reference values obtained from regions of the spectrum where glucose has
no or little absorption and where the errors due to background absorptions
by the constituents of the surrounding tissues or blood containing the
glucose are of reduced significance or can be quantitatively compensated.
BACKGROUND OF THE ART
Many methods and devices have been developed up to now for the
determination of glucose in vitro or in vivo by optical means.
For instance, in PCT application WO No. 81/00622, there is disclosed an IR
absorption method and apparatus for determining glucose in body fluids.
According to this reference, absorption spectra of serum or urine, both
transmissive or reflective, i.e. due to back-scattering effects, are
measured at two distinct wavelengths .lambda.1 and .lambda.2, .lambda.2
being typical of the absorption of a substance to be measured (for
instance glucose) and .lambda.1 being a wavelength at which the absorption
is roughly independent of the concentration of the substance of interest.
Then the pertinent measured data are derived from calculating the ratio of
the absorption values at .lambda.1 and .lambda.2, the bands of interest
being in the range of 940-950 cm.sup.-1 (10.64-10.54 .mu.m) and 1090-1095
cm.sup.-1 (9.17-9.13 .mu.m), respectively. In this reference, the source
of irradiation is provided by a CO.sub.2 laser.
Swiss Pat. No. CH-612.271 discloses a non invasive method to determine
biological substances in samples or through the skin using an attenuated
total reflection (ATR) prism. This method relies on the passing of an
infrared beam through a wave-guide (ATR prism) directly placed against a
sample to be analyzed (for instance the lips or the tongue). The
refractive index of the wave-guide being larger than that of the medium of
the sample (optically thinner medium), the beam propagates therein
following a totally reflected path, the only interaction thereof with said
thinner medium (to be analyzed) being that of the "evanescent wave"
component at the reflection site (see also Hormone & Metabolic Res./suppl.
Ser. (1979), p. 30-35). When using predetermined infrared wavelengths
typical of glucose absorption, the beam in the ATR prism is attenuated
according to the glucose concentration in the optically thinner medium,
this attenuation being ascertained and processed into glucose
determination data. U.S. Pat. No. 3,958,560 discloses a non-invasive
device for determining glucose in patient's eyes. Such device comprises a
contact-lens shaped sensor device including an infrared source applied on
one side of the cornea and a detector on the other side thereof. Thus,
when a infrared radiation is applied to the area under measurement, light
is transmitted through the cornea and the aqueous humor to the detector.
The detected signal is transmitted to a remote receiver and a read-out
device providing data on the concentration of glucose in the patient's eye
as a function of the specific modifications undergone by the IR radiations
when passing through the eye.
GB patent application No. 2,033,575 discloses a detector device for
investigating substances in a patient's regions near to the bloodstream,
namely CO.sub.2, oxygen or glucose. The key features of such detector
comprise radiation directing means arranged to direct optical radiation
into the patient's body, and receiver means for detecting attenuated
optical radiations backscattered or reflected within the patient's body
i.e. from a region below the skin surface. The detected signal is
thereafter processed into useful analytical data. Optical radiations
include UV as well as IR radiations.
Other references rather refer to the measurement or monitoring of other
bioactive parameters and components such as blood flow, metabolic
oxyhemoglobin and desoxyhemoglobin but, in reason of their close analogies
with the aforementioned techniques, they are also worth reviewing here.
Thus U.S. Pat. No. 3,638,640 discloses a method and an apparatus for
measuring oxygen and other substances in blood and living tissues. The
apparatus comprises radiation sources and detectors disposed on a
patient's body, for instance about the ear to measure the intensity
passing therethrough or on the forehead to measure the radiation reflected
therefrom after passing through the blood and skin tissue. The radiations
used belong to the red and very near infrared region, for instance
wavelengths (.lambda.) of 660, 715 and 805 nm. The number of different
wavelengths used simultaneously in the method is equal to the total of at
least one measuring wavelength typical for each substance present in the
area under investigation (including the substance(s) to be determined)
plus one. By an appropriate electronic computation of the signals obtained
after detection from absorption at these diverse wavelengths useful
quantitative data on the concentrations of the substance to be measured
are obtained irrespective of possible changes in measurement conditions
such as displacement of the test appliance, changes in illumination
intensity and geometry, changes in the amount of blood perfusing the
tissue under investigation and the like.
GB patent application No. 2,075,668 describes a spectrophotometric
apparatus for measuring and monitoring in-vivo and non-invasively the
metabolism of body organs, e.g. changes in the oxido-reduction state of
hemoglobin and cellular cytochrome as well as blood flow rates in various
organs such as brain, heart, kidney and the like. The above objects are
accomplished by optical techniques involving wavelengths in the 700-1300
nm range which have been shown to effectively penetrate the body tissues
down to distances of several mm. Thus in FIG. 14 of this reference there
is disclosed a device involving reflectance type measurements and
comprising a light source for injecting light energy into a waveguide
(optical fiber bundle) applied to the body and disposed in such way
(slantwise relative to the skin) that the directionally emitted energy
whch penetrates into the body through the skin is reflected or back
scattered by the underlying tissue to be analyzed at some distance from
the source; the partially absorbed energy then reaches a first detector
placed also over the skin and somewhat distantly from the source. Another
detector placed coaxially with the source picks up a back radiated
reference signal, both the analytical and reference signals from the
detectors being fed to a computing circuit, the output of which provides
useful read-out data concerning the sought after analytical information.
Although the aforementioned techniques have a lot of merit some
difficulties inherent thereto still exist. These difficulties are mainly
related to the optical properties of the radiations used for making the
measurements. Thus, radiation penetration into the skin depends on the
action of absorbing chromophores and is wavelength-dependent, i.e. the
light in the infrared range above 2.5 .mu.m is strongly absorbed by water
and has very little penetration capability into living tissues containing
glucose and, despite the highly specific absorption of the latter in this
band, it is not readily usable to analyze body tissue volumes at depths
exceeding a few microns or tens of microns. If exceptionally powerful
sources (i.e. CO.sub.2 laser) are used, deeper penetration is obtained but
at the risk of burning the tissues under examination. Conversely, using
wavelengths below about 1 micron (1000 nm) has the drawback that, although
penetration in this region is fairly good, strong absorbing chromophores
still exist such as hemoglobin, bilirubin ad melanin and specific
absorptions due to glucose are extremely weak which provides insufficient
or borderline sensitivity and accuracy for practical use in the medical
field. In addition, the ATR method which tries to circumvent the adverse
consequences of the heat effect by using the total internal reflection
technique enables to investigate depths of tissues not exceeding about 10
.mu.m which is insufficient to obtain reliable glucose determination
information.
DISCLOSURE OF THE INVENTION
The present invention remedies these shortcomings. Indeed it was found
quite unexpectedly that by operating at some wavelengths located in the
1000 to 2500 nm range, acceptable combinations of sufficient penetration
depth to reach the tissues of interest and sufficient sensitivity in
ascertaining glucose concentration variations could be accomplished, this
being without risks of overheating tissues. At such penetration depths of,
say, 0.5 mm to several mm, representative information on the conditions of
patients could be gained in regard to possible lack or excess of glucose
in the blood stream (hypo- or hyperglycemia). Therefore, one object of the
invention is a spectrophotometric method for the transcutaneous
non-invasive determination of glucose in patients suffering or suspected
to suffer from diabetes in which a portion of said patient's body is
irradiated with the light of a directional optical lamp source, the
resulting energy I either transmitted or diffusely reflected
(back-scattered) by a sample volume of body tissue underneath the skin of
said body portion being collected and converted into suitable electrical
signals, said collected light including at least one spectral band of a
first kind containing a measuring signal wavelength .lambda.G typical of
the glucose absorption and at least another band of a second kind with a
reference signal wavelength .lambda.R typical of the background absorption
spectrum due to the tissue containing the glucose but in which the
absorption of the latter is nil or insignificant, and in which method said
electrical signals (the value of which, IG and IR, are representative of
the absorption in said measuring and reference bands, respectively) are
fed to an electronic computing circuit for being transformed into glucose
concentration readouts, characterized in that the bands of the first and
second kind below to the 1000 to 2500 nm medium near-IR range, .lambda.G
being selected from 1575, 1765, 2100 and 2270+ or -15 nm and .lambda.R
being selected either in the range 1100 to 1300 nm or in narrow regions
situated on both sides of the measuring bands but outside the area where
glucose absorbs strongly.
BRIEF DESCRIPTION OF THE DRAWING
FIG. 1 represents schematically the main components of an apparatus for
non-invasively measuring glucose in vivo by an absorptive transmission
technique.
FIG. 2 represents schematically a detail of a variant of the apparatus of
FIG. 1 designed to operate by an absorptive reflective technique.
FIG. 3 represents schematically the components for processing the
electrical signals obtained from the light gathered after being partly
absorbed in the region of interest and for computing and converting said
signals into useful readouts of glucose determination.
FIG. 4 represents a plot of absorption measurement data versus glucose
concentration at both .lambda.G=2100 nm and .lambda.R=1100 nm.
FIG. 5 represents an infrared spectrum of glucose (1 mole/1 aqueous
solution) from which the corresponding infrared spectrum of water has been
subtracted.
FIG. 6 is like FIG. 5 but refers to blood serum and water.
FIG. 7 is like FIG. 5 but refers to human serum albumin.
FIG. 8 is like FIG. 5 but refers to keratin.
FIG. 9 is like FIG. 5 but refers to collagen.
FIG. 10 is like FIG. 5 but refers to HCO.sub.3.sup.-.
The light absorbed by the tissue subjected to analysis constitutes together
with other losses due to scattered stray radiations inherent to the
practice of the method and the apparatus components, the background
response noise from which the useful signals must be separated. The
absorbing entities in the body media containing the glucose include
peptidic substances such as albumin, keratin, collagen and the like as
well as low molecular weight species such as water, hydrogenocarbonate,
salt and the like. These substances all have characteristic absorptions
distinct from the aforementioned selected typical glucose absorptions as
shown by the infrared spectra of FIGS. 5 to 10; and compensation can thus
be afforded by subjecting the collected measuring and reference data to
computation according to programmed algorithms. Further, the time
concentration variation of the components depicted in FIGS. 5 to 10 in the
blood and/or living tissues follows a pattern different from that of
glucose in the measurement location, which difference is also usable to
determine glucose in the presence of such components. Examples of possible
computation algorithms are provided in the following reference: R. D.
ROSENTHAL, an Introduction to Near Infrared Quantitative Analysis, 1977,
Annual Meeting of American Association of Cereal Chemists.
According to one general method of computing applicable in the present
invention a normalizing factor is first established from the differences
in absorptions in the reference band when glucose is present or absent or
in insignificant quantities in the tissue to be analyzed. Then this factor
is used to normalize the measured value of glucose absorption in the
.lambda.G band, the reference value being subtracted from the normalized
value to provide the data for expressing the correct glucose concentration
in the sample. The normalizing factor can be determined for instance by
setting the reference's wavelength at an isosbestic point (i.e. a
wavelength at which there is no significant change in absorption although
the concentration of glucose may change).
Another way to obtain a normalizing factor is to focus alternately from the
place where glucose should be analyzed to a place where the amount of
glucose is either insignificant or constant and fairly well known, the
background absorption spectrum being substantially constant or comparably
shaped in the two locations. One will see hereinafter how this can be
practically implemented.
According to another way of computing the absorption measured values into
useful glucose determination results is to differentiate the IG and IR
signals with regard to .lambda. within the area of the bands of the first
and of the second kind, respectively; and then to subtract one
differential from the other and obtain the desired result from the
difference. Reference to this method is provided in T. C. O'HAVER
Potential Clinical Applications of Derivative and Wavelength Modulation
Spectrometry, Clin. Chem. 25(a), 1548-53 (1959).
The invention also concerns an apparatus for carrying out the present
analytical method.
This apparatus comprises a light source for directively applying a beam of
light on a portion of a patient's body, the spectral composition of said
light being such that it can penetrate through the skin to a region where
glucose concentration can be measured with significance regarding the
patient's conditions and from which said light can be gathered after being
partially absorbed as a function of the glucose concentration, a
collecting means for gathering the radiation transmitted or reflected
(transflected) from said region, detector means for detecting and
converting into electrical signals the gathered light as distinct
wavelengths belonging to at least two bands, one measuring band and one
reference band, and computing means to transform said electrical signals
from useful readouts representative of the desired glucose measurement
data. One characteristic feature of an embodiment of this apparatus is
that it comprises means for varying continuously or stepwise the incidence
angle relative to the body of said beam of light, said variation resulting
in a consequent variation of the depth of the center of said region
wherefrom the light is gathered after absorption.
Such an apparatus and variants thereof will now be described with the help
of the annexed drawing.
The apparatus represented in FIG. 1 consists of two main sections, a light
source section 1 and a detector section 2. The light source section 1
comprises a light source 3, for instance a halogen lamp and light
directing means, for instance a reflector 4 and a condensor 5 for
providing a directed beam 6 of light. This beam needs not be polarized or
coherent but, of course, such types of light can also be used if desired.
When using a wide band continuous spectrum of light, the apparatus also
comprises a filter or system of filters 7 to block out undesired
wavelength ranges mainly caused by higher order diffraction at the
monochromator grating; in this particular application where the signals
should be in the range of about 1000 to 2700 nm, visible ranges are
eliminated by using a SCHOTT RG780 filter (0.8-1.3 .mu.m), a silicon
filter (1.3-2.2 .mu.m) and a Ge filter (2.2-4.2 .mu.m). The lamp is a 12
V, 100 W halogen lamp with the following properties: color temperature
3300.degree. K.; maximum output at 850 nm; average luminance 3500
cd/cm.sup.2. Of course, if monochromatic light sources were used in the
present apparatus, for instance by means of tunable lasers, the blocking
filters 7 would no longer be necessary.
The apparatus further comprises a monochromator 8 with inlet slit 8a and
output slit 8b and a grating 8c. The monochromator can be of any suitable
origin, for instance a JARREL-ASH model 70000 with sine bar wavelength
drive is suitable. The monochromator can scan or repeatitively shift from
one selected wavelength to another one or, in succession, to several ones
depending on whether one or more measuring and references wavelengths are
used concurrently in the analysis. The shifting or scanning rate of the
monochromator is programmed and controlled by the computer circuits to be
described later and the signals thereto are provided by line 12a. Of
course, if the source light is provided by means of lasers of specific
wavelengths, the monochromator is no longer necessary.
The selected monochromatic beam 9 which emerges from the monochromator
passes through a chopper disk 10 driven by a motor 11 whose rotation is
controlled by a clock (not represented but conventionally inherent to any
chopper system); this system also provides timing signals, the output of
which is schematically represented by the arrow 12b, to be used for
synchronizing the analog and digital electronic processing circuits as
will be seen hereinafter. The periodical interruption of the excitation
beam of light 9 by the chopper disk is required for removing or minimizing
the background noise due to ambient light, detector dark noise, and other
stray signals, i.e. the detector will alternately signal the background
alone or the total of signal plus background from which the latter can be
evaluated and compensated by sutracting the difference. As an example the
chopper can operate with a 30 slot at a frequency of 500 Hz.
The detector section 2 of the apparatus is shown applied against an organ
of the body to be investigated, for instance the ear lobe 13 in a manner
such that the composite monochromatic beam 9 passes through that organ
before reaching the detector section whereby it is attenuated by partial
absorption and partial diffusion in the tissues under examination. As we
have seen before, the main components of the body tissues competing with
glucose as light absorbers in the spectral region of interest are the
water and the proteins of the cells and interstitial fluid; however, the
general distribution of these "background" constituents is fairly constant
and so the general "shape" of the corresponding spectrum superimposed to
that of glucose is also rather constant including the bands with points
the intensity of which is substantially independent of the glucose
concentration (isosbestic points). Therefore, as already mentioned,
correlating the absorption of the background at the isosbestic points
(wavelengths of reference) with the effective thickness of the tissue
layer of the organ under investigation traversed by the incident beam
enables to determine the reference absorption factor used for normalizing
the absorption data made at the typical glucose .lambda.G wavelengths
disclosed heretofore wherefrom the ultimate glucose concentration results
are obtained.
In this connection, it should be noted that the principle of the
aforementioned analysis can be expanded to analyze a three conponent
mixture containing glucose, serum and water. Indeed, serum contains
essentially all the dissolved constituents in blood or body fluid and, as
mentioned above, several features in the absorption spectrum of serum are
quite different from that of glucose. These features depicted from the
curves of FIGS. 5 and 6 are emphasized in Table 1 below. Therefore the
concentration of glucose can be estimated from absorbance measurements
using at least three different wavelength.
TABLE 1
______________________________________
Wavelength (nm)
Spectrum of serum
Spectrum of glucose
______________________________________
1 574 flat peak
1 732 peak slope
1 765 dip peak
2 052 peak slope
2 098 dip peak
2 168 peak slope
2 269 slope peak
2 291 peak dip
2 292 -- peak
______________________________________
By comparing FIG. 5 and FIG. 6, it is seen that very similar features still
exist in the two spectra. These features are the followings:
1 100 to 1 350 nm, flat portion and 2 223 nm, dip portion.
The detector section 2 comprises a light collecting integrating sphere or
half-sphere (sometimes referred to as Ulbricht sphere) the wall of which
is layered with a dull high reflective coating for gathering and
collecting all available light penetrating through an opening 13a of the
device which is directly applied against the organ under investigation
(the ear-lobe in this embodiment). Materials which are highly reflective
in the 1 to 2.7 .mu.m range are for instance Eastman 6080 white
reflectance coating containing barium sulfate or gold plating, the latter
having a better reflectance at the long wavelengths of the range. Using an
integrating full sphere is generally preferred unless a half-sphere is
necessary because of geometry considerations (see, for instance, the
modification of FIG. 2). When this is required because of the positioning
of the device about the ear, the integrating sphere is halved and its flat
portion consists of a highly reflective mirror (gold coating). The
performance of the half-sphere of this construction is somewhat less than
that of the full sphere but still acceptable because the mirror optically
mimics a full sphere. Differently stated, a full sphere is somewhat more
efficient for collecting light but more bulky, so a compromise between
sufficiently reduced physical size and sufficient efficiency is actually
made in this embodiment. In the present drawing the curved portion of the
half-sphere is presented as having ends somewhat flattened; however this
should not be considered physically significant; the reason thereto being
only of drafting convenience. The light collected by multiple reflection
in the half-sphere escapes through opening 13b and is condensed by means
of a condensor 14 to fall on a detector 15. Any detector sensitive to the
range of wavelengths used here can be used; an example of such detector is
a low temperature operating indium arsenide photodiode (JUDSON INFRARED
INC. Pa 18936 USA) having the following properties.
______________________________________
Model J 12-D
Peak Wavelength 2.8 .mu.m
Operating Temperature
77.degree. K.
Time constant 0.5-2 .mu.sec
Size 2 mm (diameter)
Responsivity 1 A/W
D 4.10.sup.11 cmHz.sup.1/2 W.sup.-1
Package Metal Dewar
Liquid N.sub.2 Hold Time
6-8 hr
Field of view 60.degree.
______________________________________
Light collector means different from the integrating sphere can also be
used in the present invention. Such means consist of an arrangement of
curved surface mirrors internal reflective ellipsoid or paraboloid surface
portions) which collect the light exiting from the body tissues and focus
it onto the detecting means. Such light collecting means may have improved
collecting efficiency over that of the integrating sphere because of
reduced number of reflections. Examples of collecting arrangements
suitable for application in the invention can be found in the following
references: N. W. WALLACE, the Optical Layout of off-axis paraboles:
Photonics Spectra, September 1984, p. 55; HARRICK SCIENTIFIC CORP.,
Catalog HSC-83 (IR-Vis-UV accessories), Ossining, N.Y. 10562, USA.
The operation of the present apparatus is obvious from the previous
description; for measuring the glucose in the ear tissue of a sitting
patient, the detector 2 is affixed onto the inside portion of the ear
lobe, for instance maintained by appropriate straps, and the monochromatic
light from the source section 1 (usually mounted on an appropriate stand
or rig on the side of the patient's chair) is directed on the external
side of the ear portion directly facing the detector. The light beam 9
strikes the ear portion and after traversing it penetrates into the
collecting half-sphere 13 wherefrom it goes to detector 15 whereby it is
converted into an electrical signal. The beam 9 is interrupted regularly
by the action of the chopper for the reasons explained before and, when in
its non interrupted position, it provides an alternating dual or multiple
wavelength incident light input generated in the monochromator said
incident light comprising at least one measuring signal generally centered
about the aforementioned values of 1575, 1765, 2100 or 2700 nm and at
least one reference signal in the wide reference range or at the narrow
wavelength ranges on both sides of the .lambda.G wavelengths. Thus the
electrical signal obtained from detector 15 is a multiplex signal
repetitively carrying the information relative to the optical apparatus
background, the spectral background of the volume of matter being analyzed
and the glucose absorption measurements according to a schedule under
control of the chopper system 11 (line 12b) and the computer circuits
(line 12a). We shall see hereinafter how this multiplex signal is decoded
and processed.
Before doing so we shall turn to the modification of FIG. 2. This
modification of which only a portion is represented in the drawing
constitutes an integrated light source-detector device to be placed
directly over the skin, i.e. a device that operates according to the
principle of reflection or back-scattering of light by tissues under the
skin. The reference and measuring light signal generator of this device is
similar to that used in the case of the device of FIG. 1 up to the chopper
disk; therefore the light emerging from said chopper is given the same
numeral 9 in FIG. 2.
The integrated light source-detector device of FIG. 2 is represented as
being applied on the skin 20 of a patient; said skin being arbitrarily
represented by successive layers 20a, 20b and some underneath tissue 20c.
The present device comprises a movable mirror 21 which can be displaced
horizontally continuously or stepwise while maintained in reflecting
relationship with beam 9 so that the reflected beam 22 is permanently
directed into a horizontal slit 23 of the device. In order to more clearly
illustrate this point, a ghost image 21g of the mirror 21 after being
moved in a second position is provided on the drawing. The light beam 22
reflected by mirror 21 meets the skin at an incidence angle indicated by
.lambda.. When the mirror is displaced in position 21g, its orientation is
such that the reflected beam 22g meets the skin at an angle .beta. smaller
than .lambda.. The mechanical means to move and synchronously tilt the
mirror 21 are conventional and not represented here. The present device
further comprises as in the previous embodiment a collecting half-sphere
24 with an input opening 24a, a condensor 25 and a light detector 26 for
converting the light gathered into an electric signal represented by an
arrow in the drawing.
The operation of the present device, which is fairly obvious from its
description, enables to undertake modulated depth glucose analysis below
the skin. Indeed, during analysis, the mirror 21 can be moved back and
forth so that the angle of penetration (.alpha., .beta.) of the beam 22
can be changed at will. The angles of the corresponding penetrating beams
27 and 27g will change accordingly and so will the position of the
underneath region under illumination wherefrom the back-scattered energy
will be picked-up by the halfsphere entrance aperture 24a. This is clearly
seen from the drawing in which the back-scattered light is indicated by
numeral 28 when the excitation beam 22 falls at the angle .alpha. and by
numeral 28g when the excitation beam 22g reaches the skin at an angle
.beta.. This technique permits the alternating exploration of different
zones at different depths under the skin whereby different concentrations
of glucose can be determined or monitored for a period of time. This is
particularly useful for ascertaining the general shape of the background
spectrum, i.e. the absorption of the medium in absence of glucose or when
the concentration of glucose is insignificant or of low variability as is
the case in the superficial layer of the epidermis. i.e. the
stratum-corneum. Thus, the measurement of the absorption spectrum in the
region 20a immediately under the skin surface will provide reference
results which may be continuously or periodically compared to
corresponding results obtained from deeper layers of the epidermis or the
dermis, whereby useful data about the concentration of glucose in said
deeper layers 20c can be obtained, this being directly proportional to the
blood glucose concentration. The construction of the present embodiment
also enables to block the light directly reflected by the skin surface at
the impingement point. Indeed, such surface reflected component is
parasitic since it comprises no glucose information and only contributes
detrimentally to the background noise as in the embodiments of the prior
art. Another advantage of the present invention's embodiment is that it
obviates or minimizes possible disturbances caused by foreign substances
contaminating the skin of the region under examination.
The electric signals provided from detectors 15 or 26 are analyzed and
processed in circuits which constitute also part of the apparatus of the
invention.
Such circuits shown on FIG. 3 comprise, starting from the photodetector 15
or 26 (depending on whether the embodiments of FIG. 1 or 2 are
considered), a preamplifier 30 (for instance a JUDSON INFRARED Model 700
having an amplification of 10.sup.7 V/A) a gain programmable amplifier 31
(for instance with gain varying from 1 to 200), an integrator 32 for
holding and averaging over noise and an analog to digital converter 33
(for instance a 16 bit unit). The integrator 32 is under control of a
timing unit 34 timed by the clock of the chopper 11 (see FIG. 1).
The digital signal issuing from converter 33 comprising, in succession and
according to a timing governed by said clock, the digitalized information
relative to the background noise, the glucose measurement signal and the
reference signals, is fed to a microprocessor 35 (for instance an APPLE II
microcomputer) also controlled by said clock whereby the information is
digested, computed according to a program of choice using one of the
calculating methods disclosed heretofore and displayed or stored in terms
of glucose determination data on either one or more of a monitor 36, a
printer 37 or a floppy disk recorder 38. The microprocessor 35 also
provides the signal for timing and controlling the wavelength scan or
selection of the monochromator 8 (see line 12a).
REDUCTION TO PRACTICE
The following discloses a practical test effected according to transmissive
technique (see FIG. 1). The data however apply equally well to the
reflective technique illustrated by the device of FIG. 2. The measurements
were carried out against an aqueous reference background such environment
being sufficiently close to that overall body tissues to be fully
significant. General physical considerations over absorption phenomena are
also provided for reference.
The fundamental relation between optical absorbance and the concentration
of the absorbing material is given by the Beer-Lambert law.
D=log.sub.10 (I.sub.o /I)=.epsilon..multidot.C.multidot.L
where
D=optical density, absorbance.
I.sub.o =intensity of incident light at wavelength .lambda..
I=intensity of light after passing through absorption cell.
C=concentration of the absorbing material (molar).
L=length of absorption path.
.epsilon.(.lambda.)=extinction coefficient.
The validity of this relation is generally satisfactory if the radiation is
monochromatic, if the concentrations of absorbing material are low, and if
there are no significant molecular interactions, e.g. association,
dissociation or structural changes for different concentrations. If the
measurement phenomenon involves some significant degree of scattering, the
above relation is no longer strictly valid and correction factors must be
introduced to restore its usefulness. Reference to such modification can
be found in GUSTAV KORTUM'S book Reflexionsspektroskopie, SPRINGER Verlag
(1969).
In the case of a mixture of m components, the Beer-Lambert law can be
generalized and expanded to include absorbance of each of the components
at each analytical wavelength.
##EQU1##
.epsilon..sub.i (.lambda.)=specific absorbance of a component i which is
wavelength dependent
C.sub.i =is the concentration defined as a mole fraction of the component
i, so that
##EQU2##
I.sub.o, I and L are defined as before.
In the experiments reported below, an apparatus such as that described with
reference to FIGS. 1 and 3 was used, the ear portion being replaced by
glucose solutions in water (pure water was used as reference). The
parameters were: sample concentration of glucose=C2; concentration of
water=C1 (C1+C2=1); path length for both pure water and solution=L;
extinction coefficient of water=.epsilon.1; extinction coefficient of
glucose=.epsilon.2.
The absorbances can be written in the two cases as:
##EQU3##
The absorbance difference
(.DELTA.D=log (I.sub.H.sbsb.2.sub.O)-log (I.sub.glucose-solution)
can then be written as
.DELTA.D=D.sub.2 -D.sub.1
or
.DELTA.D=LC.sub.2 (.epsilon..sub.2 -.epsilon..sub.1)
and
##EQU4##
This equation shows that the concentration of glucose C.sub.2 is
proportional to the absorbance difference .DELTA.D in the two samples
since the constant factor L(.epsilon..sub.2 -.epsilon..sub.1) is known
from operating conditions and kept constant.
As light of the incident intensity I.sub.o passes alternately through
samples 1 and 2 causing intensities I.sub.1 and I.sub.2, this can be
written:
C.sub.2 .about.D=log (I.sub.o /I.sub.1)-log (I.sub.o /I.sub.2)=log (I.sub.2
/I.sub.1)=log I.sub.2 -log I.sub.1 (1)
This means that it is sufficient to measure the difference of the
absorbance in the samples 1 and 2. The incident intensity I.sub.o need not
be measured.
Thus, the following three detected signals are processed in the
microprocessor 35. (The proportionality constant between light intensity
and detector signals is g).
S.sub.B =gB: background when there is no light falling onto the samples. B
is the background equivalent light intensity from ambient light plus the
detector noise.
S.sub.1 =g(I.sub.1 +B): signal caused by test sample 1=(intensity I.sub.1
plus background).
S.sub.2 =g(I.sub.2 +B): signal caused by reference sample 2=(intensity
I.sub.2 plus background).
For each sample, the difference between signal and background is taken,
resulting in .DELTA.S.sub.1 and .DELTA.S.sub.2.
sample 1: .DELTA.S.sub.1 =g(I.sub.1 +B)-gB=gI.sub.1
sample 2: .DELTA.S.sub.2 =g(I.sub.2 +B)-gB=gI.sub.2
These operations were synchronized by the chopper system (500 Hz) to
eliminate drifts of the background (zero) signal B which normally occur at
very slow rate.
The quantities .DELTA.S.sub.1 and .DELTA.S.sub.2 were measured
automatically for a number of times (100).
To find out the glucose concentration in sample 2, the absorbance value
from water in sample 1 was used as a reference and the two values were
subtracted from each other (see equation (1)):
.DELTA.D=log I.sub.2 -log I.sub.1 =log (.DELTA.S.sub.2 /g)-log
(.DELTA.S.sub.1 /g)
.DELTA.D=log .DELTA.S.sub.2 -log .DELTA.S.sub.1 (2)
The equation shows that the actual light intensities in equation 1 can be
replaced by the electrical detector signals. The result is not dependent
on the proportionality constant g.
The data processing program used in this embodiment is also able to compute
the errors of a set of measurements by using classical algebraic equation
of error propagation theory.
Four different glucose concentrations 0M, 0.05M, 0.5M and 1M and two
different wavelengths 1100 nm (reference wavelength .lambda.R) and 2098 nm
(test wavelength .lambda.G) were chosen.
At 1100 nm, the glucose spectrum is flat. The water absorbance has its
lowest value. At about 2100 nm (more precisely 2098 nm), the glucose
spectrum exhibits a characteristic absorption peak. The water absorbance
here is about two absorbance units higher than an 1100 nm.
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